British Journal of Radiology

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BJR The British Journal of Radiology 2006, Volume 79

The British Journal of Radiology January 2006, Volume 79, Issue 937 February 2006, Volume 79, Issue 938 March 2006, Volume 79, Issue 939 April 2006, Volume 79, Issue 940 May 2006, Volume 79, Issue 941 June 2006, Volume 79, Issue 942 July 2006, Volume 79, Issue 943 August 2006, Volume 79, Issue 944 September 2006, Volume 79, 945 October 2006, Volume 79, 946 November 2006, Volume 79, 947 September 2006, Volume 79, Special Issue 1 Volume 79 (2006), Case reports

BJR The British Journal of Radiology

January 2006 Volume 79 Issue 937

January 2006, Volume 79, Issue 937

The President’s Conference 2005: ‘‘Technology in Imaging and Radiotherapy – towards improved workflow and productivity’’ ● CT scanning the early days ● Cardiac applications of multislice computed tomography ● Technology solutions for better outcomes: integrated information management in key to productivity increases in medicine ● The case for particle therapy ● The contribution of PET/CT to improved patient management ● Mesenteric panniculitis in oncologic patients: PET-CT findings ● Diagnostic efficacy of SonoVueH, a second generation contrast agent, in the assessment of extracranial carotid or peripheral arteries using colour and spectral Doppler ultrasound: a multicentre study ● Lymphoepithelioma-like carcinoma of salivary glands: treatment results and failure patterns ● Comparison of patient doses in 256-slice CT and 16-slice CT scanners ● Assessment of tube current modulation in pelvic CT ● Radiosurgical palliation of aggressive murine SCCVII squamous cell carcinomas using synchrotron-generated X-ray microbeams ● Solitary pulmonary nodule with growth and contrast enhancement at CT: inflammatory pseudotumour as an unusual benign cause ● Non-haemorrhagic subdural collection complicating rupture of a middle cranial fossa arachnoid cyst ● Correspondence ● A deformed skull with enlarging hand and feet in a young female ● Acknowledgment to Referees ●

The British Journal of Radiology, 79 (2006), 1–4 DOI: 10.1259/bjr/19232533

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2006 The British Institute of Radiology

Commentary

The President’s Conference 2005: ‘‘Technology in Imaging and Radiotherapy – towards improved workflow and productivity’’ P P DENDY

Gu¨nter Dombrowe, the President of the British Institute of Radiology (BIR), introduced the theme of this year’s Conference, and explained its dual purpose – to highlight the contributions of medical and information technologies towards improving clinical practice, patient outcome and health economics; and to pay tribute to the pioneering work of Sir Godfrey Hounsfield, the inventor of CT scanning, perhaps the key technology of the digital imaging age. This Commentary provides an overview of some of the important topics discussed at the Conference. Some of the key presentations are also included in this issue. Elizabeth Beckmann reminded the audience of the early days of CT – the excitement generated by the images of the brain shown at the 32nd Congress of the BIR on 20 April 1972, the delightfully understated title of Sir Godfrey’s lecture – ‘‘Computerised axial tomography, a new means of demonstrating some of the soft tissue structures of the brain without the use of contrast media’’, and the subsequent publications in the BJR [1, 2]. The enduring memory of this and other early developments is that so much was achieved with so little money. Was Sir Godfrey one of the last brilliant, intuitive, string and sealing wax physics brigade? The first of the two nominated Hounsfield lecturers, Willi Kalender gave a comprehensive review of the past, present and future of CT from a physics and technology standpoint. He pointed out that there had been three distinct phases of development: (1) the 1970s had been a time of rapid development with second, third and fourth generation scanners; (2) the 1980s had been a period of stagnation with the competing development of MRI (the late 1980s was the only time during a 30 year period when there was no increase in the number of CT scanners in Germany); (3) the 1990s were the renaissance years, particularly with the introduction of spiral CT and multidetector arrays. Scan times are now typically 0.3 s to 0.5 s per full 360 ˚ scan and 10–30 s for the whole body. The first figure is important for temporal resolution, especially in cardiac applications, and one of the limitations on faster times is the centrifugal force to which sensitive components such as the X-ray tube are subjected [3]. To achieve better temporal resolution increased electronic control of the beam and possibly multiple tube designs are being explored. The British Journal of Radiology, January 2006

Improvements in total scan time will be achieved through further development of wider detector arrays, possibly towards flat panel detectors. This will in turn require X-ray tubes with an even higher peak output, as the total flux of photons required to image a given volume remains roughly the same. Like for like, patient doses have been reduced with tube current modulation both on rotation from anteroposterior (AP) to lateral projections and as the beam traverses the body from high to low attenuating regions. Achieving the same counting statistics on all data is a worthwhile goal [4, 5]. Since 1990 the emphasis has been on scanning volumes rather than slices and one of the landmarks has been to achieve isotropically uniform spatial resolution, typically in the range 0.4–0.6 mm [6]. It is important to recall that for isotropic resolution, radiation dose to the patient increases with the fourth power of the resolution element. These improvements must also be seen in the context of global use of radiology. CT is a relatively high dose technique, now accounting for 25% of all radiation exposure, and there must be strong clinical justification for its use, and in particular serial, repeat whole body scans. The future for CT is hidden from view but there are many possibilities and it is worthwhile to summarize Kalender’s predictions – more detector rows; shorter effective scan times; higher resolutions and more tissue parameters (there is renewed interest in superimposing, e.g. a calcium density map on a real density map obtained by dual energy CT [7]); lower doses (of course!). The second nominated Hounsfield lecturer, Adrian Dixon, reviewed the clinical advances in CT. Two important issues in particular were addressed:

(1) Do the ‘‘advances’’ in CT technology make any difference to the patient? (2) Many cutting-edge CT investigations are still charitably funded and if the NHS is to become responsible for their provision, they must be shown to be cost-effective. As a specific example of the clinical issues, he considered the impact of multidetector CT on abdominal problems. The improved anatomical resolution of modern helical CT scanners enables the diagnosis of acute appendicitis or the cause of small bowel obstruction to be made with a high degree of accuracy [8]. Consequent on its multitasking 1

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abilities, CT is increasingly being used as a means of triaging patients and facilitating early discharge for those without serious disease – with obvious benefits to the patients and cost savings to the NHS [9]. CT has become so good that in many areas of radiology the real questions are now (a) is there a role for plain film radiography? (b) when should ultrasound be used? (c) is there a role for MR other than to avoid the use of ionizing radiation? This success has come at a price: clinicians are tending to request a CT scan without fully examining the patient; surgeons are reluctant to operate without high quality imaging; for outpatients in oncology the number of requests for CT staging is starting to approximate the number of visits to hospital. However, Dixon was able to conclude on a positive note. For the patient CT has replaced some very unpleasant investigations. The tribute to Hounsfield concluded with a more specialized lecture from Albert de Roos on cardiac CT. Roos summarized the technical considerations for multislice CT in cardiac scanning – low contrast detection, spatial resolution at high contrast, temporal resolution, scan time and patient dose. The choice of acquisition variables and reconstruction characteristics is very dependent on the clinical problem under investigation. De Roos then reviewed a wide range of applications including: the quantitative assessment of coronary artery calcification [10, 11]; the assessment of coronary artery morphology; stent and graft patency; the selection of patients for invasive therapy; assessment of the anatomy of pulmonary veins and the investigation of acute chest pain. In the last of these applications there is now a one-stage protocol, i.e. the nirvana of the ‘‘one stop shop’’ to diagnose accurately both cardiac and non-cardiac causes of chest pain [12]. The Mackenzie Davidson lecture, delivered by Nicola Strickland, touched on many aspects of modern imaging but concentrated on information technology, especially PACS. PACS has now become a mature technology, especially as a result of the DICOM standard and network protocols. It clearly has the potential to improve workflow and productivity but does not, in itself, solve departmental inefficiencies and may highlight them. It is not a ‘‘quick fix’’ and must be an integral part of workflow engineering. Looking to the future, speech recognition and web browsers will be developed further. The electronic patient record remains a major challenge, since the facilities provided need to match the service being provided. A good example is home reporting – a full work load requires a full diagnostic service, emergency reporting needs only more limited facilities. Strickland concluded that technology provides the means for improving workflow and productivity – the challenge is to optimize the use of technology to maximize productivity in a clinically efficient way. Manufacturers’ views of the use and development of technology were also presented. Hermanns Requardt from Siemens Medical Solutions reminded us that, worldwide, challenges to healthcare systems are dominated by two main topics – demographic factors and progress in medicine. In diagnostic radiology, as in some other branches of medicine, for example molecular/genetic medicine, the challenge now is not a lack of information 2

but a flood of information. Drawing an analogy from industry where knowledge management systems are commonplace, Requardt predicted that information technology would bring about a paradigm shift in medicine if it could facilitate the formation of a clinical knowledge database and enable this to be used to complement the data from the individual patient. Jacques Souquet from Philips Medical Systems considered some other aspects of the impact of future technology on medical imaging. Picking up a theme from the previous speaker on progress in medicine, he pointed out that knowledge doubling times have fallen from about 8 years in 1970 to 1 year in 2001. Increased use of computer-aided decisions is one way to improve management of data, for example nodule identification in a radiograph, using embedded medical knowledge to reduce avoidable medical errors, genetic algorithms to discover diagnostic patterns in huge data sets. Souquet reminded us that much remains to be done. There are still several diseases for which no diagnostic test is available and the development of drugs to correct specific genetic flaws that are biological causes of cancer has a long way to go. In conclusion, he threw out two challenges:

(1) How can the translation from cell to mouse to man be speeded up? (2) How can the multidisciplinary constituencies contributing to progress (basic sciences, engineering, medicine, industry) be coordinated? This is a challenge that is close to one of the fundamental aims of the BIR. Jane Guinn from Kodak Ltd concluded the session by comparing the techniques of computed radiography (CR) and digital radiography (DR) from the standpoint of workflow patterns. She listed 16 distinct stages in the production of a traditional analogue film, many involving radiographer movement. CR removed only two steps, DR removed nine. This had a big impact on average examination time and in a busy general radiography room, on patient waiting time. Unfortunately DR does not provide the flexibility of CR for several examinations. Peter Williams delivered the Silvanus Thomson Memorial Lecture. With the somewhat enigmatic title ‘‘Things can only get better’’ he reviewed the development of external beam radiotherapy treatment delivery, concentrating on current developments and future promises. Early examples of ‘‘things getting better’’ included megavoltage therapy with Co-60; isocentric mounting; electrons as well as X-rays; anatomical data from the CT scanner for treatment planning. For a few years the ability to model tumours exceeded the ability to treat, which was restricted to a cylinder. In 1987 the multileaf collimator (MLC) became available for beam shaping and as with most really worthwhile medical developments, there were no formal health quality assessments or clinical trials. MLCs led to intensity-modulated radiotherapy (IMRT), essentially conformal therapy for difficult targets [13, 14], and at the same time electronic portal imaging was being developed to provide active control of beam direction rather than a passive verification system. Williams then discussed the current development of real time tumour tracking to counteract patient movement by The British Journal of Radiology, January 2006

Commentary: President’s conference 2005

mounting a diagnostic machine with fluoroscopic, radiographic and CT capabilities onto the treatment linear accelerator. Examples of improved set-up were shown for lung and bladder treatments – image-guided radiotherapy will certainly make things better! For the future, although physicists and engineers are not yet spent (vide the next topic of proton therapy), they will need help from other disciplines, e.g. molecular biologists and geneticists (biological targeting for anoxia and metabolism, and selective targeting of tumour cells), and from radiobiologists (for example to exploit the information on bystander effects coming from microbeam studies). As a fitting sequel to the Silvanus Thomson Memorial Lecture, Bleddyn Jones presented the case for particle therapy, especially with protons. The theoretical advantages of using the Bragg dose peak to improve the therapeutic ratio have been known for many years. Unfortunately, for a 60 MeV beam the peak is at only 3 cm depth and treatment is limited to quite superficial tumours. Notwithstanding, over 1200 choroidal melanomas have been treated successfully at the Clatterbridge Hospital. Work by Lomax et al [15] has shown that for treatment of the breast and regional nodes, a 9-field photon IMRT approach can either produce similar dose homogeneity across the planning treatment volumes to that of a proton plan, or similar sparing of dose to both lungs and the heart, but not both. Jones estimated that 10–20% of patients might be better treated by particle radiotherapy and believes that technical improvements in physics, bioengineering and computing, especially in robotics and particle delivery, now make treatment with a 200 MeV beam, with Bragg peak depths approaching 20 cm, a practical proposition. It is anticipated that this will lead to a big increase in demand for particle therapy in the UK [16]. The Conference concluded with two further papers in diagnostic imaging. Catherine Owens gave a wide-ranging review of the changing practice of paediatric imaging. The diagnostic capability and accuracy of multidetector CT (MDCT) angiography was compared with echocardiography, cardiac catheterization and surgery in the assessment of the great vessels in 40 consecutive patients (mean age 5 years) with congenital heart disease. MDCT was accurate, showing good agreement with interventional catheter and surgery and provided additional information. Effective doses of radiation were low – ranging from 0.97 mSv in neonates to 1.7 mSv in adolescents [17]. Magnetic resonance coronary angiography and lateenhancement imaging have been shown to be feasible in children who had undergone arterial switch for transposition of the great arteries. Diagnostic quality images were acquired in 72% of the coronary arteries imaged and this rose to 100% in subjects over 10 years old [18]. Finally, Peter Ell discussed the contribution of PET/CT to improved patient management. Whilst acknowledging the important contribution in neurology and cardiology, in the limited time available and in the context of the Conference, Ell concentrated on oncology. Four distinct areas were covered, diagnosis, staging, radiotherapy planning and treatment monitoring. Two very different challenges for this wonderful technique were highlighted. At the cutting edge of research The British Journal of Radiology, January 2006

there are almost unlimited opportunities for PET/CT to be used to assess the biology of individual response to treatment [19]. Whilst recognizing the importance of F-18 fluorodeoxyglucose in oncology, Ell emphasised the need to look at a wide range of other novel markers that are being developed, aimed at imaging proliferation [20, 21], hypoxia, angiogenesis, apoptosis, etc. At the other extreme there is the huge problem of diffusion of technology in a cost-effective way so that, on a day-to-day basis, many more of the millions of cancer sufferers can benefit from the power of multimodality imaging. Ell’s concluding remarks were:

N N N N

PET/CT has changed patient management; It is best at assessing extent and severity of cancer; It informs radiotherapy planning; and It combines the power of CT with the unique metabolic mapping obtained with PET.

These remarks were, of course, addressed to PET/CT but, in many respects, with suitable changes of wording, could be applied to the impact of other technological advances discussed during the 2005 President’s Conference. We commend to you the full articles contributed by the speakers in this issue of the Journal.

Acknowledgments I am grateful to Fergus Gleeson and Gu¨nter Dombrowe for helpful contributions to this Commentary.

References 1. Hounsfield GN. Computerised transverse axial scanning (tomography). Part 1 description of system. Br J Radiol 1973;46:1016–22. 2. Ambrose J. Computerised transverse axial scanning (tomography). Part 2 clinical application. Br J Radiol 1973;46:1023–47. 3. Shardt P, Deuringer J, Freudenberger J, Hall E, Knipfer W, Mattern D, et al. New X-ray tube performance in computed tomography by introducing the rotating envelope tube technology. Med Phys 2004;31:2699–706. 4. Kalender WA, Wolf H, Seuss C. Dose reduction in CT by an anatomically adapted tube current modulation. Med Phys 1999;26:2248–53. 5. Greess HR, Wolf H, Suess C, Lutze J, Kalender WA, Bautz WA. Automatic exposure control to reduce dose in subsecond multislice spiral CT – Phantom measurements and clinical results. Radiology 2002;225 Suppl. RSNA programme p 593. 6. Kalender WA. Thin-section three dimensional spiral CT. Is isotropic imaging possible? Radiology 1995;197:578–80. 7. Kalender WA, Klotz E, Suess C. Vertebral bone mineral analysis: an integrated approach with CT. Radiology 1987;164:419–23. 8. See TC, Ng CS, Watson CJE, Dixon AK. Appendicitis: spectrum of appearances in helical CT. Br J Radiol 2002;75:775–81. 9. Ng CS, Watson CJE, Palmer CR, See RC, Beharry NA, Housden BA, et al. Evaluation of early abdominopelvic computed tomography in patients with acute abdominal pain of unknown cause – prospective randomised study. BMJ 2002;325:1387–9. 10. Girshman J, Wolff SD. Techniques for quantifying coronary artery calcification. Semin Ultrasound CT MR 2003;24:33–8.

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P P Dendy 11. Thompson GR, Partridge J. Coronary calcification score: the coronary-risk impact factor. Lancet 2004;363: 557–9. 12. White CS, Kuo D, Keleman M, Jain V, Musk A, Zaidi E, et al. Chest pain evaluation in the emergency department; can MDCT provide a comprehensive evaluation? AJR Am J Roentgenol 2005;185:533–40. 13. Williams PC. IMRT: delivery techniques and quality assurance. Br J Radiol 2003;76:766–76. 14. James HV, Scrase CD, Poynter AJ. Practical experience with intensity modulated radiotherapy. Br J Radiol 2004;77:3–14. 15. Lomax AJ, Cella L, Weber D, Kurtz JM, Mirabell R. Potential role of intensity-modulated photons and protons in the treatment of the breast and regional nodes. Int J Radiat Oncol Biol Phys 2003;55:785–92. 16. Jones B, Burnet NG, Price P, Roberts JT. Modelling the expected increase in demand for particle therapy: implications for the UK. Br J Radiol 2005;78:832–5.

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17. Benson C, Taylor A, Ross UG, et al. Three-dimensional anatomy of the great vessels defined by 16-slice multi-detector CT angiography in neonates, infants, children and adolescents with congenital heart disease. Presented at the 42nd Congress of the European Society for Paediatric Radiology, Dublin, June 2005. 18. Taylor AM, Dymarkowski S, Hamaerkers P, et al. MR coronary angiography and late-enhancement myocardial MR in children who underwent arterial switch surgery for transposition of great arteries. Radiology 2005;234:542–7. 19. Bugarolas J, Clark JW, Chabner B. Using ‘‘rationally designed drugs’’ rationally. Lancet 2003;361:1758–9. 20. Shields AF, Grierson JR, Dohmen BM, et al. Imaging in vivo proliferation with 18FLT and positron emission tomography. Nature Medicine 1998;11:1334–6. 21. Francis DL, Visvikis D, Costa DC, Croasdale I, Arulampalam TH, Luthra SK, et al. Assessment of recurrent colorectal cancer following 5-fluorouracil chemotherapy using both 18FDG and 18FLT. Eur J Nucl Med Mol Imaging 2004;31:928.

The British Journal of Radiology, January 2006


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President’s conference paper

CT scanning the early days E C BECKMANN, BSc(Eng) Lanmark, Beaconsfield, Bucks, UK

Abstract. CT scanning has become an established diagnostic tool within the radiology department. This article covers some of the history of the development and early days of CT scanning. It is based upon the lecture given on the Memorial Day for Sir Godfrey Hounsfield during the British Institute of Radiology President’s Conference 2005.

It is less than 34 years ago, on 20th April 1972, that an unknown engineer from EMI Ltd, the company better known at the time for publishing the Beatles records, gave a presentation at the 32nd Congress of the British Institute of Radiology. The Engineer, Godfrey Hounsfield, was lecturing with Dr James Ambrose from Atkinson Morley’s Hospital on ‘‘Computerised Axial Tomography (A new means of demonstrating some of the soft tissue structures of the brain without the use of contrast media)’’ [1, 2]. Many people attending that BIR congress will never forget the experience of hearing a presentation on CT scanning for the first time. In fact Hounsfield had presented the results of some of his animal experiments the previous year at the 2nd congress of the European Association of Radiology, in Amsterdam, but they had not excited much interest. The same might have happened in the USA because at a Neuro Postgraduate Course at the Albert Einstein College of Medicine, New York, on Monday 15th May 1972, only about a dozen people stayed to hear an extra lunchtime lecture by Hounsfield and Dr Bull, where they showed the first clinical images. However these people realised the significance of what they had seen and the news spread rapidly.

The beginning In the mid 1960s Hounsfield was working on the pattern recognition of letters when he began to consider whether he could reconstruct a three-dimensional representation of the contents of a box from a set of readings taken through the box at randomly selected directions. He found that by considering the three-dimensional object within the box as a series of slices, reconstruction was easier than treating the content as a volume. He tested the theoretical principal by working with a matrix of numbers set to zero with a square in the middle where each number was set at 1000. He entered these data into a computer programme to get simulated absorption values and then reconstructed the picture using another programme. Hounsfield recalled his surprise at how accurate the result was.

Here he stated ‘‘The purpose of the study was to investigate the employment of a computer to make better use of the information obtained when an object is examined by gamma rays or X-rays’’. In this proposal Hounsfield compared the classic conventional X-ray technique producing a confused and fuzzy picture to the clear outline produced by the proposed system. Hounsfield proposed a system as shown in Figure 1 based upon reconstructing pictures of slices through an object and in detailing the expected benefits he indicated a theoretical accuracy of detection better than 1%.

The lathe bed model The initial test rig was built on the bed of an old lathe which Hounsfield had been using in a previous project working on computer stores. Hence the early test unit became referred to as the ‘‘Lathe bed model’’. The initial rig utilized a gamma source, Americium 95, with a photon counter as the detector. On this rig, the source made 160 traverses of the object, which was rotated 1 ˚ at the end of each traverse for a total of 180 ˚. It took 9 days to collect sufficient information, and 2.5 h to reconstruct the image on an ICL 1905 mainframe computer. However, the resultant images proved the feasibility of the technique and with the replacement of the gamma source by an X-ray source as shown in Figure 2, the scanning time was reduced to 9 h. Initial images were of inert objects, then specimens from an abattoir, including bullocks brains and pigs bodies as shown in Figure 3. Due to the long scan times, particularly with the gamma source, many of these specimens decayed while the

The project proposal Once Hounsfield had proved the theoretical principle he went on to generate the original project proposal in 1968. Received 12 September 2005 and accepted 16 September 2005.


Figure 1. Extract of the original 1968 project proposal. 5

E C Beckmann

Figure 2. The original lathe bed model (copyright EMI Ltd).

Figure 4. First image of a brain specimen.

pictures were acquired producing gas bubbles which caused artefacts in the images. This initial work was done by a very small team comprising Hounsfield, Stephen Bates (programming), Peter Langstone (electronics) and Mel King (mechanics) working on a very low budget of £25 000. Dr James Ambrose recalls that, in about 1969, he received a call from an old acquaintance, Dr Evan Lennon then principal medical officer in Radiology at the Department of Health, asking if he would see a man called ‘‘Godfrey Hounsfield’’ and listen to him. Lennon had found him confusing but was reluctant to dismiss him as a crank (Ambrose later learnt that other eminent radiologists had already dismissed him as a crank!). Ambrose recalls that when he and his senior physicist Dr John Perry met Hounsfield, the conversation was difficult. Hounsfield would only say that the method was fundamentally different from other methods of X-ray imaging, more efficient in photon usage and likely to be more sensitive to small density variations. In order to demonstrate a clinically relevant image, Ambrose arranged for a bottled specimen of a brain to be borrowed from a museum and was amazed at the image Hounsfield showed him 5 weeks later. An image of the first brain scanned is shown in Figure 4.

the foresight to place an order for a machine with a theoretical specification which included a 4–5 min scan time and an 0.5% pixel accuracy, and this enabled the project to continue. This order was for a prototype and three clinical machines that would generate sufficient income to fund a fifth machine for Hounsfield and his team to keep and work on. The Department of Health order would also fund half the remaining research costs and in exchange they would receive a small royalty on sales. At the time it was calculated that it would cost £69 000 to build a complete working system and so it was agreed that the Department of Health would pay £150 000 for each of the four systems.

Building the prototype Having shown some clinically interesting images the project was then ready to move to the next stage of building a full prototype. However funding was an issue. It was Gordon Higson at the Department of Health who had

Figure 3. Early scan of a pig. 6

The first clinical patient The prototype was installed at Atkinson Morley’s Hospital in South London where the first patient, a middle aged lady with a suspected frontal lobe tumour, was scanned on 1st October 1971. The surgeon who operated on her shortly afterwards reported that ‘‘it looks exactly like the picture’’ shown in Figure 5. Hounsfield remained cautious. He recalled ‘‘I’ve had this before, first time is always lucky and then everything else goes wrong after that. So I thought, the next ones are not going to be any good, but they did another ten more patients and every one of them came out as being obvious diseases of the brain showing up in various forms. Dr Ambrose found that, by injecting iodine-based contrast agent that would localize the particular spot where the tumour was and it showed up even better’’. Hounsfield took some of the contrast enhanced images and subtracted without contrast images to compare the blood flow on either side of the brain. In the original system the patient’s head was placed in a rubber cap surrounded by water. This water bag was used to reduce the dynamic range of the detected X-rays and improve the absolute values of the attenuation figures. Using one sodium iodide (NaI) crystal and photomultiplier tube detector per slice, plus one as a reference detector with a scan time of 4.5–20 min per 180 ˚ scan, the system acquired two contiguous slices per scan each with a 80680 matrix of 3 mm63 mm613 mm voxels. Early images showed the ability to meet the pixel density accuracy of 0.5% in the absorption coefficient as defined in the theoretical specification. The British Journal of Radiology, January 2006

CT scanning the early days

done at a central processing unit on a suitable large and fast main frame machine. But the introduction of the mini computer and the implementation of the new improved reconstruction algorithms were to change this.

CT1010 scanner A challenge with the original EMI Mk1 scanner was the water bag, both as regards the ease of use with patients and also due to the occasional water leak! Replacement of the water bag with shaped carbon fibre wedges and bean bags was a significant improvement. This was further enhanced by the increase to eight detectors per slice in the CT1010 which was still a two contiguous slice scanner offering 1606160 and 3206320 matrix sizes over a 210 mm scan diameter and with the minimum scan time improved to 1 min. The prototype of this system was installed in 1975 at Atkinson Morley’s Hospital and showed significant improvement in clinical image quality.

Body scanning

Figure 5. First patient image scanned on the prototype EMI scanner at Atkinson Morley’s Hospital on 1st October 1971.

The three systems ordered by the Department of Health were installed at the National Hospital for Neurology and Neurosurgery in London, Manchester and Glasgow. After this, the first CT scanners were installed in the USA at the Massachusetts General Hospital and the Mayo Clinic, where the first scan in the USA was done on 19th June 1973.

Reconstructing the picture Early scan data were actually taken back to EMI on tape for processing overnight which took 20 min per image on an ICL 1905 computer. In production this was done on a minicomputer which fortuitously had emerged at the right time. Images were taken back the next day on tape to Atkinson Morley’s Hospital to be displayed. The early images were displayed in three ways; paper printout, cathode ray tube (CRT) display or as a Polaroid picture of the CRT display. The early images were generated using iterative algebraic reconstruction implemented by Steve Bates on the ICL 1905 mainframe. Subsequently reconstruction used the filtered back projection or convolution method invented and patented by Chris Lemay, one of the many patents filed and held by Hounsfield and his team. On the original EMI Mk1 scanner an 80680 image took 7 min to process, with filtered back projection on the same computer a 1606160 image could be processed in 30 s after the end of the scan. It had been thought that image reconstruction and processing was so complicated that it would have to be The British Journal of Radiology, January 2006

The feasibility of body scanning was proved when a slim member of the EMI team, Tony Williams, was scanned in a head scanner. The first body images taken in the body prototype machine were of Hounsfield himself on 20th December 1974. The first body images were shown to a meeting at the first International Conference on CT Scanning in Bermuda on Friday 14th March 1975, one of these images is shown in Figure 6. All the research machines were named after stones: Opal, Pearl, Garnet and the body prototype was Emerald. This Emerald system was first installed clinically at Northwick Park Hospital in March 1975. The first body scan carried out in the USA was in October 1975 at the Mallinkrodt Institute St Louis. Dr Ron Evans recalled that this was a jaundiced patient, in whom it had been difficult to differentiate between medical and surgical jaundice. The CT scans showed that it was surgical jaundice which was subsequently clinically confirmed. Initially known as the CT5000, the body scanner was developed into the commercial production machine, the CT5005. These body scanners were single slice machines using a gantry with 30 detectors plus a reference detector to reduce scan time to 20 s. The matrix had been increased to 3206320 over a selectable 240 mm, 320 mm or 400 mm scan field.

The generation game All these early scanners were the so called 1st or 2nd generation utilizing the translate/rotate technology where the gantry scanned across the patient before indexing by one degree and scanning back. An early problem in CT scanner design was detector stabilization and the need for calibration. The EMI scanners were using NaI crystal photon detectors and photo multiplier tubes, and the translate/rotate technology enabled detector calibration by taking air readings at the end of each translate movement. This gave high accuracy but limited the speed of the scan. By 1976 there were 17 companies offering CT scanners with 3rd generation rotate/rotate scanners having 7

E C Beckmann

Figure 6. Body scan of Hounsfield taken on the prototype scanner in the laboratories and shown at Bermuda conference on 14th March 1975.

been introduced, to offer fast scan times, most based upon xenon gas detectors arranged in an arc [3]. Hounsfield realised the need for a system that was faster than translate/rotate and that could overcome the calibration and artefact issues of rotate/rotate systems.

Topaz The patent for a scanning focus system to produce a true volume scanner was filed on 19th October 1976. The Topaz research system, also named after a stone and shown in Figure 7, was a 3rd generation system with a flying X-ray spot. The X-ray flying spot scanned in a direction opposite to the direction of rotation of the machine which meant that the body could be scanned with arcs of detector readings which overlapped in such a way that they could be compared and continuously calibrated. Built with 612 detectors including a central zoom region, Topaz had a resolution in the x-y plane of 0.65 mm. Volume scans taken in June 1980 were displayed in three dimensions in real time as 1200612006270 pixels.

Recognition Initially the scale for describing the attenuation coefficients was referred to as EMI numbers. This was then expanded by a factor of two and became known as Hounsfield units (H) where H~

ktissue {kwater |1000 kwater

and m is the linear attenuation coefficient. Each Hounsfield unit is equivalent to 0.1% of the attenuation of water [3]. In addition to giving his name to the unit of attenuation, Hounsfield received many awards including the BJR Barclay prize jointly with Ambrose in 1974, the Nobel Prize for Physiology or Medicine in 1979 [4] and a Knighthood in 1981. 8

Figure 7. Topaz 3rd generation flying focal spot scanner.

Hounsfield and his team created the CT scanner, which has had an explosive impact on diagnostic radiology, with little money and few resources. By the end of the 1970s they already had plans for many of the technologies which were to develop the CT scanner over the next 30 years, including helical multislice scanners and high power continuously rated scanned beam X-ray tubes. They developed many of the techniques which formed the foundation of modern imaging including image subtraction. By 1976 the reconstruction techniques used in CT were already being applied to other areas including ultrasound and nuclear magnetic resonance.

Acknowledgments The author is indebted to many people especially those members of the original EMI team who worked with Sir Godfrey Hounsfield for their input to the original lecture and material used in this article.

References 1. Hounsfield GN. Computerised transverse axial scanning (tomography): Part 1. Description of system. Br J Radiol 1973;46:1016–22. 2. Ambrose J. Computerised transverse axial scanning (tomography): Part 2. Clinical application. Br J Radiol 1973;46:1023–47. 3. Brooks RA, Di Chiro G. Principles of computer assisted tomography (CAT) in radiographic and radioisotropic imaging. Phys Med Biol 1976;21:689–732. 4. Computed medical imaging. Nobel lectures in physiology or medicine 1971–1980; 568–86.



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Cardiac applications of multislice computed tomography 1

A DE ROOS, MD, 1L J M KROFT, MD, 2J J BAX, MD, 1H J LAMB, MD and 1J GELEIJNS, PhD

Departments of 1Radiology and 2Cardiology, Leiden University Medical Center, Albinusdreef 2, 2333 ZA Leiden, The Netherlands

Multislice CT (MSCT) is gaining clinical acceptance for cardiac imaging owing to improved temporal and spatial resolution of the latest 16-slice and 64-slice technology. Although the cardiac MSCT applications are promising, there is still room for further technical improvements and optimization of post-processing techniques for cardiac evaluation. Interestingly, the data acquired for CT angiography of the coronary arteries can also be used to create volumetric cine loops of cardiac function. The functional data are available without the need for repeat scanning or for administration of additional contrast material [1]. Furthermore, MSCT allows assessment of first-pass perfusion and delayed enhancement imaging in patients with subacute myocardial infarction. Recently, it has been reported that MSCT reveals microvascular obstruction or the so-called no-reflow phenomenon as a late perfusion defect in patients with re-perfused acute infarctions, similar to observations made by other techniques like MRI [2]. With further development MSCT may allow combined assessment of the presence and extent of coronary atherosclerosis, the percent diameter stenosis, plaque characterization and the effect of the lesion on perfusion and myocardial function. In this review, the technical requirements of cardiac MSCT and some frequent clinical applications are discussed.

MSCT imaging requirements Requirements for cardiac MSCT image acquisition depend strongly on the clinical problem. For example, CT coronary angiography requires excellent spatial and temporal resolution, whereas only modest spatial and temporal resolution is sufficient for the assessment of the anatomy of pulmonary veins and the left atrium. In general, the higher the requirements for image quality become, the more complex the acquisition, the longer scan time and the higher patient dose. Main aspects with regard to imaging performance are low-contrast and spatial resolution, temporal resolution, and scan time. Patient dose and radiation risk should always be considered as the counterpart of image acquisition and image quality.

Low-contrast resolution and spatial resolution Low-contrast resolution is the ability to visualize structures that demonstrate only a small difference in Received 22 September 2005 and accepted 5 October 2005.


Hounsfield units compared with their direct environment. In cardiac applications of CT, native tissue contrasts are in general not sufficient to differentiate between, for example, the vessel wall and its unenhanced lumen, or the heart and the inner chambers. Contrast enhancement is thus mandatory for visualizing the lumen of coronary arteries, the heart chambers, pathology of the myocardium or anatomy of pulmonary veins. Low-contrast resolution depends on tube current (mA), the reconstructed slice thickness, tube voltage, beam filtration and the reconstruction algorithm, and is strongly correlated to radiation exposure. In general, low-contrast resolution performance of CT scanners is not a limitation for the application of cardiac CT. Spatial resolution, or high-contrast resolution, determines the ability to visualize contours of small structures within the scanned volume. Small objects can only be resolved when there is a rather large contrast with the direct environment. Considerable improvement of spatial resolution in clinical acquisitions was achieved with the latest generations of multislice CT scanners. This is of importance, particularly for the application of CT coronary calcification scoring and CT coronary angiography. The actual diameters of the lumen of normal coronary artery segments range from 5 mm in the proximal segments to less than 1 mm in the distal segments [3]. This means that spatial resolution of 1.0 mm in all three dimensions should be sufficient for imaging of the coronary arteries, except for distal segments that would require a spatial resolution of at least 0.5 mm. Bypass graft diameter typically ranges from 4 mm to 6 mm. A spatial resolution of 2 mm3 (voxel size) might thus be sufficient for imaging the lumen of bypass grafts. For imaging of small structures within the coronary arteries, such as atherosclerotic plaque and stents, excellent spatial resolution, even better than 0.5 mm3, might be required. Voxel size is often used as an indicator of spatial resolution. However, voxel size should be interpreted with care since smaller voxel size does not necessarily imply better spatial resolution. Spatial resolution is preferably expressed as the response of a delta-function; in CT, this response is either called a pointspread-function (spatial resolution in the axial plane) or a slice sensitivity profile (spatial resolution along the z-axis). Spatial resolution is limited by the acquisition geometry of the CT scanner, the reconstruction algorithm and the reconstructed slice thickness. The performance of current 64-slice scanners with regard to spatial resolution, expressed as the full-width half-maximum of the response of a delta-function, is within the range 0.6–1.0 mm in all three dimensions. 9

A de Roos, L J M Kroft, J J Bax et al

Temporal resolution Temporal resolution determines whether fast moving objects can be resolved in the CT image. Good temporal resolution limits motion artefacts and blurring of the image. Principally, good temporal resolution can be achieved by a short reconstruction window providing snap shots of the beating heart and coronary arteries. Good temporal resolution in cardiac CT is realised by fast data acquisition (fast rotation of the X-ray tube), but even more importantly by a dedicated reconstruction algorithm. A recent paper [4] provides information on the rest period of the heart, which is a measure for the required reconstruction window. The rest period is defined as the time during which the 3D motion of a coronary artery is less than 1 mm. It was reported that, for patients with a heart rate of 64¡9 beats per minute (BPM), the endsystolic rest period duration was 76¡34 ms; and the middiastolic rest period duration was 65¡42 ms for the proximal to middle segments of the right coronary artery. For the left coronary artery tree, the end-systolic rest period duration was 80¡25 ms; the mid-diastolic rest period duration 112¡42 ms. From these data it is concluded that the duration of a ‘‘snap shot’’ of the coronary arteries, or in other words the reconstruction window, should be shorter than 65–110 ms. This is in good agreement with earlier papers; in one paper it is suggested that the reconstruction window should be lower than 100 ms for coronary angiography in mid-diastole at 62¡10 BPM [5], and in another paper it is stated that a 100 ms reconstruction window is relatively optimal for most patients at heart rates up to 90 BPM [6]. All of these considerations assume image reconstruction at the cardiac phase point that is associated with least motion, e.g. a reconstruction window starting between 60% and 80% of the interval between two consecutive R-waves. More strict criteria for the reconstruction window apply if the heart should be assessed at more than one cardiac phase point, including those that are associated with rapid movement of the heart wall, e.g. for studying the dynamics of the myocardium. More strict criteria apply as well when a 1 mm displacement of a coronary artery within the duration of the snap shot becomes unacceptable. This may happen, for example, when imaging small distal parts of the coronary arteries, quantifying coronary stenoses and assessment of coronary atherosclerotic plaque. General reconstruction algorithms that are used for general CT applications provide, in principle, a temporal resolution equal to the rotation time (360 ˚ rotation, full reconstruction), the best achievable temporal resolution with general reconstruction algorithms is slightly longer than 50% of the rotation time (180 ˚ rotation, half reconstruction). Current 64-slice scanners that are used for cardiac applications provide a rotation time of 330– 400 ms. These typical rotation times are not short enough for achieving a 100 ms or shorter snap-shot of the heart, even if a 180 ˚ rotation half-reconstruction is applied. Therefore, dedicated reconstruction algorithms are used in cardiac CT that allow for reconstruction of synchronized images from transmission data acquired during two or more successive heart cycles according to a method described already in 1977 [7]. These so-called segmented (multicycle) reconstruction algorithms allow for merging synchronized transmission data from successive heart 10

cycles. The more heart cycles that can be included in the reconstruction, the better the temporal resolution. A low pitch factor, which is typical for cardiac CT acquisition, is required to acquire data from more than one heart cycle. A pitch factor as low as 0.2 is required to record at least two heart cycles and to achieve a temporal resolution in the order of magnitude of 100 ms for typical heart rates between 60–80 BPM. Figure 1 shows, as an example, the temporal resolution that is achievable with a reconstruction algorithm that can merge transmission data from an unlimited number of heart cycles. The figure illustrates the dependence of the reconstruction window on rotation time and heart rate and was calculated for a pitch factor of 0.2. From Figure 1 it can be concluded that, for achieving the shortest reconstruction window, rotation time should be adapted to the heart rate.

Scan time Scan time is the time interval between the start and the end of one acquisition, sometimes referred to as a sequence. To avoid breathing artefacts and to limit the amount of contrast material in contrast enhanced acquisitions, scan time in cardiac CT should remain at least below 30 s, but preferably below 20 s. The extent of the target volume, as well as acquisition parameters such as rotation time, pitch factor, section thickness and number of simultaneously acquired sections, define scan time. In general a much shorter scan time than 20 s can now be realised with the current generation of 64-slice scanners for typical cardiac CT examinations; for example, a typical

Figure 1. Temporal resolution of CT coronary angiography. The temporal resolution depends strongly on the rotation time and the reconstruction algorithm. In segmented (multiphase) reconstructions, temporal resolution depends also strongly on the pitch factor. The lower the pitch factor, the more cardiac phases are captured during the acquisition and the better temporal resolution. The graphs are calculated for a pitch factor of 0.2. The graphs clearly show the dependence of temporal resolution on heart rate and rotation time.



Half reconstruction 250 2.7 120 0.83 300 100 0.4 6460.5

100 Multisegmental 9.0 120 0.19 40 120 0.4 1662.0

100 Multisegmental 17.3 240 0.19 300 120 0.4 6460.5

100 Multisegmental 9.5 120 0.19 300 120 0.4 6460.5

Half reconstruction 250 2.5 Not applicable 120 200 120 0.25a

Partial rotation. 60 beats per minute assumed. b

a

2.5/40 Spiral

5/70

4/100 Spiral

Pulmonary veins Spiral

4/100

Tube Pitch Acquisition Rotation Tube configuration times (s) voltage (kV) current (mA) (n6T mm)

463

Prospective triggering Retrospective gating Retrospective gating Retrospective gating No synchronization Spiral

Coronary artery calcification is a marker for atherosclerotic lesions in the coronary arteries. The amount of coronary artery calcification is correlated to the risk of coronary events. However absence of coronary artery calcification does not rule out atherosclerosis. Applications

CTA coronary arteries CTA coronary bypass RV function

Quantitative assessment of coronary artery calcification

Synchronization Acquisition Contrast (ml s-1, ml)

MSCT provides special opportunities for cardiovascular CT in addition to angiography of the coronary arteries and coronary bypass grafts. These options include assessment of left ventricular (LV) and right ventricular (RV) function, coronary calcification score, myocardial infarction imaging and assessment of the anatomy of pulmonary veins in patients with atrial fibrillation. Each of these applications can be characterized by their specific techniques for acquisition and reconstruction. Table 1 provides information about typical acquisition and reconstruction parameters for some clinically established cardiac CT applications.

Examination

Clinical applications

Table 1. Typical acquisition and reconstruction characteristics of some cardiac CT examinations

Radiation protection of patients is based on justification and optimization. Justification implies that the benefit for the patient outweighs the risk of radiation exposure. Patient dose assessment is required for balancing harm and benefit of the CT examination and to assess the effect of measures for optimization of cardiac CT. Nowadays, most CT scanners provide the user with an indication of patient dose in the form of the CT dose index (CTDI) and dose–length product (DLP). Effective dose can be derived from these dose quantities. Effective dose from cardiac CT coronary angiography is relatively high, mainly due to the need to catch more than one cardiac cycle and the resulting low pitch factor. On the other hand, effective dose from an ungated acquisition, such as in ungated pulmonary vein CT angiography, is relatively low due to the high pitch factor. Effective dose for calcium scoring, assessment of ventricle function or pulmonary veins is in the range 1–3 mSv, effective dose for CT coronary angiography is considerably higher, e.g. in the range 10– 15 mSv. Concern about radiation exposure stimulates the development of methods for dose reduction in cardiac CT coronary angiography. The field of view of interest in cardiac CT is rather small and therefore radiation exposure of tissue outside this field of view can be limited by means of a special ‘‘small field’’ beamshaping filter. Another method for dose reduction is to reduce X-ray output during the systolic phases that are expected to be of less interest for the evaluation of the coronary arteries (ECG triggered modulation of dose). Pitfalls of small field scanning are the occurrence of artefacts and reduced image quality. A pitfall of tube modulation is reduced image quality at certain relevant phases of cardiac cycle, e.g. due to an irregular heart rate.

Scan range Scan Reconstruction (mm) time (s) algorithm

Patient dose in MSCT

none

scan time for calcium scoring is 2.5 s, for coronary angiography 10 s and for an ungated acquisition of the pulmonary veins 3.0 s.

Calcium scoring Sequential

Reconstruction windowb (ms)

Cardiac MSCT

11


of quantitative assessment of coronary artery calcification are screening of asymptomatic individuals with risk factors for coronary artery disease and follow-up of patients who received medication for the treatment of coronary artery disease. Coronary artery calcification is well visualized with X-ray techniques such as radiography but only CT provides a non-invasive method for detecting and quantifying coronary artery calcification [8]. Coronary calcification is best detected and measured in a plain CT acquisition without contrast enhancement. Quantification of coronary calcium was introduced in 1990 by Agatston et al [9]. They used electron beam tomography and established the ‘‘Agatston score’’. The Agatston score requires an acquisition with a special protocol (3 mm contiguous slices, 130 kV). The Agatston score is achieved by setting a threshold for the Hounsfield unit (130 HU) and for the size of the lesion (1 mm2). Then a pragmatic weighting of the calcified area is applied depending on the maximum HU in the lesions for each image. The total calcium score is calculated by summing the weighted areas for all images (Figure 2). With the introduction of MSCT, new acquisition protocols came into use; prospective ECG triggering in combination with a half (180 ˚) reconstruction at 120 kV is now generally used for calcium scoring. In prospective ECG triggered MSCT acquisitions, the patient is only exposed within the 170–200 ms acquisition window at diastole and radiation exposure is therefore significantly less compared with retrospective gated MSCT cardiovascular examinations. The application of MSCT for quantification of coronary calcium made it mandatory

Figure 2. Coronary artery calcification imaging at 64-row multidetector CT (MDCT). 64-row MDCT of a 52-year-old male patient with risk factors for coronary artery disease. Small calcifications in the left anterior descending artery. The total calcium score according to Agatston was 21, and the total volumetric score was 25, indicating mild atherosclerotic plaque with mild or minimal coronary artery narrowings likely. CT-angiography revealed no coronary artery stenoses. 12

to switch to new quantification methods that can be compared for different scanners and that are robust with respect to different scanners and acquisition protocols. Alternatives for the Agatston score are the volume score (the volume of all voxels exceeding a certain threshold) and calcium mass (mg) [10]. The latter quantity holds the promise of providing the best physical measure for coronary artery calcification. Unfortunately, there is still a lack of standardization of the MSCT techniques with regard to image acquisition as well as to the methodologies for quantitative coronary calcification scoring. The development of standardized and reproducible protocols is a technical prerequisite for coronary calcification scoring to become a useful clinical tool. In addition, for screening purposes, the coronary calcification score will have to be established as an independent predictor of existing risk factors for cardiovascular disease [11].

Coronary angiography MSCT has rapidly evolved through different stages of technological innovation, allowing high-quality non-invasive 3D imaging of coronary artery morphology (Figures 3 and 4). Recently the diagnostic accuracy of 64-slice MSCT for the identification and quantification of coronary artery stenoses has been reported [12, 13]. The patient-based analysis revealed that 94% of patients who required revascularization were correctly diagnosed by CT. Although excellent accuracy for stenosis detection was noted, technical restrictions for exact quantification of the degree of stenosis and reliable visualization of small vessel segments remain [12]. In an accompanying editorial the authors express the expectation that MSCT will be used in the near future on a routine basis for the identification of patients who do not need revascularization therapy despite the presence of symptoms [14]. The potential value of MSCT for stenosis quantification is currently under active investigation. Recently, a good correlation between MSCT and quantitative coronary X-ray angiography was shown for stenosis quantification with the use of 16-slice technology, although MSCT revealed a systematic overestimation as compared with the reference standard [15]. Perfusion defects related to previous myocardial infarction or ischaemia may be well visualized with the use of MSCT (Figure 5). In CT coronary angiography, beta-blockers may be used to reduce the heart rate to a lower range, e.g. to 50– 60 BPM to increase the cardiac rest period and with this to reduce motion artefacts. The resulting imaging performance is more predictable and of more consistent quality when using such medication. Special reconstruction algorithms for the reconstruction pose an alternative to the use of medication. The segmented reconstruction algorithm yields good temporal resolution even at higher heart rates. Also, when total scanning time is short, e.g. below 10 s, the quality of the scan improves since, due to the reduction of the total amount of heart beats in the scan, less variation can be expected in the heart rate during the acquisition. Hyperventilation and administration of oxygen may be used to stabilize heart rate particularly at scan times of approximately 20 s scanning time or longer. The British Journal of Radiology, January 2006

Cardiac MSCT

Figure 3. Normal coronary artery anatomy at 64-row multidetector CT (MDCT). 64-row MDCT of a 62-year-old male patient with risk factors for coronary artery stenosis. No stenoses were found at MDCT coronary angiography. Left anterior (a) oblique view and (b) caudal view. LAD, left anterior descending coronary artery; D, diagonal branch of the LAD; IM, intermediate coronary artery branch; Cx, circumflex coronary artery; MO, obtuse marginal branch (of the Cx); DP, descending posterior branch (of the right coronary artery).

Figure 4. Bypass imaging at 64-row multidetector CT (MDCT). 64-row MDCT of a 78-year-old male patient after coronary artery bypass graft operation (CABG). Occlusion of multiple venous bypass grafts (nr 1 in a). Left internal mammarian artery bypass graft (nr 2 in a,b) with open anastomosis (nr 3 in a,b,c) on the left anterior descending coronary artery (nr 4 in a,c). Poor quality native coronary artery system with multiple stenoses and poor contrast enhancement (nr 4 in a,c). b and c are displayed in two perpendicular longitudinal directions. The British Journal of Radiology, January 2006

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Figure 5. Multiple perfusion defects imaged with 64-row multidetector CT (MDCT). Same patient (78-year-old male) as in Figure 4 after coronary artery bypass graft operation and multiple venous bypass graft occlusions. Multiple perfusion defects with regional wall thinning.

Assessment of ventricular function With retrospective gated 180 ˚ segmented sinogram space reconstruction the data can be reconstructed for evaluation of ventricular function [16]. Diastolic and systolic images can easily be extracted and reconstructed in any orientation for functional evaluation (Figure 6). Global ventricular function is generally measured as the end-systolic and end-diastolic volume (ESV, EDV). Subsequently, stroke volume (SV) and ejection fraction (EF) can easily be derived from ESV and EDV. Semiautomatic software may be used for ventricular cavity contour detection and for the calculation of global ventricular function. Regional LV wall motion

can be assessed by visual scoring of cinematic loops of well described myocardial segments [17]. Integrated CT assessment of the coronary arteries and regional myocardial function allows assessment of the functional consequences of a coronary artery stenosis leading to ischaemia and contraction abnormalities. The usefulness of this combined approach has been reported in patients with hypertension and diabetes mellitus [18, 19]. From the same data set global function and left ventricular mass can also be determined, which have clinical relevance in patients with hypertension for prognosis and guidance of therapy. Several studies have shown that right ventricular function can also be accurately measured by gated MSCT. The assessment of right ventricular function may have special interest in patients with acute pulmonary embolism. Right ventricular enlargement on chest CT has been shown to be a predictor of early death in patients with acute pulmonary embolism [20, 21]. Even the dimensions of the right ventricle in non-gated CT images may be predictive for mortality in this setting. The potential value of gated MSCT for assessing right ventricular function in patients with pulmonary embolism is now under investigation.

Assessment of pulmonary veins Atrial arrhythmias often originate in the pulmonary veins and can be treated with percutaneous radiofrequency catheter ablation. With this technique, the arrhythmic foci are electrically disconnected from the left atrium by means of catheters placed in the left atrium [22]. Pre-procedural MSCT examination is helpful to depict the anatomy of the pulmonary veins and left atrium and particularly to demonstrate additional pulmonary veins (e.g. middle lobe vein), which is important for planning the interventional procedure. Variations in pulmonary venous anatomy are quite common and comprise variation in the number of veins as well as the occurrence of common ostia

Figure 6. Ventricular function imaging at 64-row multidetector CT (MDCT). 26-year-old male patient after surgery for congenital heart disease. Ventricular function can be assessed after drawing the endocardial ventricular contours in (a) end-diastolic and (b) endsystolic phases at multiple cardiac levels, thereby including the ventricular volumes. 14


Cardiac MSCT

evaluation of patients with chest pain presenting to the emergency department was reported [26]. It was shown that MSCT is feasible to evaluate chest pain patients comprehensively. During one comprehensive MSCT protocol cardiac and non-cardiac causes of chest pain can accurately be diagnosed. It is expected that MSCT will become a gatekeeper in patients presenting with chest pain from various sources.

References

Figure 7. Pulmonary vein imaging at 64-row multidetector CT (MDCT). 64-row MDCT, non-ECG-synchronized imaging. 59year-old male patient. Pre-interventional assessment of pulmonary veins for radiofrequency ablation. Posterior view of the patient’s heart. Common ostium for the left pulmonary veins, i.e. the pulmonary veins join before entering the left atrium. Separate ostia for the right pulmonary veins. LS, left superior pulmonary vein; LI, left inferior pulmonary vein; RS, right superior pulmonary vein; RI, right inferior pulmonary vein; LA, left atrium; LPA, left pulmonary artery; RPA, right pulmonary artery; VC, inferior vena cava.

and early branching [23]. Three-dimensional surface rendering reconstructions provide a quick overview of the pulmonary venous anatomy, but cross-sectional reconstruction in coronal, sagittal and transverse orientations is necessary for full appreciation of the morphology of the pulmonary veins (Figure 7) [24]. Post-procedural MSCT also offers an opportunity for follow-up of the pulmonary vein after ablation [25]. MSCT pulmonary venography requires a contrast enhanced helical acquisition. To avoid motion artefacts a half reconstruction is generally performed, yielding a reconstruction window of about 165–200 ms. This is sufficiently short for imaging the rather large pulmonary veins with diameters well above 10 mm. Reliable images can be acquired without the use of ECG gating. Breathhold acquisitions with a high pitch factor and resulting rather low patient dose are routinely obtained. The potential additional value of ECG synchronized MSCT is under investigation.

Conclusion MSCT is a highly accurate tool for the non-invasive detection of coronary artery disease. Further technical advances are expected in acquisition techniques as well as post-processing of the CT data. Detector technology and arrays may be further expanded, allowing shorter imaging times. Improved temporal and spatial resolution will contribute to better stenosis quantification and plaque characterization. Integration of coronary artery imaging and functional data are feasible with current MSCT. Shorter scanning times may allow integration of coronary imaging, first-pass perfusion imaging as well as wall motion analysis from the same data set. Other cardiovascular applications also benefit from the improvements in CT technology. Recently, the value of MSCT for the The British Journal of Radiology, January 2006

1. Schuijf JD, Bax JJ, Salm LP, Jukema JW, Lamb HJ, van der Wall EE, et al. Noninvasive coronary imaging and assessment of left ventricular function using 16-slice computed tomography. Am J Cardiol 2005;95:571–4. 2. Paul JF, Wartski M, Caussin C, Sigal-Cinqualbre A, Lancelin B, Angel C, et al. Late defect on delayed contrast-enhanced multi-detector row CT scans in the prediction of SPECT infarct size after reperfused acute myocardial infarction: initial experience. Radiology 2005;236:485–9. 3. Dodge JT Jr, Brown BG, Bolson EL, Dodge HT. Lumen diameter of normal human coronary arteries. Influence of age, sex, anatomic variation, and left ventricular hypertrophy or dilation. Circulation 1992;86:232–46. 4. Shechter G, Resar JR, McVeigh ER. Rest period duration of the coronary arteries: implications for magnetic resonance coronary angiography. Med Phys 2005;32:255–62. 5. Hofman MB, Wickline SA, Lorenz CH. Quantification of in-plane motion of the coronary arteries during the cardiac cycle: implications for acquisition window duration for MR flow quantification. J Magn Reson Imaging 1998;8:568–76. 6. Lu B, Mao SS, Zhuang N, Bakhsheshi H, Yamamoto H, Takasu J, et al. Coronary artery motion during the cardiac cycle and optimal ECG triggering for coronary artery imaging. Invest Radiol 2001;36:250–6. 7. Harell GS, Guthaner DF, Breiman RS, Morehouse CC, Seppi EJ, Marshall WH Jr, et al. Stop-action cardiac computed tomography. Radiology 1977;123:515–7. 8. Girshman J, Wolff SD. Techniques for quantifying coronary artery calcification. Semin Ultrasound CT MR 2003;24:33–8. 9. Agatston AS, Janowitz WR, Hildner FJ, Zusmer NR, Viamonte M Jr, Detrano R. Quantification of coronary artery calcium using ultrafast computed tomography. J Am Coll Cardiol 1990;15:827–32. 10. Ulzheimer S, Kalender WA. Assessment of calcium scoring performance in cardiac computed tomography. Eur Radiol 2003;13:484–97. 11. Thompson GR, Partridge J. Coronary calcification score: the coronary-risk impact factor. Lancet 2004;363:557–9. 12. Leber AW, Knez A, von Ziegler F, Becker A, Nikolaou K, Paul S, et al. Quantification of obstructive and nonobstructive coronary lesions by 64-slice computed tomography: a comparative study with quantitative coronary angiography and intravascular ultrasound. J Am Coll Cardiol 2005;46:147–54. 13. Leschka S, Alkadhi H, Plass A, Desbiolles L, Grunenfelder J, Marincek B, et al. Accuracy of MSCT coronary angiography with 64-slice technology: first experience. Eur Heart J 2005;26:1482–7. 14. Achenbach S, Daniel WG. Computed tomography of the coronary arteries: more than meets the (angiographic) eye. J Am Coll Cardiol 2005;46:155–7. 15. Cury RC, Pomerantsev EV, Ferencik M. Comparison of the degree of coronary stenoses by multidetector computed tomography versus by quantitative coronary angiography. Am J Cardiol (In press).

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A de Roos, L J M Kroft, J J Bax et al 16. Dirksen MS, Bax JJ, de Roos A, Jukema JW, van der Geest RJ, Geleijns K, et al. Usefulness of dynamic multislice computed tomography of left ventricular function in unstable angina pectoris and comparison with echocardiography. Am J Cardiol 2002;90:1157–60. 17. Cerqueira MD, Weissman NJ, Dilsizian V, Jacobs AK, Kaul S, Laskey WK, et al. Standardized myocardial segmentation and nomenclature for tomographic imaging of the heart: a statement for healthcare professionals from the Cardiac Imaging Committee of the Council on Clinical Cardiology of the American Heart Association. Circulation 2002;105:539–42. 18. Schuijf JD, Bax JJ, Jukema JW, Lamb HJ, Vliegen HW, van der Wall EE, et al. Noninvasive evaluation of the coronary arteries with multislice computed tomography in hypertensive patients. Hypertension 2005;45:227–32. 19. Schuijf JD, Bax JJ, Jukema JW, Lamb HJ, Vliegen HW, Salm LP, et al. Noninvasive angiography and assessment of left ventricular function using multislice computed tomography in patients with type 2 diabetes. Diabetes Care 2004;27:2905–10. 20. Schoepf UJ, Kucher N, Kipfmueller F, Quiroz R, Costello P, Goldhaber SZ. Right ventricular enlargement on chest computed tomography: a predictor of early death in acute pulmonary embolism. Circulation 2004;110:3276–80. 21. van der Meer RW, Pattynama PM, van Strijen MJ, van den Berg-Huijsmans AA, Hartmann IJ, Putter H, et al. Right ventricular dysfunction and pulmonary obstruction index at helical CT: prediction of clinical outcome during 3-month follow-up in patients with acute pulmonary embolism. Radiology 2005;235:798–803.

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22. Pappone C, Rosanio S, Oreto G, Tocchi M, Gugliotta F, Vicedomini G, et al. Circumferential radiofrequency ablation of pulmonary vein ostia: A new anatomic approach for curing atrial fibrillation. Circulation 2000;102:2619–28. 23. Ghaye B, Szapiro D, Dacher JN, Rodriguez LM, Timmermans C, Devillers D, et al. Percutaneous ablation for atrial fibrillation: the role of cross-sectional imaging. Radiographics 2003;23 Spec No:S19–S33. 24. Jongbloed MR, Dirksen MS, Bax JJ, Geleijns K, Lamb HJ, Van der Wall EE, et al. Multislice computed tomography to evaluate pulmonary vein anatomy prior to radiofrequency catheter ablation of atrial fibrillation. Radiology 2005. 25. Maksimovic R, Cademartiri F, Scholten M, Jordaens LJ, Pattynama PM. Sixteen-row multislice computed tomography in the assessment of pulmonary veins prior to ablative treatment: validation vs conventional pulmonary venography and study of reproducibility. Eur Radiol 2004;14:369–74. 26. White CS, Kuo D, Kelemen M, Jain V, Musk A, Zaidi E, et al. Chest pain evaluation in the emergency department: can MDCT provide a comprehensive evaluation? AJR Am J Roentgenol 2005;185:533–40.



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Technology solutions for better outcomes: integrated information management in key to productivity increases in medicine H REQUARDT, PhD Group Executive Management, Siemens Medical Solutions, Henkestrasse 127, 91052 Erlangen, Germany

Abstract. The challenges to healthcare systems around the world are primarily impacted by two topics: demographic factors and progress in medicine. An ageing population inherently needs more medical services which add financial burdens, in particular, to public healthcare. Since the level of medical education is growing at the same time, we are observing an increased demand for sophisticated (in general expensive) medicine. Drastic changes in financing seem unavoidable. Multiple diagnoses, repeated examinations, trial-and-error, overcapacities and other signs of missing economical considerations are reinforced by reimbursement systems. In a world where, in principle, all information is available everywhere, more than a patient’s history should be accessible. Other industries have knowledge management systems in place that make state-of-the-art expertise available everywhere. Intelligent patient databases could consist of learning cycles that (i) enable the individual to benefit from structured knowledge, in addition to personal experience of the physician, and (ii) use the knowledge generated from the individual to extend the database. The novel area of molecular medicine fits perfectly well into these scenarios. Only attached to an IT backbone can the flood of information be managed in a beneficial way. Efficiency improvements in healthcare address the needs of all parties in the system: patients, providers, and payers. The opportunities, however, can only materialize if everyone is prepared to change. IT will set the standards for the biggest challenge in healthcare: The paradigm shift in medicine.

Introduction Demographic developments are placing tremendous pressure on healthcare systems around the world. Although age distribution varies significantly in different countries (e.g. China’s one-child policy versus India’s firtree distribution), problems come down to one common denominator: We are all living longer. Figure 1 [1, 2] shows the age distribution in more developed regions and the prognoses for 2025. It is obvious that health is a major macroeconomic factor. If we want to avoid the situation that fewer and fewer payers have to support more and more users of healthcare services, we will need to see more elderly people working. The prerequisite for this development is that they stay healthy. Healthcare systems thus would need to prove that the investment in them pays off as a productivity factor. A related challenge is reflected in the fact that a growing population is increasingly demanding to actively participate in medical progress. Mass media and the Internet depict what is possible today; with the majority of research being funded by the public purse. Thus, it is a natural desire that the same paying public also wants to enjoy the benefits that are generated. The basic question is: How can all of this remain affordable? Cutting cost by cutting services is not helpful for addressing both the need for higher quality care and the necessity to reduce cost. Instead, all contributors to the delivery of healthcare need to ask themselves ‘‘How can we do more with less?’’ If we draw an analogy with industry, Received 16 August 2005 and accepted 16 September 2005.


this question translates to ‘‘What levers do we see to improve efficiency?’’

Innovations drive efficiency Medical industries are developing not only more cost effective and reliable systems, but are also generating more and more relevant patient information in less examination time. Figure 2 shows a standard way of looking at CT datasets. The approximately 2 GB of raw image data that are typically acquired in a 5 s scan are stored in cache memories, are post-processed with volume renderers and can be displayed according to the interpreter’s comfort view. A different example is given in Figure 3: Not only has the amount of data dramatically increased, but so has the quality. In this case, a high-resolution three-dimensional (3D) image of the moving heart displays the stent structures with superb resolution. The broadening of the application scope is typical for each of the imaging modalities: Angio suites do excellent 3D imaging with cone beam reconstruction algorithms, linear accelerators deliver kV and MV images, magnetic resonance scanners have left the domain of pure morphologic imaging, and now measure functions in various ways. As an example, Figure 4 shows colour coded diffusion spectral imaging that is highly correlated with the directions of nerve bundles. The international medical industry has developed many technologies that can be utilized to improve efficiency in 17

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Figure 1. The change of age distribution in more developed regions. The qualitative cost curve reflects the current status. If nothing changes, the real overall cost can be the integral over the age distribution multiplied by the cost curve.

Figure 3. High resolution CT image of the heart. The stent structure is clearly delineated.

potential is not only a matter of technology; reimbursement systems and workflow structures have to be adjusted accordingly.

Figure 2. Volume-rendered abdominal CT image. The underlying image dataset consists of approximately 800 images.

diagnostic and therapeutic processes. Figure 5 shows in a schematic diagram how these developments can be locked into the learning cycles of healthcare providers. The potential for cost savings without sacrificing quality of care is clear. It is, however, evident that leveraging this 18

Is more always better? The basic question ‘‘Do I get enough information about my patient?’’ is no longer appropriate from a technology perspective. The medical industry has established time-tomarket cycles that can rapidly turn a novel clinical parameter into a product standard. Only 6 weeks elapsed between the identification of the SARS virus and the availability of a clinical test. The problem is no longer the The British Journal of Radiology, January 2006

Technology solutions for better outcomes

Figure 6. PET images overlaid to a volume-rendered CT dataset. The primary breast cancer is clearly delineated. Metastasis search is done within the same dataset. Figure 4. Diffusion spectral MR image. Colours code for spatial directions.

lack of data; the problem lies in filtering out the relevant information. There are various technological solutions for filtering. A widely practiced method uses overlay of images with different measurement parameters. Figure 6 shows an example in which a positron emission tomography (PET) image shows us where to focus in a set of hundreds of CT images. The overlaid images help us to select the slices of interest. A totally different approach with similar outcome is represented by ‘‘computer-aided detection’’ (CAD)

algorithms. Figure 7 shows CAD-detected polyps within a virtual colonoscopy dataset acquired with CT. These algorithms have now reached a performance level that is comparable with human readers. It is, however, still applicable only for simple structures, but can help us to focus our attention on the more complex features. Progression of CAD into more complex structures will be subject to the availability of standardized reference cases. It can be implied that innovation pressure for medical devices will in future not only focus on the generation of more data, but more and more on the generation of

Figure 5. Contiguous improvement cycles stimulated by technology (examples). The British Journal of Radiology, January 2006

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Figure 7. Computer-aided detection (CAD) algorithms detect polyps in a virtual colonoscopy. The sensitivity for polyps ¢ 6 mm is on average 90%; and the median false positive rate is a manageable 3 per volume [3].

‘‘smarter data’’. Yes, there will be CT scans that do 256 slices. But at the same time there will be an industry focus on systems with two or three X-ray detector systems that can generate not only increased temporal resolution, but also open up new degrees of freedom with respect to contrast by applying different anode voltages in the subsystems. Figure 8 shows a basic set-up for such a system.

Figure 8. Multitube CT set-up. The system enables a new degree of freedom allowing for double temporal resolution and/or novel contrast opportunities. 20

Overall, the focus of industry will move from ‘‘generation of data’’ towards ‘‘exploitation of data’’. It is evident that information technology is a key enabler for that shift.

IT enables process optimization In a patient-centric system, the ultimate outcome of the treatment is reflected by the status of the patient. The typical patient process in a hospital usually starts with diagnostic steps (radiology, ECG, lab, …), iterates with various therapeutic procedures (medication, surgery, radiation, …), and terminates with the recovery of the patient (ICU, ward, rehab, …). The most competitive healthcare provider will be the one that optimizes the entire procedure chain rather than the individual steps (this does not relieve the individual departments from delivering the best quality; ‘‘best’’ according to cost optimization criteria means ‘‘adequate and intelligent’’). In industrial analogy this means analysis, mapping and continuous improvement of workflow. Workflow optimization comprises the moving of patients, resources and information within the healthcare continuum according to certain rules. Everything (including the rules) is subject to best practice shared across all relevant healthcare participants throughout the world. Workflow can be referenced in ‘‘hospital information systems’’ by so-called workflow engines. An example of what a workflow engine can contribute is given in Figure 9: The emergency treatment of an acute stroke patient is managed by a computer network. The state-ofthe-art workflow engine would not only draft a work list, it would also monitor all activities in feedback loops. Cross-checks with rules engines ensure that the patient experiences state-of-the-art stroke treatment procedures. Figure 9 gives an impression how a workflow engine can be programmed according to the local conditions. It is obvious that workflow engines not only synchronize The British Journal of Radiology, January 2006


Figure 9. Workflow engine editor. The various decision steps reflect the time-critical diagnosis and treatment of an acute stroke. The time window for initiating thrombolysis is computer controlled.

Figure 10. Steps for cancer development. Today’s procedures detect cancer at a very late stage associated with high treatment cost and reduced prognosis. Early detection schemes lead to cellular and molecular levels; one of the exciting novel areas of ‘‘molecular medicine’’.


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Figure 11. Process chart of future treatment planning. Data access for both the patient’s individual electronic patient record and a comprehensive knowledge data base are crucial to enable state-of-the-art medical treatment for everyone, everywhere.

clinical activities, but also other day-to-day operations, e.g. discharge (paper work needs to be ready, transportation needs to be arranged, room needs to be made up, bed needs to be cleaned, etc.). Workflow engines will not only change the way care is delivered, but will also define the requirements for newly developed systems. Requirements and job descriptions in both arenas, industry and healthcare services, will be affected.

The patient is an individual The process chain within healthcare environments (prevention R diagnosis R therapy R care) is obviously not limited to hospitals. If we look at a schematic development of cancer in Figure 10, we realise that with today’s diagnostic methods we detect cancer only at a very late stage with higher cost and lower quality of life. Patient-focused healthcare systems will bring the intervention point forward to an earlier stage of the disease. With early detection and prevention capabilities, healthcare will increasingly be looked at just like every other service industry. The patient will behave like any other customer, but still with one fundamental difference: He/she is not free in selecting the disease. To shift the intervention point in an efficient way, much basic research remains to be done: The complexity of the ‘‘omics’’ (genomics, proteomics, metabolomics) needs to be understood and standardized with respect to the development of individual diseases. The potential, however, is big and every single day new cancer genes are being discovered or proteins identified that originate in specific tissue alterations. The diagnostic industry is asked not only to 22

deliver blood sample tests, but also software modules that make the associated knowledge available. The individualization, however, is not only subject to the diagnosis of the individual patient. It also needs to give clear recommendations for an optimized treatment. The entire arena of pharmacogenomics will be closely associated with ‘‘omics’’ analysis. Also, specific tumour metabolisms can be clearly understood and thus individually treated. It becomes evident that in scenarios like these, the diagnostic process moves from primary diagnostic to optimized treatment planning and follow up.

The holistic scenario The topics discussed so far lead to a few characteristics of future healthcare systems:

(1) they will be patient-focused and workflow-driven; (2) the patient’s entire history will be accessible through an electronic patient record (EPR); (3) the providers will be in a competitive situation and thus will publish proven outcome statistics to differentiate themselves; (4) the capability of sharing best practices with best-inclass providers will be a differentiating factor. The patient of the future will no longer rely just on the individual experience of his physician, but on the entire medical knowledge that is available. Obviously, the individual experience becomes part of that knowledge, but there are also other contributors. Figure 11 shows a scenario of how the individual patient information can be matched with the available knowledge. The individual The British Journal of Radiology, January 2006


treatment plan for the patient is mainly impacted by two elements: (1) the clinical knowledge database with rules for utilization of equipment and drugs, contraindications, standardizations, procedures and others; (2) the EPR consisting of images, lab data, structured reports, ‘‘omics’’, etc. Those databases will be mined by software agents for reference cases with proven outcome data to derive the most promising treatment plans. This enables the primary care physician (PCP) to match his individual experience with all the information that is available in the data stores. The databases will not only be filled with expert knowledge from medicine, but will also include related disciplines like pharmacology, radiation biology, biomechanics and others. In short, the PCP has a real, powerful tool that leaves him with a high degree of confidence that he has done all he can to help the patient. It will certainly be a long way to reach this scenario, but at the same time it is worth defining and working towards a common vision. Enabling technologies are there to help make this vision reality. Many new problems will come up including topics like data protection, ethics, business models or simply operational realization, and a social consensus will be required to address them all.


Medicine will never become deductive, but managing its complexity will become easier. Although basic work remains to be done, the technological solutions are available today. It is now a question of political desire to launch the paradigm shift in medicine.

References 1. Population Division of the Department of Economic and Social Affairs of the United Nations Secretariat. World Population Prospects: The 2004 Revision Population Database. [Online]. 2005 [cited 2005 March 15]. Available from: URL: http://esa.un.org/unpp/ 2. Economic Policy Committee (EPC). Budgetary challenges posed by ageing populations: the impact on public spending on pensions, health and long-term care for the elderly and possible indicators of the long-term sustainability of public finances. Brussels. 2001 October 24 (EPC/ECFIN/655/01-EN final). p. 34. 3. Bogoni L, Cathier P, Dundar M, Jerebko A, Lakare S, Liang J, et al. Computer-aided detection (CAD) for CT colonography: a tool to address a growing need. Br J Radiol 2005;78:57– 62.

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The case for particle therapy B JONES, MD, FRCR, MedFIPEM Queen Elizabeth University Hospital, Birmingham B15 2TH, UK

Abstract. Among the most important decisions facing the British Government regarding the treatment of cancer in the National Health Service (NHS) is the purchase of charged particle therapy (CPT) centres. CPT is different from conventional radiotherapy: the dose is deposited far more selectively in Bragg Peaks by either protons or ‘‘heavy’’ ions, such as carbon. In this way, it is possible to ‘‘dose paint’’ targets, voxel by voxel, with far less dose to surrounding tissues than with X-ray techniques. At present the UK possesses a 62 MeV cyclotron proton facility at Clatterbridge (Wirral), which provides therapy for intraocular cancers such as melanoma; for deeper situated cancers in the pelvis, chest etc., much higher energies, over 200 MeV are required from a synchrotron facility. There is an impressive expansion in particle beam therapy (PBT) centres worldwide, since they offer good prospects of improved quality of life with enhanced cancer cures in situations where conventional therapy is limited due to radioresistance or by the close proximity of critical normal tissues. There is a threat to UK Oncology, since it is anticipated that several thousand British patients may require referral abroad for therapy; this would severely disrupt their multidisciplinary management and require demanding logistical support.

The benefits of an increase in charged particle therapy (CPT) centres in the UK would be not only for children and young adults with cancer, where a reduced risk of radiation induced malignancy is predicted, but also in older patients where it is necessary to avoid abnormal tissues such as an enlarged heart/restricted lung irradiation and where artificial (metallic) joints may cause difficulties in the use of conventional radiotherapy techniques. The results of phase I and II clinical studies are extremely encouraging. The UK must obtain at least one CPT centre with protons/ions in order to conduct research and development; it is suggested that quality adjusted life years should be used to assess outcomes. It is anticipated that the UK might eventually require 7–8 such centres in 10–15 years from now. In the meantime, healthcare purchasers and providers need to put in place mechanisms and personnel for patient referrals abroad, as well as the establishment of UK CPT facilities.

NHS became disadvantaged in terms of expensive technological acquisition. Dr R D Errington related the history of cyclotron radiotherapy at the BIR President’s Day conference in 2003. He detailed how the initial promising results obtained with neutron therapy at The Hammersmith Hospital were not subsequently confirmed in randomized trials at Edinburgh and at the Clatterbridge facility [1, 2], which produced neutrons that matched a 5 MeV X-ray beam. The latter facility was converted to produce protons on the recommendation of the late Prof. Arthur Jones of St Bartholomew’s Hospital. This enabled patients with choroidal melanoma of the eye to receive radical radiotherapy using protons; this technique was the first example of three-dimensional (3D) radiotherapy in the UK. Over 1400 patients have by now received this therapy with a local control rate of 98% – an outstanding achievement within British medicine [3].

Background

Past attempts to obtain a higher energy facility in the UK

The connection between subatomic particles and health delivery improvements may seem rather tenuous, but the narrative begins in 1879, when J J Thompson discovered the negatively charged electron in Cambridge, and Aneurin Bevan was born in Wales. The subsequent discoveries of the positively charged proton (a term coined by Ernest Rutherford in 1920) and the uncharged neutron by James Chadwick in 1931, confirmed the preeminence of our science. Bevan, with similar precision of thought, digested the wide recommendations of the Beveridge Report (1942) and transformed most of its principles to practical achievements, including the National Health Service Act of Parliament (1946) and the inception of the service in 1948. Subsequently, Britain was at the forefront of practical applications of physics and engineering developments in cancer therapy until the early 1990s, when the reorganized 24

Since 1992, Clatterbridge, Oxford and the National Physical Laboratory at Daresbury (near Warrington) have all unsuccessfully attempted to obtain a higher energy CPT facility [4]. All these bids were rejected because of perceived lack of clinical support, intermittent beam availability, the lack of clinical trial evidence, the recommendation that a facility should be sited in a University Hospital campus and perhaps mostly, the expected high initial costs incurred at a time when NHS reforms discouraged large-scale projects, even the provision of new (replacement) linear accelerators. More recently, there has emerged a more collective response from clinical oncologists and medical physicists who appreciate that obtaining a CPT facility is essential The British Journal of Radiology, January 2006

The case for particle therapy

for the advancement of radiation oncology standards in the UK. The Royal College of Radiologists (RCR), British Institute of Radiology (BIR) and Institute of Physics and Engineering in Medicine (IPEM) for example all support the case for a CPT facility. Recent improvements in the quality of cancer imaging and the availability of industrially produced turnkey facilities, has allowed the question to be carefully re-considered and better understood, particularly in relation to the rapid expansion in CPT facilities abroad. Figure 2. Approximate depth dose positions of partially spread

Technical aspects The velocity of heavy charged particles (electrons are considered to be light) is reduced as they traverse deeper through tissues. The interaction probability to cause ionization increases as the velocity falls, so that a peak of dose occurs at a depth proportional to the energy imparted to each particle. William Bragg, a British physicist, described this phenomenon over 100 years ago [5]. The so called Bragg peak can be ‘‘spread out’’ to achieve a plateau of uniform dose that covers a target by use of rotating range-shifting modulators of variable thickness. In the past, passively scattered beams were used in this way to provide wide circular or rectangular beams with spread out Bragg peaks (Figure 1). More recently, the spot scanning method allows smaller beams to deposit their peaks within individual voxel targets defined by good imaging techniques: by the use of ‘‘wobbler’’ magnets and particle energy selection, the raster scanning system allows cancer bearing voxels (defined by x, y, z, co-ordinates), to be ‘‘dose painted’’. The Bragg peak position will depend on the initial energy imparted to the particles as well as their mass and charge; the Bethe-Bloch equation contains all the necessary parameters. It can be seen from Figure 2 that the range for clinical use should be at least 200 MeV in the case of protons; higher energies – up to 400 MeV – for carbon ions.

out Bragg peaks for protons of different energies.

Gantries and robots Within treatment rooms there are options for beam arrangements. The simplest approach is to have either fixed horizontal or vertical beams, or a combination of the two for the simplest treatments. An isocentric rotating gantry is required for more complex geometrical problems. These consist of large cylindrical rotating structures that contain the beam bending magnets: they weigh 100 tonnes for protons and 200 tonnes for ions and require movement with 1 mm precision of beam placement. Future engineering innovations may reduce the tonnage and costs. Robotic treatment couches are desirable in order to rapidly position the patient at predetermined angles relative to the beams; they may also transport patients in fixed positions from image guided or other localization devices in the treatment rooms to the actual treatment location. Radiographers may feel sensitive about robotics, but it will always be the radiographer who commands the robot and remotely monitors their performance.

Typical centre The typical layout of a centre is illustrated in Figure 3. The particles are injected from a small linear accelerator and further accelerated to higher energies around the synchrotron, then extracted and delivered selectively to different rooms; the beam switching time between rooms is

Figure 1. Schematic depth dose diagram of a proton beam Bragg peak, the spread out Bragg peak and a megavoltage X-ray beam (modified from Suit et al [12]). The grey shaded areas indicate the extent of dose reduction within normal tissues situated proximal and distal to the tumour target.


Figure 3. A schematic diagram of a synchrotron treatment centre.

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as short as 10–20 s. A high throughput of patients can be achieved by efficient placement and preparation of patient position in advance of the beam availability in each room. Larger synchrotrons can deliver carbon ions or protons. Some rooms may be equipped with positron emission tomography (PET) scanning facilities and other image guided devices. The overall arrangement is quite different from standard radiotherapy departments where there is a linear accelerator in each treatment room. For more detailed plans see various chapters in Supplement 2 of Radiotherapy & Oncology (volume 73), 2004 [10].

The dose distribution advantages Many authors have made important contributions by means of comparative dose distributions using X-rays and protons, which are summarized elsewhere [6, 7]. The essential principles may be better realised by inspection of relatively simple depth dose diagrams as seen in Figure 4. In Figure 4A, the spread out Bragg peak (SOBP) is seen from a single beam entering from the left hand side. In contrast, the X-ray fall off of dose is pseudo-exponential as shown in Figure 4D. When two opposed fields are used there is approximately uniform dosage in the case of X-rays (as in Figure 4E), whereas for particles there is a preferential dose deposition where the SOBPs coincide, as in Figure 4B. For three intersecting beams, there is now some degree of selectivity for X-rays as seen in Figure 4F, but the ratios of dose in the centre to that near the surface is considerably better for the particles as shown in Figure 4C. Inspection of axial views of three intersecting beams, as in Figure 5, shows the different dose distributions achievable. These figures can be normalized to give the same dose in the central region, with resulting lower peripheral doses for particles. The absence of dose in one direction beyond the target is striking – this arrangement may be used to reduce exposure to critical structures such as

Figure 4. (A–C) Simplified schematic diagrams of protons and (D–F) X-ray percentage depth dose distributions for three simple field arrangements. In B, C, E, F depth is measured along the direction of opposing fields. Relatively small changes in dose are not included in these fields. 26

Figure 5. (a,b). Axial views of simplified schematic dose distributions for three field coplanar techniques using X-rays and protons.

rectum, spinal cord, etc. Rotation of the beams may also be used to avoid beam traversion through, or scattered radiation from metal prostheses, which cause dose uncertainties in treatment planning. The reduction in the so called integral dose, which is an assessment of dose to wider volumes within a patient, is considerable – proton beams generally reduce this by 50% and frequently by more in some cases [7]. This effect alone should reduce the risk of second cancer formation [8], which may be enhanced with the use of some modern linear accelerator based techniques such as intensitymodulated radiotherapy (IMRT), where there is a ‘‘dose bath’’ effect due to increased integral dose. Not only is the risk of second cancers reduced, but also substantial reductions occur in dose commitment to organs that are sensitive to radiation, e.g. kidneys, eyes, lung, heart, and parts of the nervous system. Low doses to substantial proportions of these organs can cause functional problems. For example, consider the treatment plans shown in Figure 6, where multiple field IMRT is compared with single field spot scanning protons. Whilst the target volume is covered equally well with both techniques, the dose bath effect is readily seen for IMRT, with significant dose to spinal cord and kidneys; the proton plan effectively spares these critical organs. Even a tissue such as bone is highly relevant: bone marrow cell production is not supported at doses above 30 Gy and longer term effects include osteoporosis, micro-fractures and fractures; in practice, low backache is not infrequent following pelvic radiotherapy, and bone density changes, revealed by MRI, are seen to exactly correspond to the beam portals used. For a wide variety of cancers the advantages of the improved dose distributions should provide substantial improvements in the quality of life where normal tissue doses are reduced and improved cure potential when tumour dose is increased. These are considered in further detail in Table 1, although the generic reduction of second malignancy is not included. Meticulous studies in Japan, using carbon ions, with respiratory movement gating compensation, have shown two extremely important results. They are:

(1) Cure of small peripheral screen detected lung cancers in a single exposure and without loss of lung function; similar The British Journal of Radiology, January 2006


(a)

(b)

Figure 6. (a, b) Comparative dose distributions for IMRT and protons for a recurrent sarcoma in a young 12-year-old boy (reproduced by kind permission of Dr A Lomax, PSI, Switzerland and Prof. P Hoskins, Editor of Clinical Oncology).

Table 1. The advantages of charged particle therapy (CPT) in a range of anatomical situations Cancer bearing region

Advantage of CPT

Breast Head and neck Pelvis (e.g. prostate, bladder, rectum)

Avoid irradiation of heart, lung and brachial plexus Reduced dose to spinal cord, salivary glands, eyes, bone and brain Reduced irradiation of bone, sparing of organs such as bladder, rectum; large sarcomas are safely treated without sacral plexus damage As in pelvis, but also improved dose to lateral parametrium, better distribution for vulvar cancers; can be used where brachytherapy not feasible; field extension to para-aortic region with less toxicity Reduced lymphoedema and deformities Better preservation of lung and heart function Marked reduction in acute effects, can safely dose escalate for radio-resistant cancers, e.g. hepatoma, cholangiocarcinoma Sparing of small bowel, spine and kidneys Reduction of irradiation to sensitive structures such as hypothalamus, pituitary, reduced risk of stroke Reduction of collateral irradiation to tissues outside the CNS, e.g. all tissues anterior to spine and reduced irradiation of appendages e.g. external auditory apparatus and eye, etc.

Gynaecological system

Limbs Lung Liver/pancreas Paraspinal/para-aortic CNS

cure rates can be achieved by surgery, but with inevitable loss of lung function [9]. (2) Cure of patients with primary liver cancers treated in four exposures; again similar rates of cure can be achieved following surgery but with considerable morbidity and some mortality [10]. These results suggest that radiotherapy might eventually replace radical surgery in deeply situated anatomical locations. The risks and costs of radical surgery are likely to increase with time in an ageing population. In addition, these results confirm previous theoretical predictions based on radiobiological modelling that as dose is better localized to the target and markedly reduced in a wider range of surrounding tissues, the principles of fractionation become less important [11]. Thus treatment can be delivered in far fewer exposures; the economics of CPT then become more favourable. In addition, the treatment is more elegant, involves fewer beams and is potentially less liable to errors made in treatment delivery. Owing to space constraints it is only possible to show a limited number of treatment plans. Figure 7 shows the advantages of a four field proton plan which could be used The British Journal of Radiology, January 2006

to treat a hepatoma or cholangiocarcinoma. The colour wash dose distribution shows how restricted the dose is to target; this spares the patient the acute side effects of nausea, vomiting and severe malaise which occur with X-ray traversion of the stomach, duodenum and liver.

Figure 7. Comparisons of dose distributions for a 4 field X-ray (photon) plan and a proton plan for treatment of hepatocellular cancer (courtesy of Dr J Munzenrider, Northwest Proton Therapy Centre, Boston, USA). 27

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Figure 8. An example of a single field application of protons to treat a posterior orbital cancer (courtesy of Dr J Munzenrider, Northwest Proton Therapy Centre, Boston, USA). The colours denote different dose levels with red being the full prescribed dose, with fall off to the limits of the beam.

The next example (Figure 8) shows how the brain and other bony structures in the head and neck can be spared due to the sheer elegance of a single field proton approach to treat cancers in the posterior orbit, such as lachrymal gland cancer or rhabdomyosarcoma. To obtain equivalent uniformity of dose across the target region, at least 2 or 3 X-ray fields would be required, with resultant exit doses into the brain.

The existing evidence base The clinical evidence base consists of phase I/II dose escalation studies. There are no randomized control trials that compare CPT with conventional radiotherapies [6], although there are randomized phase II ‘‘dose searching’’ studies. One example is the randomization between 72 Gy and 78 Gy cobalt Gray equivalent (CGE) for skull base chordomas at Massachusetts General. Some international authorities consider that randomized studies that compare conventional X-ray therapy with protons are not justified because of the advantageous dose distributions for the latter [12]. Whereas this may be true for skull base tumours and in hepatic cancers, there must be greater justification elsewhere, e.g. the comparison of IMRT/ implants with protons in prostate cancer. Whether phase III studies (comparisons with conventional radiotherapy) will be performed remains to be seen: some authorities consider that such research would be unethical [12]. It is inevitable that randomized comparisons of CPT against radical surgery will have to be done for small screen detected cancers in deeply situated tissues (see below).

Misconceptions It is not surprising that misconceptions abound when referring to CPT. Comparisons are often made with neutrons due to their production from similar 28

sophisticated equipment. It must be remembered that neutrons are neutral particles and consequently do not have Bragg peak characteristics: the additional toxicity seen with neutron therapy was due to the higher relative biological effect (RBE) and high integral doses. Precision is another issue: are protons and ions too precise? Certainly, the dose can be painted onto any safe volume, so that tumour margins can be fully respected. There is no reason why, in certain tumours, one cannot do wide initial volumes, shrinking down to smaller targets with increasing dose; protons could be used with three definite dose volume regions, e.g. 55 Gy, 65 Gy and 75 Gy volumes defined around a target simultaneously. Many oncologists assume that the advantages are only seen in tumours such as skull base chordomas. It must be realised that such tumours were treated because of poor results with conventional therapy and with limited proton beam time coupled with relatively low energy beams that precluded treatment of deeper structures. Greater beam availability has allowed testing of CPT in a wider variety of tumours in different locations.

Added value for science research and teaching A clinical facility could also be used for radionuclide production: the particles can activate stable elements to become radioactive, with applications in healthcare and industry. Overnight production allows income generating use of short-lived radionuclide on the following day. Synchrotron radiation, essentially mono-energetic bremstrahlung emitted when the particles are deviated by magnets, can be used for X-ray crystallography studies. Particle micro-beam analysis of solid state and biological material can also be pursued, e.g. intracellular diagnostic capacity at nanometre levels, testing of materials for their resistance to cosmic rays prior to space flights. A detailed case is presently being written by the Engineering and Physical Sciences Research Council (EPSRC) Medical Applications of Ion Beams Network.

Contributions from molecular biology The vast expansion in knowledge gained by research in molecular biology applied to oncology will inevitably result in more reliable early diagnosis of cancer. Screening of a population by ‘‘PCR (polymerase chain reaction) amplification’’ techniques and proteomic techniques should detect aberrant DNA and protein products from quite small cancers in body fluids. Further gene specific or target protein imaging using sophisticated forms of PET scanning may be sufficient to confirm the presence of small cancers in deeply situated organs. Image guided biopsies may also be necessary in some cases. These approaches are probably more practical than the more distant Holy Grail of cancer cure following the application of such approaches. This is not to say that such approaches will not be useful, particularly in modifying cancer growth patterns and metastatic potential; but when used alone, molecular approaches may be doomed to failure because of the capacity of a cancer to produce further mutations and to bypass metabolic blockade even when multiple approaches are used. However, the reliable earlier diagnosis of cancer would create a high demand for The British Journal of Radiology, January 2006


surgery and radiotherapy, particularly highly focal forms of radiotherapy that enable a high localized dose to be delivered with good sparing of normal tissues, as in CPT. The decisive clinical trials of the future may be those that compare CPT with surgery, particularly in sites where the latter has a high morbidity, mortality and cost, e.g. hepatic, pancreatic and renal surgery.

Contributions from medical oncology The reduction of exit dose radiation to skeletal regions that contain active bone marrow will reduce the risk of severe neutropenia and the morbidity and mortality that follow septicaemia. Thus CPT radiotherapy may be combined with more aggressive chemotherapy regimens. In addition, the risk of subsequent organ failure on exposure to certain classes of radiotherapy may be reduced. For example, the cochlear sparing associated with medulloblastoma proton-therapy is likely to reduce the high tone deafness associated with the use of Cisplatinum treatment [13]; the risk of renal failure may be reduced when using protons instead of IMRT to treat the para-aortic nodes in metastatic or advanced local cervix cancers. Also, the risk of severe cardiomyopathy may be reduced – even in the case of later exposure to anthracycline drugs – if the heart has not been exposed to significant radiation dose by use of CPT, e.g. in the case of left sided breast cancer. There is clearly a wide prospectus for research with a major input from medical oncologists with an interest in radiotherapy in this important area of oncology.

Contributions from surgery The increasing future role of radiotherapy in small volume deep-seated cancers has already been mentioned. For larger cancers, volume reduction using surgery may still be desirable, as might the concept of ‘‘improving treatment geometry’’ by selective resection and restoring a finite space between tumour and critical normal tissues. Prolonged surgery will always reduce tissue tolerance owing to accumulated vascular damage. Decisions regarding operability, the extent of surgery and the necessary dose of radiation will always need careful consideration according to circumstances. The possibility of preoperative CPT in some situations would be useful: in Massachusetts General Hospital there is already some experience of pre-operative proton therapy to paraspinal bone tumours in order to reduce the potential for brachytherapy catheter implantation of tumour cells when radio-iodine seed implants are made into the adjacent bone situated distally to the tumour. There is clearly considerable scope for research in the degree to which surgery and CPT can be combined.

Research and development: quality adjusted survival end points There is increasing disquiet that very large trials are required to detect small incremental changes in outcomes, with a tendency to favour patient survival as the primary end point, possibly with inclusion of some separate quality of life study. This stance is not unreasonable for The British Journal of Radiology, January 2006

comparisons of chemotherapy schedules, where severe acute toxicity is life threatening and influences survival. Such approaches are far from ideal for the assessment of new radiation techniques where subtle long-term differences in a wide spectrum of tissues are more relevant. Newer forms of trial assessment will probably be necessary. One such approach is considered here. In a computer generated survival curve with only 100 patients in each treatment arm, with a survival advantage of ,10% for CPT c.f. X-rays, the p-value exceeds 0.05 using the logrank test (p.0.05). The side effect profiles (graded in four categories according to ascending severity) show subtle improvements with CPT, although when tested using a contingency table the Chi-squared statistic shows a non significant trend (p.0.05) because of the low numbers in each category. But when survival is adjusted by using the toxicity grade factor F defined as (5-x)/5, where x is the toxicity grade with five categories, the quality adjusted survival (F times the actual survival) becomes highly significant (p,0.0001). More work is required to justify and encourage these approaches, but the potential advantages in terms of cost and rapidity of obtaining results with a greater number of trial arms containing different doses/treatment combinations are readily apparent from the example given. Such a novel approach could be used within CPT studies.

The threat to British oncology If the UK will not invest sufficiently rapidly in CPT facilities, there is a real risk of there being between 5000 and 12 000 patients who will require or demand therapy abroad in around 10 years from now [14]. These estimates were arrived at using the logistic equation to simulate supply and demand with best and worst case scenarios for overall capacity to accept UK referrals abroad. Treatment abroad would undoubtedly cause severe disruption of multidisciplinary cancer care as well as anticipated social and linguistic problems. In terms of staff retention, there is a risk that many British physicists, radiographers and oncologists might be attracted to work abroad. Also, the UK clinical trial portfolio may not contain state of the art radiotherapy and consequently our trials may become irrelevant and ignored elsewhere in the world.

Costs It has become politically incorrect to mention costs in medical circles, although cost effectiveness is deemed respectable and quotable. Such restrictive criteria are, for example, accepted by The British Medical Journal for its publications. One cannot escape the fact that the costs for synchrotron commissioning are large, of the order of £70–100 million depending upon the specifications for protons and the more expensive ions and how many large gantries are required. Some consideration has already been given to cost benefit and patient demand in Switzerland, Sweden, France and Austria [15–18]. Cost benefit will be most accurately measured prospectively within clinical trials. The costs charged will vary with the number of exposures: presently around £12 000 for 4 exposures at Clatterbridge; but with some economies of scale and improved throughput one can envisage CPT for around 29

B Jones

£8000–25 000 per patient, depending on the fractionation used; this is less than the cost of renal dialysis necessary to keep a patient alive for 1 year and compares favourably with the cost of prolonged radical surgery. A single UK centre should recoup its own initial and running costs within 6 years providing it can treat 2500 patients by its third year of operation. However, the UK would depend on a multitude of healthcare purchasing agreements – a most unsatisfactory system for the provision of complex healthcare. Definitive cancer treatment using radiation should be separated from these cumbersome procedures, with a clear assurance that all British patients with a diagnosis of cancer will receive equal access to more complex therapy where necessary. Dr Neil Burnet has estimated from Swedish data (Burnet N, personal communication) that the proportion of total cancer care costs spent on radiotherapy would increase from the present 5% to 6% if 15% of all radiotherapy is given by protons [18]. This is likely to be cost effective in the long term because of the reduced side effects and compares well with the present expenditure on cytotoxic chemotherapy, which accounts for around 12% of total cancer care. It remains unclear as to how funding can be achieved without a high level political decision. Even the new Foundation NHS Trusts cannot borrow the necessary monies to enable CPT. Our NHS needs better structures that can arrange finance, whether public or private: perhaps a return to regional and supraregional systems for cancer care?

Logistics for a National Centre The NHS has developed impressive Cancer Networks as part of its Cancer Plan, and CPT will need to be imaginatively superimposed on this framework. These existing networks are essential to ensure equity of access for CPT. Each local Network should form the basis of referral to special multidisciplinary team (MDT) meetings concerned with CPT. When a clinical indication is identified, then appropriate dose planning assessments are necessary: this might be achieved by electronic transfer of data to a national reference centre which itself might be virtual, i.e. it can be envisaged that all cases of tumour type X might be independently assessed in City A, and for tumour type Y in City B as for the physical appropriateness of IMRT or CPT. The referring city could also plan with the two modalities and confer with the national CPT centre. Encouragement for physicists and oncologists to attend a National Centre on a rotational or frequent basis, e.g. for specific MDT and treatment planning meetings, should also be encouraged. A national service will need to have strong links with other centres abroad for the treatment of rare conditions.

Logistics for referral abroad The prospect of referring hundreds or thousands of patients abroad is daunting. The time taken to assess and counsel, and to send all diagnostic information away is significant. There is an immediate need for full time staff devoted to these logistics, with attention to transfer funding for provision of appropriate care abroad. British 30

staff should be put in place to support patients and families whilst abroad and also to promote training in how to deliver CPT. Eventually, the number of treatment facilities in the UK should become appropriate to meet the needs of the British people. However, UK healthcare planners should urgently apply themselves to these problems and produce appropriate plans that meet the most likely short and long-term requirements.

Politics/Government/Research Councils and Charities CPT needs to be fully researched, with major UK participation. At least one high-energy UK CPT facility should be established to conduct clinical research and trials, with equitable patient referral via the Cancer Networks. The immediate questions for the UK authorities are ‘‘when’’ and ‘‘how many facilities’’ do we need? These important decisions confront the UK Government for future cancer care, and must be judged in the context of the proposed increased investment in the scientific base of this country [19]. The concept of joined up working across the various Research Councils (EPSRC, MRC, Accelerator Science, N-Tech), and linked to the major cancer charities (Cancer Research UK) should allow the UK to further develop the technology that underpins the most sophisticated form of radiation therapy against cancer. It would be tragic to wait until public awareness forces the issue. Bevan, an astute politician and cancer sufferer, would surely have sensed that the NHS should possess the weapon of particle radiotherapy within its arsenal against cancer, in the same way as he bravely supported an independent nuclear deterrent. He wanted only the best for the British people and so should we.

Acknowledgments The author is indebted to the following for discussions and their encouragement. Oncologists & Physicians: Prof. Pat Price, Dr Neil Burnet, Dr Trevor Roberts, Dr R D Errington, Dr P R Blake, Dr F Saran, Dr D Dearnaley, Dr D A L Morgan, Dr R Taylor, Dr A Cassoni, Dr M Gaze, Dr K I Hopkins, Dr D J Cole, Dr P N Plowman, Dr R Beaney, Dr R Rampling, Dr D Spooner, Dr A Crellin, Dr N G Glaser, Dr R H Phillips, Dr D V Ash, Prof. A Price, Prof. W Duncan, Prof. A Munro, Dr J Staffurth, Prof. C Coombes, Dr R K Coker, Prof. W Littler. RCR: Dr A Barrett, Dr R Hunter, Dr M V Williams, Prof. P Hoskins, Dr F Calman. Surgeons: Mr I McIndoe, Mr R E Kingston, Prof. G Cruickshank. Medical Physicists: Dr A Kacparek, Dr Roger Dale, Dr Ivan Rosenberg, Dr Stuart Green, Prof. A Beddoe, Dr S Blake, Dr D Thwaites, Dr A Nahum, Dr A Carabe. Academic Physicists: Dr K Kirkby, Dr D Parker, Prof. J Nelson. National Physics Laboratory: Dr H Palmens, Dr D Rayner. Declarations: BJ is a Trustee of The Cyclotron Trust (UK Charity) and a member of the EPSRC Medical Applications of Ion Beams Network.

References 1. Errington RD, Ashby D, Gore SM, et al. High energy neutron treatment for pelvic cancers: study stopped because of increased mortality. Br Med J 1991;302:1045–51.


The case for particle therapy 2. Maor MH, Errington RD, Caplan RJ, et al. Fast neutron therapy in advanced head & neck cancer: a collaborative internal randomised trial. Int J Radiat Oncol Biol Phys 1995;32:99–604. 3. Damato B, Lecuona K. Conservation of eyes with choroidal melanoma by a multimodality approach to treatment: an audit of 1632 patients. Opthalmology 2004;111:977–83. 4. Price P, Errington RD, Jones B. The clinical and scientific case for a high energy proton therapy facility in the UK. Clin Oncol 2003;15:S1–S9. 5. Brown A, Suit H. The centenary of the discovery of the Bragg Peak. Radiother Oncol 2004;73:265–8. 6. Jones B, Rosenberg I. Particle therapy Cooperative Oncology Group (PTCOG40), Institut Curie 2004. Br J Radiol 2005;78:99–102. 7. Jones B, Burnet NG. Radiotherapy for the future. Br Med J 2005;330:979–80. 8. Miralbell R, Lomax A, Cella L, Schneider U. Potential reduction of the incidence of radiation-induced second cancers by using proton beams in the treatment of pediatric tumors. Int J Radiat Oncol Biol Phys 2002;54:824–9. 9. Miyamoto T, Yanamoto N, Nishimura H, Koto M, Tsujii H, et al. Carbon ion radiotherapy for stage 1 non-small cell lung cancer. Radiother Oncol 2003;66:127–40. 10. Tsuji H. Overview of clinical experiences with carbon ions at NIRS. Radiother Oncol 2004;73(Suppl. 2):S41–9.


11. Jones B, Dale RG. Radiobiologically based assessments of the net costs of fractionated focal radiotherapy. Int J Radiat Oncol Biol Phys 1998;41:1139–48. 12. Suit H, Goldberg S, Niemerko A, Trofimov A, Adams J, et al. Proton beams to replace photon beams in radical dose treatments. Acta Oncologica 2003;42:800–8. 13. Plowman PN, Usher C. Authors’ reply. Br J Radiol 2005;78:285–6. 14. Jones B, Price P, Burnet NG, Roberts JT. Modelling the expected increase in demand for particle radiotherapy: implications for the UK. Br J Radiol 2005;78:832–5. 15. Goitein M, Jermann M. The relative costs of proton and X-ray radiation therapy. Clin Oncol 2003;15:S37–50. 16. Lundkvist J, Ekman M, Ericsson SR, Jonsson B, Glimelius B. Cost-effectiveness of proton radiation in the treatment of childhood medulloblastoma. Cancer 2005;103:793–801. 17. Mayer R, Mock U, Jager R, Potter R, et al. Epidemiological aspects of hadron therapy. Radiother Oncol 2004;73(Suppl. 2):S24–8. 18. Ringborg U, Bergqvist D, Brorsson B, Cavallin-Stahl E, Ceberg J, Einhorn N, et al. The Swedish Council on Technology Assessment in health care (SBU) systematic overview of radiotherapy for cancer including a prospective survey of radiotherapy practice in Sweden 2001 – summary and conclusions. Acta Oncologica 2003;42:357–65. 19. Jones B. Science and Innovation Summit 2004. Br J Radiol 2005;78:17–9.

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E



The contribution of PET/CT to improved patient management P J ELL, FMedSci, FRCP, FRCR Institute of Nuclear Medicine, UCL, London, UK

Abstract. With the introduction of both SPET/CT and PET/CT, multimodality imaging has truly entered routine clinical practice. Multiple slice spiral CT scanners have been incorporated with multiple detector gamma cameras or PET systems, such that the benefit of these modalities can be achieved in one patient sitting. The subject of this manuscript is PET/CT and its impact on patient management. Applications of PET/CT span the whole field of medical and surgical oncology since very few cancers do not take up the labelled glucose tracer, 18 F-FDG. Given the contrast achieved, high-quality data can be obtained with FDG PET/CT. This technology has now spread worldwide and has been the subject of intense interest, as witnessed by the vast body of published evidence. In this short overview, only a brief discussion of the main clinical applications is possible. Novel applications of PET/CT outside the field of oncology are expected in the near future.

Introduction The technologies of positron emission tomography (PET) and spiral computed tomography (CT) have been combined in a single multimodality detection instrument. The PET/CT scanner provides, in a single patient sitting, both the data to be expected from a high-end advanced spiral CT scanner and information recorded by a top of the range PET scanner, capable of depicting the distribution of positron-labelled tracers such as fluorodeoxyglucose (FDG). Routine image fusion is obtained, CT data being merged with PET data to aid in the exact localization of the site of FDG uptake. CT information is also used for the purpose of attenuation correction, which is now almost instantaneous; as a consequence, whole-body PET/CT studies can be obtained in less than 30 min. This has led to an increase in patient acceptance and throughput (30% over that achieved with PET alone). Scanning times are expected to improve further in the near future. With PET/CT studies obtained from a flat bed, this information can be used to improve radiotherapy planning, a novel and rapidly evolving application of this technology. PET/CT leads to improved lesion detection and localization and a faster learning curve for all involved; it has achieved significant acceptance at multidisciplinary case conferences [1, 2].

Applications Tables 1 and 2 summarize the present and predicted areas of application of PET/CT, and anticipated changes Table 1. Present and predicted areas of applications of PET/CT Present

Future

Oncology 97% Infection 2% Cardiology 1%

Oncology 70% Infection + Musculoskeletal 5% Cardiology 15% Psychiatry 10%

in tracer use. There are realistic expectations that a number of novel tracers, labelled with, for example, 18F or even 68Ga (to mention just two radionuclides), will lead to useful clinical studies on atherosclerosis [3], angiogenesis, hypoxia and detection of amyloid plaque in Alzheimer’s disease. Other tracers such as 18F-labelled thymidine (FLT: a marker of TK1 activity and indirectly of cellular proliferation) and 18F-labelled dopamine have already been applied in the fields of oncology (FLT and dopamine) [34, 35] and movement disorders (dopamine). The discussion below will, however, be restricted to the use of FDG in oncology. Labelled FDG provides some of the highest signal-tonoise ratios to have been observed in nuclear medicine. This is the result of a number of factors which play a role in the cellular uptake of FDG: over-expression of membrane GLUT transporters, increased glucose transport in malignancy, increased glycolysis, and increased hexokinase activity coupled with a decrease in glucose-6phosphatase activity. It is also now well known that maximal FDG uptake in the lesion is not reached within the first hour of intravenous administration. Invariably, a further increase in the signal-to-noise ratio can be observed at 2 h, and a plateau is reached much later. It must also be stressed that FDG is not a cancer-specific ligand: macrophages actively take up FDG [4, 5], and granulomas and inflammatory lesions can be falsely interpreted as malignant. From a practical point of view, the unit most often used to quantitate FDG uptake is the standardized uptake value (SUV). This normalizes the FDG taken up in a region of interest to the total amount of tracer injected and the patient’s body weight. The SUV is time dependent, since FDG continues to accumulate during the period of imaging. For each study, SUVs have to be measured at the Table 2. Anticipated changes in tracer use Present

Future

FDG 95%

FDG 85% 68 Ga + Others : 15%

Received 31 May 2005 and accepted 6 September 2005.

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PET/CT to improve patient management

same time after administration of FDG. From a region of interest, average or maximum SUVs can be obtained, the maximum SUV being the most reproducible value for comparative purposes. SUVs greater than three are most often associated with malignancy. Whilst this cut-off is somewhat arbitrary, it is of value since it helps to distinguish malignant from benign nodal disease: enlarged nodes on CT with low SUVs are almost always benign.

PET/CT and FDG in oncology As one might expect, the main areas of interest are in diagnosis, staging, treatment monitoring and radiotherapy planning [6].

Diagnosis PET/CT is infrequently used to offer or aid in the diagnosis of a patient’s primary condition, the principal indications for this purpose being suspected paraneoplastic syndrome, pyrexia of unknown origin and unresolved suspicion of a CNS tumour (the more frequent application is for differential diagnosis of post-treatment radiation necrosis versus recurrence, rather than diagnosis at presentation). Impressive data have been obtained in the diagnosis of paraneoplastic syndromes and a variety of vasculitides and arteritides [7]. Occasionally PET has helped in the evaluation of patients with malignant paragangliomas and carcinoid tumours [8], and PET/CT holds promise for this indication.

Staging and re-staging It is in the setting of cancer staging that PET/CT comes into its own. Combined PET/CT has been shown to be superior to other imaging modalities in most tumour types. A gain of 20% was documented when the TNM tumour classification was used as the comparator and PET/CT was compared with whole-body MRI [9]. In lymphoma, PET/CT is better than CT in the diagnosis of both nodal and extranodal disease, and can detect disease in normal-sized lymph nodes that will be overlooked by CT. As a consequence, PET/CT upstages approximately 40% of all cases of lymphoma. PET/CT is also better than CT for the purpose of post-therapy evaluation owing to its greater predictive value: a positive post-treatment PET study is associated with poorer prognosis, whilst a scan performed after the first cycle of treatment is often predictive of response, especially in cases of aggressive Hodgkin’s disease and non-Hodgkin’s lymphoma [10–12]. FDG PET is useful to guide adoptive immunotherapy with donor lymphocyte infusions post transplant [13]. With regard to non-small cell lung cancer (NSCLC), three major studies have shown that PET/CT prevents unnecessary surgery in one out of five patients deemed operable by other criteria [14–16]. This is because PET/CT upstages a large proportion of patients by demonstrating both soft tissue and skeletal involvement. A further study found that PET/CT resulted in a change in management in 30% of patients with NSCLC [17]. Recently, Goren et al [18] discussed the relative roles of CT, PET and The British Journal of Radiology, January 2006

endoscopic-guided ultrasound with needle aspiration in the management of patients with lung cancer. There is a clear clinical role for PET/CT in colorectal cancer. It is of value for staging of recurrent disease, detection of liver involvement, detection of local recurrence, differential diagnosis of recurrent disease from scar and assessment of patients who present with rising tumour markers [19–21]. A meta-analysis carried out over a 5-year period showed that FDG PET changed the management in approximately 35% of patients in the setting of colorectal cancer. Often PET/CT demonstrates multiple liver deposits not seen on other imaging modalities [22]. A case could now be made that PET/CT should be the first imaging modality to be employed in the staging and re-staging of colorectal cancer. PET/CT is also applied to the staging and re-staging of patients with cancers in the head and neck, breast, oesophagus, pancreas, cervix and testicle, as well as patients with sarcomas and melanomas. In the head and neck, PET/CT misses micrometastatic disease (as do all imaging modalities) but it is useful in the context of upstaging N0 disease [23, 24]. In patients who present with cervical adenopathy and negative crosssectional imaging (CT/MRI), PET/CT is a useful investigation [25]. Patients with advanced disease tend to be upstaged with PET/CT. PET/CT is useful in disease monitoring after therapy (surgery, chemotherapy or radiation), but the optimal timing of this application remains controversial. The possibility of a false positive inflammatory response must be borne in mind. In thyroid cancer, PET/CT should be restricted to the re-staging of patients with raised serum markers (thyroxine-binding globulin, calcitonin, carcinoembryonic antigen) who present with negative cross-sectional imaging and negative 131I scans [26, 27]. In breast cancer, PET/CT is not used to stage the axilla owing to its failure to detect micrometastatic disease. PET/CT is, however, useful in re-staging, in the detection of nodal disease and in the visualization of distant disease in unsuspected sites. PET/CT scanning uncovers deposits in the skeleton and can be helpful in the evaluation of internal mammary and mediastinal node involvement. It also appears useful in the evaluation of response to treatment, absence of response on PET/CT carrying a worse prognosis. Scarring and fibrotic masses can be distinguished from active disease on the basis of FDG uptake. In the curative setting, PET/CT is used for the investigation of the nodal spread of oesophageal cancer. Here, PET/CT is better than CT alone. A growing body of evidence shows the utility of PET/CT in the evaluation of response to therapy. A study by Weber et al [28] investigated 40 patients. A PET study was performed at baseline and 2 weeks after initiation of chemotherapy. The first scan had a sensitivity of 93% and a specificity of 95%. Patients who responded to therapy had a reduction in FDG uptake by 54%, whilst in non-responders the reduction in FDG was of the order of 15% or less. In a similar study by Brucher et al [29], 27 patients with oesophageal carcinoma were given chemotherapy and radiotherapy. Patients responding to the treatment had a reduction in FDG uptake of 72%, whereas those who did not respond had a reduction of only 22%. Most studies of this type now point to the utility of FDG PET in the assessment of early response to treatment. 33

P J Ell

In melanoma patients, PET/CT is not useful for initial staging or in early disease, but it is of value for re-staging of more advanced disease. Melanoma metastases are intensely FDG avid. PET/CT is also used in the re-staging of patients with carcinoma of the cervix. Recurrent disease can be distinguished from non-viable necrotic or fibrotic post-therapy tissue. PET/CT has been used in a variety of other cancer types, such as GIST tumours, mesotheliomas, multiple myelomas and sarcomas. Pancreatic cancer, neuroendocrine tumours and germ cell tumours and their deposits can all exhibit intense FDG uptake. In contrast, prostate cancer and deposits from this tumour often exhibit poor FDG avidity; hence PET/CT with FDG is not useful in this context.

record data identical to those that would be obtained using a conventional CT. In patients with cancer, radiation exposure should often be considered of secondary importance, given their age, survival rates and therapeutic aspects. It can therefore be argued that PET/CT should become the first imaging study in a significant proportion of patients with cancer. From the above it can be seen that the CT information obtained from the PET/CT instrument can also be used for the purpose of volume planning and that the available PET information can be similarly used to better delineate tumour margins, whilst also distinguishing viable from non-viable tumour and aggressive from less aggressive disease. Ultimately, a more rational approach to radiotherapy planning is an achievable goal. Data are beginning to accrue that confirm this approach and its utility [41].

Treatment monitoring In part, this application has already been alluded to. It is evident that a metabolic response can precede a change in tumour size, and a reduction in FDG uptake can be seen within a matter of hours in patients with lymphoma or germ cell cancer in whom treatment is effective [30, 31]. Eventually PET/CT will be used to assess the biology of the individual tumour and its response to treatment [32], with novel markers aimed at imaging proliferation [33–36], hypoxia, angiogenesis, apoptosis, etc. PET/CT is useful to assess the efficacy of novel therapies. This has been demonstrated with Gleevec in the treatment of germ cell cancers, but PET/CT will have wide applicability in a number of new settings. It will be used as a surrogate marker for drug response, and this might imply yet another revision of the established but still insufficiently used RECIST criteria for tumour response to therapy. Eary et al [37], studied the effect of tumour heterogeneity, reflected in heterogeneity in FDG uptake, in patients with sarcomas. A 30% increase in risk of death was observed for every increase of 1 standard deviation (SD) in tumour heterogeneity, and there was a 12% increase in risk of death for every increase of 1 SD in the maximum SUV. However, the concept of a metabolic response as assessed by FDG will need to be validated in larger studies. In breast carcinoma patients treated with Tamoxifen, a flare response, albeit transient, has been described [38]. When such a response occurs it tends to do so 8–10 days after the commencement of Tamoxifen, and is usually an indicator of subsequent patient response to the treatment [38]. It is also recognized that patients studied soon after radiotherapy may exhibit an increase in FDG activity owing to an inflammatory response [39]. MacManus et al [40] have nevertheless shown the utility of evaluation of the metabolic response by PET in patients with NSCLC.

Radiotherapy planning A PET/CT scanner can be used to inform radiotherapy planning. The CT component of the instrument is identical to a conventional spiral CT and modern PET/CT scanners are available with 4-, 8- or 16-slice spiral CT scans. The CT component can be used for attenuation purposes only, in order to aid in the localization of the abnormality seen on the PET scanner, or it can be used at high power to 34

Imaging the skeleton With PET, it is possible to obtain data from skeletal metastases via two tracers: 18F-labelled fluoride ion, which is directly taken up by the skeleton, and 18F-FDG. More data need to be obtained before final recommendations can be made regarding the use of these two tracers for skeletal imaging. It is already apparent, however, that in many cancers, FDG can demonstrate both soft tissue and skeletal involvement; indeed, it has been advocated that conventional bone scanning is no longer required when staging NSCLC patients with FDG. In multiple myeloma, FDG is superior to conventional bone scanning in the detection of bony deposits. If scanner availability and tracer costs were not limiting factors, 18F-fluoride scanning of the skeleton would come to replace the conventional bone scan owing to the merits of PET/CT co-registration in the context of both malignant and benign bone disease [42–44].

Future developments Multimodality imaging is here to stay and image fusion will become routine. The first truly routine implementation of image fusion involving a large number of patients has been achieved with PET/CT. The design properties of PET/MRI are under consideration, and progress has already been made in this field with small animal scanners. The next generation of PET/CT technology is likely to make use of new radiation detectors and electronics. Discussions are now focusing, for example, on the reduction of whole-body imaging times to less than 15 min and the introduction of routine respiratory and cardiac gating for improvement of lesion localization and margin definition. Multiple slice spiral CT scans will open the way for cardiac imaging, and interesting developments are expected in this field, which, as with nuclear medicine in general, is heavily dependent on the emergence of new, clinically useful ligands. There is realistic hope that these new ligands will lead to novel practical applications in neurology, cardiology and oncology. As individually tailored medicines begin to impact on healthcare, these technologies will find special relevance in determining patient response to these therapies. An early indicator of lack of response may be not only beneficial but also immensely important in economic terms. The future for The British Journal of Radiology, January 2006

PET/CT to improve patient management

PET/CT imaging as a surrogate endpoint for novel therapeutic interventions is bright. This will imply a rethink of traditional criteria for lesion response – conventional RECIST criteria will need to be re-assessed in the light of the metabolic parameter made available by PET [45].

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17. Keidar Z, Haim N, Guralnik L, et al. PET/CT using 18-FDG in suspected lung cancer recurrence: diagnostic value and impact on patient management. J Nucl Med 2004;45:1640–6. 18. Goren HJM, Slijfer DTh, deVries EGE. Positron emission tomography, computerized tomography, and endoscopic ultrasound with needle aspiration of lung cancer. ASCO Educational Book 2005:579–86. 19. Arulampalam THA, Costa DC, Loizidou M, Visvikis D, Ell PJ, Taylor I. Positron emission tomography in colorectal cancer. Br J Surg 2001;88:176–89. 20. Arulampalam THA, Francis DL, Visvikis D, Taylor I, Ell PJ. FDG-PET for the pre-operative staging of colorectal liver metastases. Eur J Surg Oncol 2004;30:286–91. 21. Huebner RH, Park KC, Shepherd JE, Schwimmer J, Czernin J, Phelps ME, et al. A meta-analysis of the literature for whole-body FDG PET detection of recurrent colorectal cancer. J Nucl Med 2000;41:1177–89. 22. Dietlein M, Weber W, Schwaiger M, Schicha H. [18FFluorodeoxyglucose positron emission tomography in restaging of colorectal cancer. Nuklearmedizin 2003;42:145–56. 23. Hyde NC, Prvulovich E, Newman L, Waddington WA, Visvikis D, Ell PJ. A new approach to pre-treatment assessment of the NO neck in oral squamous cell carcinoma: the role of sentinel node biopsy and positron emission tomography. Oral Oncol 2003;29:350–60. 24. Myers LL, Wax MK, Nabi H, Simpson GT, Lamonica D. Positron emission tomography in the evaluation of the N0 neck. Laryngoscope 1998;108:232–6. 25. Hughes SJ, Prvulovich EM, Witherow H, Kalavrezos N, Ell PJ. A comparison of FDG PET/CT and MRI versus histology for staging of primary head and neck cancers and detection of recurrent disease. J Nucl Med 2004;45:80. 26. Chung JK, So Y, Lee JS, Choi CW, Lim SM, Lee DS, et al. Value of FDG PET in papillary thyroid carcinoma with negative 131I whole-body scan. J Nucl Med 1999;40:986–92. 27. Schluter B, Bohuslavizki KH, Beyer W, Plotkin M, Buchert R, Clausen M. Impact of FDG PET on patients with differentiated thyroid cancer who present with elevated thyroglobulin and negative 131I scan. J Nucl Med 2001;42:71–6. 28. Weber WA, Ott K, Becker K, Dittler HJ, Helmberger H, Avril NE, et al. Prediction of response to preoperative chemotherapy in adenocarcinomas of the esophagogastric junction by metabolic imaging. J Clin Oncol 2001;19:3058–65. 29. Brucher BL, Weber W, Bauer M, Fink U, Avril N, Stein HJ, et al. Neoadjuvant therapy of esophageal squamous cell carcinoma: response evaluation by positron emission tomography. Ann Surg 2001;233:300–9. 30. Hoekstra OS, Ossenkoppele GJ, Golding R, van Lingen A, Visser GW, Teule GJ, et al. Early treatment response in malignant lymphoma, as determined by planar fluorine-18fluorodeoxyglucose scintigraphy. J Nucl Med 1993;34:1706–10. 31. Oliver T, Shamash J, Powles T, Somassundram U, Ell PJ. 20 years phase K study of single agent carboplatin in metastatic seminoma: could it have been accelerated by 72 hour PET scan response? ASCO, New Orleans, June 2004. 32. Brugarolas J, Clark JW, Chabner B. Using ‘‘rationally designed drugs’’ rationally. Lancet 2003;361:1758–9. 33. Francis DL, Visvikis D, Costa DC, Croasdale I, Arulampalam TH, Luthra SK, et al. Assessment of recurrent colorectal cancer following 5-fluorouracil chemotherapy using both 18FDG and 18FLT PET. Eur J Nucl Med Mol Imaging 2004;31:928. 34. Francis DL, Freeman A, Visvikis D, Costa DC, Luthra SK, Novelli M, et al. In vivo imaging of cellular proferation in colorectal cancer using Positron Emission Tomography. Gut 2003;52:1602–6.

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P J Ell 35. Francis DL, Visvikis D, Costa DC, Arulampalam THA, Townsend C, Luthra I, et al. Potential impact of [18F]39deoxy-39-fluorothymidine versus [18F] fluoro-2-deoxy-D-glucose in Postron Emission Tomography for colorectal cancer. Eur J Nucl Med 2003;30:988–94. 36. Shields AF, Grierson JR, Dohmen BM, et al. Imaging in vivo proliferation with 18FLT and positron emission tomography. Nature Med 1998;11:1334–6. 37. Eary JF, Brenner W, Vernon C, and O’Sullivan F. Tumor heterogeneity in sarcoma patients is a significant predictor of survival. Eur J Nucl Med Mol Imaging 2004;31 Suppl. 2:S232. 38. Dehdashti F, Flanagan FL, Mortimer JE, Katzenellenbogen JA, Welch MJ, Siegel BA. Positron emission tomographic assessment of ‘‘metabolic flare’’ to predict response of metastatic breast cancer to antiestrogen therapy. Eur J Nucl Med 1999;26:51–6. 39. Strauss LG. Fluorine-18-deoxyglucose and false-positive results: a major problem in the diagnostics of oncological patients. Eur J Nucl Med 1996;23:1409–15. 40. MacManus MP, Hicks RJ, Matthew JP, et al. Positron emission tomography is superior to computed tomography scanning for response-assessment after radical radiotherapy or chemotherapy in patients with non-small-call lung cancer. J Clin Oncol 2003;21:1285–92.

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41. Scarfone C, Lavely WC, Cmelak AJ, Delbeke D, Martin WH, Billheimer D, et al. Prospective feasibility trial of radiotherapy target definition for head and neck cancer using 3-dimentional PET and CT imaging. J Nucl Med 2004;45:543–52. 42. Schirrmeister H, Guhlmann A, Kotzerke J, Santjohanser C, Kuhn T, Kreienberg R, et al. Early detection and accurate description of extent of metastatic bone disease in breast cancer with fluoride ion and positron emission tomography. J Clin Oncol 1999;17:2381–9. 43. Cook GJ, Houston S, Rubens R, Maisey MN, Fogelman I. Detection of bone metastases in breast cancer by 18FDG PET: differing metabolic activity in osteoblastic and osteolytic lesions. J Clin Oncol 1998;16:3375–9. 44. Gayed I, Vu T, Johnson M, Macapinlac H, Podoloff D. Comparison of bone and 2-deoxy-2-[18F]fluoro-D-glucose positron emission tomography in the evaluation of bony metastases in lung cancer. Mol Imaging Biol 2003;5:26–31. 45. Schuetze SM, Eary JF, Griffith KA, Rubin BP, Hawkins DS, Vernon CB, et al. FDG PET but not RECIST agrees with histological response of soft tissue sarcoma to neoadjuvant chemotherapy. ASCO 2005.



E


Mesenteric panniculitis in oncologic patients: PET-CT findings 1,2,3

R ZISSIN, MD,

1,3

U METSER, MD, 4D HAIN, MD and

1,3

E EVEN-SAPIR, MD, PhD

1

Department of Nuclear Medicine, Tel-Aviv Sourasky Medical Center, and the 2Department of Diagnostic Imaging, Sapir Medical Center, Kfar Saba, both affiliated to the 3Sackler Faculty of Medicine, Tel-Aviv University, Tel-Aviv, and 4 Nuclear Medicine Institute, Shaare Zedek Medical Center, Jerusalem, Israel

Abstract. The aim of this study is to assess the role of PET/CT in differentiating between mesenteric panniculitis (MP) and co-existing MP and mesenteric tumoural involvement. A total of 33 PET/CT examinations, of 19 oncologic patients (16 men and three women with ages ranging from 48 years to 83 years) with findings of MP on the CT part of the study were retrospectively reviewed. The FDG uptake in mesenteric nodules was recorded. The final diagnosis of malignant mesenteric involvement was based on clinical and imaging follow-up. Based on the FDG uptake in mesenteric nodules, patients were categorized as group A: increased mesenteric uptake (n58) and group B: no mesenteric uptake (n511). In seven of the eight patients in group A, a co-existing MP and mesenteric tumour involvement was found: one patient had a recurrent cervical carcinoma and the other six patients had lymphoma. In four of these six patients, the positive PET findings disappeared on followup PET/CT with complete remission while the CT findings of the MP remained unchanged. In the other two, the PET findings progressed along with clinical deterioration. In the last patient of group A, with rectal carcinoma without evidence of recurrence, the mesenteric FDG uptake was a false positive uptake. In all 11 patients with CT findings of MP and negative PET, no malignant involvement of the mesentery was diagnosed. To conclude, a negative PET has a high diagnostic accuracy in excluding tumoural mesenteric involvement while increased uptake suggests the co-existing of mesenteric deposits, particularly in patients with lymphoma.

Mesenteric panniculitis (MP), also entitled liposclerotic mesenteritis, mesenteric lipodystrophy, mesenteric lipomatosis and lipogranuloma of the mesentery, is a benign condition characterized by non-specific inflammation involving the adipose tissue of the mesentery, with acute inflammatory changes and fat necrosis being the predominant histological findings. In its chronic phase when fibrosis is dominant, the disease is known as retractile mesenteritis [1–4]. Sclerosing mesenteritis seems the most appropriate diagnostic term of this entity, characterized by a spectrum of histological findings [4]. The specific aetiology of the disease is unknown, although various causes have been suggested, including infection, trauma or ischaemia of the mesentery. The disease has been related to other pathological processes such as vasculitis, granulomatous disease, pancreatitis and malignancy [2]. Its prevalence in abdominal CT examinations is approximately 0.6%, commonly appearing as an incidental finding, mostly in middle or late adulthood [5]. An association between MP and pre-existing malignancy has been reported [5, 6]. The CT features of MP are well recognized and may suggest the diagnosis, but they are non-specific and can appear in other conditions such as mesenteric oedema, granulomatous diseases, primary or secondary abdominal neoplasms and lymphoma [1]. In cases of MP and known intra-abdominal malignancies, differentiating MP from tumoural involvement of mesenteric lymph nodes (LNs) is of crucial importance. 18 F-Fluorodeoxyglucose (FDG)/PET imaging has been introduced in addition to conventional cross-sectional Received 17 March 2005 and in final form 3 May 2005, accepted 1 June 2005. Address correspondence to: Einat Even-Sapir, Department of Nuclear Medicine, Tel-Aviv Sourasky Medical Center, 6 Weizman Street, TelAviv, 64239 Israel.


imaging methods in the routine practice of oncologic patients. Recently, hybrid systems composed of PET and CT have been introduced and its use is increasing steadily [7]. PET and CT are performed at the same clinical setting resulting with generation of fused PET/CT images, which provides both functional and anatomical data. The potential role of PET/CT in differentiating benign MP from tumoural mesenteric involvement is the topic of the current study. We have reviewed 19 patients with a history of known malignancy, who had incidental MP on the CT component of the PET/CT study, and report the PET/CT features of MP in this oncologic population.

Material and methods The clinical data and PET/CT findings of 19 consecutive patients with MP incidentally diagnosed on the CT part of the study were retrospectively reviewed. The patient group consisted of 16 men and 3 women with ages ranging from 48 years to 83 years (mean age 62¡11 years). Five of the 19 patients underwent a PET/CT study for staging and 14 for suspected recurrence or for monitoring response to treatment. Known malignancies included lymphoma (n510), colorectal cancer (n55), melanoma (n52) and lung and cervical carcinomas, one patient each. A total of 33 PET/CT examinations were performed and reviewed in these 19 patients as 11 of them had one to three follow-up (F/U) studies. The patients fasted at least 4 h prior to the intravenous (IV) injection of 370–666 MBq (10–18 mCi) FDG. Iodinated oral contrast material was administered prior to FDG injection. Glucose levels had been checked prior to the injection of FDG. A PET/CT study was performed only when blood glucose levels were bellow 8.32 mmo l21. Scanning from the base of the skull through the mid-thigh 37

R Zissin, U Metser, D Hain and E Even-Sapir

was performed using the Discovery LS PET/CT system (GE Medical Systems, Milwaukee, WI). Low-dose CT acquisition was performed first with 140 kV, 80 mA, 0.8 s per CT rotation, a pitch of 6 and a table speed of 22.5 mm s21, without any specific breath-holding instructions. A PET emission scan was carried out immediately following acquisition of the CT, without changing the patient’s position. From 5 to 8 bed positions were performed with an acquisition time of 5 min for each one. CT data were used for attenuation correction. Images were generated and interpreted on work station (Xeleris Elgems, Haifa, Israel) equipped with fusion software that enables the display of PET, CT and fused PET/CT images. The CT criteria for the diagnosis of MP included a welldefined, inhomogeneous fatty mass with higher attenuation than the normal retroperitoneal fat, occasionally with preserved perivascular fat, that contained small nodules and surrounded by a tumoural pseudocapsule (Figure 1) [5]. When interpreting the PET/CT, the uptake of FDG in the mesenteric nodules was reported. Standardized uptake value (SUV) was measured for any focal increased uptake within the CT mesenteric abnormalities. The latter semiquantitative parameter was automatically obtained on the patient’s final report and was calculated as the ratio of activity in tissue per millilitre to the activity in the injected dose per patient body weight. The final diagnosis of the mesenteric pathology was based on clinical and imaging (PET/CT and/or diagnostic CT) F/U: co-existing MP and malignancy was concluded when improvement or disappearance of the mesenteric abnormalities were seen on F/U imaging together with clinical evidence of a favourable response to therapy, or when mesenteric abnormalities progression was seen on F/U imaging along with clinical and imaging evidence of disease progression. The MP was regarded as benign if the mesenteric findings remained stable in a patient who was clinically disease-free or if the findings remained stable on

F/U imaging while other sites of disease resolved in response to therapy and the patient was clinically considered in complete remission.

Results Based on the FDG uptake within the MP, the study patients were divided into two groups: group A consisting of eight patients with MP and focal increased FDG uptake within mesenteric nodules and group B consisting of 11 patients with MP without increased FDG uptake.

Group A: FDG uptake within MP FDG uptake was demonstrated within CT mesenteric abnormalities, indistinguishable from a benign MP, in eight patients. The clinical and imaging findings of these patients are summarized in Table 1. A conclusion of malignant mesenteric involvement was made in seven of these patients, one with a metastatic cervical cancer and six with non-Hodgkin’s lymphoma (NHL): In four lymphoma patients both the increased FDG uptake and the nodules themselves resolved following chemotherapy, while other CT findings of MP remained unchanged on F/U PET/CT (Figure 2). In the other two, clinical F/U and repeat PET/CT were consistent with tumour progression (Figure 3). In the case of metastatic cervical carcinoma new mesenteric PET findings appeared within known MP, seen previously on two PET/CT studies, along with clinical evidence of tumour recurrence. In coexisting MP and mesenteric tumoural involvement, the MP changes, seen on the CT part of the examination, which were not associated with increased FDG uptake, remained unchanged on F/U PET/CT studies. In the remaining patient with rectal cancer and focal mesenteric FDG uptake (SUV – 3.5) the positive PET was regarded as a false positive study as there was no clinical or imaging evidence of active tumour, together with stability of the MP findings on F/U diagnostic CT studies during a long disease-free period of 28 months.

Group B: MP without increased FDG uptake In 11 patients no FDG uptake was seen within typical features of MP. The mesenteric soft-tissue nodes ranged between immeasurable, numerous small nodules, to discrete nodes measuring up to 0.9 cm in the short axis and 1.9 cm in the long axis (Figure 4). In all these patients the mesenteric abnormalities seen on the CT part of the study were stable on imaging F/U of a mean of 10.5 months (range: 5–30 months) and we therefore believe that the mesenteric findings were benign.

Discussion Figure 1. CT findings of mesenteric panniculitis (MP). Nonenhanced abdominal CT at the mid-abdomen shows a welldefined, inhomogeneous fatty lesion, with higher attenuation than the normal retroperitoneal fat, confined by a highlyattenuated stripe representing a tumoural pseudocapsule (thick arrows), with an engorged mesenteric vessel and scattered discrete nodules of soft-tissue density, some of which are engulfed by a hypodense fatty halo (thin arrow).

38

MP is a non-neoplastic inflammatory process of unknown aetiology, affecting the small bowel mesentery. It was rarely diagnosed before the era of ultrasound and CT, but currently it is not uncommonly encountered, often as an incidental imaging finding. Male predominance, as was found in our group, has been previously reported [1–3] though a slight female predominance has been reported in ?a single publication [5]. Most cases of MP are The British Journal of Radiology, January 2006

PET-CT finding of mesenteric panniculitis Table 1. Clinical, PET/CT and F/U of eight patients with CT findings of mesenteric panniculitis (MP) and increased FDG uptake within mesenteric nodules Patient no., sex, age (years)

Primary tumour

Indication for the 1st PET/CT

PET/CT findings (on the 1st study)

Final diagnosis and imaging F/U

1. M, 76

NHL

Staging at diagnosis

CT: increased-attenuation fat, pseudocapsule and numerous, slightly enlarged soft-tissue nodules PET: uptake (SUV – 13.2) in a 1.3 cm mesenteric nodule

Co-existing MP and mesenteric lymphoma

2. M, 61

3. M, 50

4. M, 55

5. F, 57

6. M, 58

NHL

NHL

NHL

NHL

NHL

Monitoring response to treatment. - No baseline study

Restaging for suspected recurrence



Monitoring response to treatment. - No baseline study

CT: markedly increased attenuation of the mesenteric fat, enlarged mesenteric LNs (up to 2.5 cm62 cm), pseudocapsule, ‘‘fat ring’’ sign. PET: diffuse uptake (SUV – 2.7) in enlarged mesenteric nodules, and in mesenteric fat

CT: increased-attenuation fat, pseudocapsule and numerous, slightly enlarged soft-tissue nodules PET: uptake (SUV – 3.6) in a 0.8 cm mesenteric nodule

CT: increased-attenuation fat, pseudocapsule and numerous, slightly enlarged soft-tissue nodules PET: uptake (SUV – 4.5) in a 1 cm61.3 cm mesenteric nodule

CT: increased-attenuation fat, pseudocapsule and numerous, slightly enlarged soft-tissue nodules PET: uptake (SUV – 8.2) within markedly enlarged mesenteric nodules, up to 2.3 cm61.7 cm

CT: increased-attenuation fat, pseudocapsule and numerous, slightly enlarged soft-tissue nodules PET: uptake (SUV – 3.8) in several mesenteric nodules, up to 0.8 cm.

F/U PET/CT (4 M later): the hypermetabolic mesenteric nodule enlarged to 3.8 cm64 cm (SUV – 18.8) with the appearance of two new hypermetabolic nodules (1.1 cm61.5 cm, SUV – 7.5), indicating disease progression No change in the other MP findings Co-existing MP and mesenteric lymphoma

F/U PET/CT (5 M later-without treatment): no change in the mesenteric abnormalities and their uptake. The patient was clinically considered with active disease Co-existing MP and mesenteric lymphoma

F/U PET/CT (17 M later-following chemotherapy): no focal increased uptake. No change in the other MP findings. Co-existing MP and mesenteric lymphoma

F/U PET/CT (6 M later – following chemotherapy): No focal increased uptake No change in other MP findings Co-existing MP and mesenteric lymphoma

F/U PET/CT (4 M later-following chemotherapy): the mesenteric lymphadenopathy decreased in size to 1 cm, FDG uptake disappeared No change in other MP findings Co-existing MP and mesenteric lymphoma

F/U PET/CT (2 M later- following chemotherapy): the hypermetabolic mesenteric nodules and FDG uptake disappeared No change in other MP findings (Continued)


39

R Zissin, U Metser, D Hain and E Even-Sapir Table 1. (Cont.) Clinical, PET/CT and F/U of eight patients with CT findings of mesenteric panniculitis (MP) and increased FDG uptake within mesenteric nodules Patient no., sex, age (years)

Primary tumour

7. F, 56

Metastatic cervical carcinoma

8. F, 55

Rectal carcinoma

Indication for the 1st PET/CT Monitoring response to treatment. - No baseline study

Misty mesentery on a diagnostic CT, performed for F/U, negative markers

PET/CT findings (on the 1st study)

Final diagnosis and imaging F/U

CT: increased-attenuation fat, pseudocapsule and numerous, slightly enlarged soft-tissue nodules PET: negative

CT: increased-attenuation fat, pseudocapsule and numerous, enlarged LNs up to 1.4 cm62.5 cm PET: uptake (SUV – 3.5) in 3 small mesenteric nodules. No uptake in the enlarged nodes

Co-existing MP and metastases

F/U PET/CT (11 M later): no change F/U PET/CT (after 5 M): disease progression with a new 1.7 cm60.9 cm mesenteric nodule with FDG uptake (SUV – 5.7) No change in other MP findings No clinical evidence for active disease: A false positive PET

- No change in the MP findings in a previous CT study, 2 years earlier and in a F/U diagnostic CT after 4 M, negative markers

F/U, follow-up; M, months; LN, lymph node.

asymptomatic and are incidentally detected on abdominal CT performed for unrelated conditions [5]. On CT, MP appears as a mass of increased-attenuation mesenteric fat containing small soft-tissue nodes, with a maximal transverse diameter directed toward the left abdomen consistent with the orientation of the jejunal mesentery. The infiltrated fat typically engulfs the mesenteric vessels and displaces adjacent bowel loops without invading them [1, 2, 5]. Hypodense, cystic-like areas and calcifications due to fat necrosis are infrequently seen within this mass [1]. Increased fatty attenuation and small mesenteric nodules, also termed ‘‘misty mesentery’’ may, however, be seen in any pathological process infiltrating the mesentery, such as inflammation, oedema, haemorrhage or metastases [8]. Two CT findings are considered more specific for the diagnosis of MP as they have not been reported in other mesenteric diseases: the presence of tumoural pseudocapsule (found in up to 60% of MP cases) and the ‘‘fat ring’’ sign of hypodense fatty halo surrounding mesenteric nodules and vessels (seen in up to 75% of cases) [1, 2, 4]. Daskalogianki et al have reported the co-existence of MP and various neoplastic diseases, especially lymphoma and gastrointestinal and urogenital adenocarcinomas, in up to 69% of patients with MP [5]. 10 of the 19 study patients with CT findings of MP had lymphoma as the underlying malignancy. Co-existing MP with malignant mesenteric involvement was found in six of the lymphoma patient (60%), representing 85.7% of the 7 study patients with malignant mesenteric involvement. In oncologic patients, therefore, the small soft-tissue mesenteric nodules typically seen within the infiltrated mesenteric fat of MP may be misdiagnosed as metastatic implants. On the other hand, metastatic deposits within a pre-existing MP can also be present, as was described in a single case report on a patient with uterine papillary serous adenocarcinoma in whom multiple nodular metastases were detected on CT within typical MP findings [9]. 40

The results of our study emphasise the potential role of PET/CT in differentiating benign MP and MP with mesenteric tumoural involvement. Fused PET/CT images provide both metabolic and anatomic information with a high accuracy. On CT, lymph node pathology is based on size criteria alone. Enlarged lymph nodes may be reactive while normal-sized nodes may contain early metastatic deposits, which can be reliably detected by the functional (PET) part of the study. The majority of our patients, including those with a malignant mesenteric involvement, had only subtle CT findings and the differentiation between benign and malignant causes could not be made with confidence based on the CT alone. Our results suggest a potential role for integrated PET/CT in the assessment of MP detected on CT in oncologic patients. PET/CT study can be used to correctly exclude mesenteric tumoural involvement when no FDG uptake is seen within typical CT features of MP. Alternatively, in a patient with an oncologic history, the demonstration of FDG uptake, even in small-sized nodules within characteristic CT findings of MP, is highly suggestive of neoplastic involvement of the mesentery. In PET/CTs of co-existing MP and mesenteric metastatic deposits, the increased FDG uptake was detected in nodules smaller than the benign nodules of the MP that had no increased uptake. The increased FDG uptake of these malignant mesenteric deposits resolved on a F/U study following a favourable response to treatment while the findings of the benign MP remained unchanged. Increased FDG uptake is, however, not tumour-specific as FDG uptake may be seen in benign inflammatory conditions [10], as was the case in one of our patients in whom a slightly increased FDG uptake was detected within MP findings without evidence of malignancy on a long-term F/U of 28 months. As the most consistent histological finding of MP is the presence of an inflammatory infiltrate, it may explain the uptake in that The British Journal of Radiology, January 2006

PET-CT finding of mesenteric panniculitis

(a)

(b) Figure 2. A 50-year-old man with follicular lymphoma: mesenteric panniculitis (MP) with meseneteric tumoural involvement before and after a favourable response to chemotherapy. (a) A fused PET/CT image shows increased 18F-FDG uptake in an 8 mm nodule (dashed arrow) in the background of MP (arrows). The latter appears as a mesenteric mass of inhomogeneous fatty tissue containing scattered soft-tissue nodules which are not 18FDG-avid. (b) PET/CT images at diagnosis (top images) and following chemotherapy (lower images) show regression in size of the nodule and disappearance of 18FDG uptake (arrows). No change is seen in the other findings of the MP.


41

R Zissin, U Metser, D Hain and E Even-Sapir

(a)

(b) Figure 3. A 76-year-old man with non-Hodgkin’s lymphoma (NHL): mesenteric panniculitis (MP) with meseneteric tumoural involvement with disease progression. PET/CT images at diagnosis (top images) and 4 months later (lower images): (a) At the mid-abdomen typical findings of MP with no 18F-FDG uptake are seen, stable on F/U. (b) More caudally, increased 18F-FDG uptake is detected at diagnosis within a 1.2 cm61.5 cm mesenteric nodule (SUV – 13.2) (arrows). On F/U the hypermetabolic node, most likely involved with lymphoma, enlarged to 3.8 cm64 cm with increasing 18F-FDG uptake (SUV – 18.8) (arrows). 42


PET-CT finding of mesenteric panniculitis

Acknowledgments The authors wish to thank Mrs Limor Zuriel, MSc, for her assistance in the preparation of the manuscript.

References

Figure 4. A 78-year-old man with diffuse B cell lymphoma and co-existing mesenteric panniculitis (MP). A fused PET/CT image shows typical findings of MP, including inhomogeneous fatty density mass with well-defined nodules of soft-tissue density, confined by a pseudocapsule (arrows), without 18F-FDG uptake. The findings remain stable on a F/U PET/CT, 5 months later, with no clinical evidence of active lymphoma.

case [4]. We have found a single case report in the English literature regarding the FDG/PET in a patient with sclerosing mesenteritis; a large, speculated, soft-tissue mesenteric mass showed peripheral increased FDG uptake, probably representing the peripheral high metabolic inflammation and inactive central area of fibrosis [11]. The limitations of our study are the relatively small number of patients and the lack of a pathological proof for all lesions. However, as often happens in tumour imaging, not all detected lesions have histological diagnosis and their nature is sometimes based on clinical and imaging F/U. Validation of our findings in larger patient groups is warranted. To conclude, if MP is suspected on the CT part of the PET/CT study, special attention should be paid to the 18 F-FDG-avidity of the findings. A negative PET has high diagnostic accuracy in excluding tumoural mesenteric involvement while increased uptake may suggest the coexisting of metastatic deposits, particularly in patients with lymphoma.


1. Horton KM, Lawler LP, Fishman EK. CT findings in sclerosing mesenteritis (panniculitis): spectrum of disease. Radiographics 2003;23:1561–7. 2. Sabate JM, Torrubia S, Maideu J, Franquet T, Monill JM, Perez C. Sclerosing mesenteritis: imaging findings in 17 patients. AJR Am J Roentgenol 1999;172:625–9. 3. Parra-Davila E, McKenney MG, Sleeman D, Hartmann R, Rao RK, McKenney K, Compton RP. Mesenteric panniculitis: case report and literature review. Am J Surg 1998;64:768–71. 4. Emory TS, Monihan JM, Carr NJ, Sobin LH. Sclerosing mesenteritis, mesenteric panniculitis and mesenteric lipodystrophy: a single entity? Am J Surg Pathol 1997;21:392–8. 5. Daskalogiannaki M, Voloudaki A, Prassopoulos P, Magkanas E, Stefanaki K, Apostolaki E, et al. CT evaluation of mesenteric panniculitis: prevalence and associated diseases. AJR Am J Roentgenol 2000;174:427–31. 6. Kipfer RE, Moertel CG, Dahlin DC. Mesenteric lypodystrophy. Ann Intern Med 1974;80:582–8. 7. Schoder H, Larson SM, Yeung HW. PET/CT in oncology: integration into clinical management of lymphoma, melanoma, and gastrointestinal malignancies. J Nucl Med. 2004;45 Suppl. 1:72S–81S. 8. Mindelzun RE, Jeffrey RB Jr, Lane MJ, Silverman PM. The misty mesentery on CT: differential diagnosis. AJR Am J Roentgenol 1996;167:61–5. 9. La Fianza A, Alberici E, Di Maggio EM, Preda L, Campani R. Multiple nodular metastases in mesenteric panniculitis by uterine papillary serous adenocarcinoma [UPSC]: CT appearance of a case. Clin Imaging 1999;23:90–3. 10. Kazama T, Faria SC, Varavithya V, Phongkitkarun S, Ito H, Macapinlac HA. FDG PET in the evaluation of treatment for lymphoma: clinical usefulness and pitfalls. Radiographics 2005;25:191–207. 11. Nguyen BD. F-18 FDG PET demonstration of sclerosing mesenteritis. Clin Nucl Med 2003;28:670–1.

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E


Diagnostic efficacy of SonoVueH, a second generation contrast agent, in the assessment of extracranial carotid or peripheral arteries using colour and spectral Doppler ultrasound: a multicentre study 1

P S SIDHU, FRCR, 2P L ALLAN, FRCR, 3F CATTIN, MD, 4D O COSGROVE, FRCR, 5A H DAVIES, MD, D D DO, MD, 7S KARAKAGIL, MD, 8J LANGHOLZ, MD, 9D A LEGEMATE, MD, PhD, 10 A MARTEGANI, MD, 11J-B LLULL, MD, 12C PEZZOLI, PhD and 11A SPINAZZI, MD 6

1

Department of Radiology, King’s College Hospital, Denmark Hill, London SE5 9RS, UK, 2Department of Radiology, Edinburgh Royal Infirmary, Edinburgh, Lothian, UK, 3CHU de Besancon, Service de Radiologie B, Boulevard Fleming, F-25030 Besancon, Belgium, 4Department of Imaging, Royal Postgraduate Medical School, Hammersmith Hospitals Trust, Du Cane Road, London W12 OHS, UK, 5Department of Surgery, Imperial College School of Science, Technology & Medicine, Charing Cross Hospital, Fulham Palace Road, London W6 8RP, UK, 6Inselspital-Division of Angiology, Department of Internal Medicine, University of Berne, Freiburgstarsse 10, 3010 Berne, Switzerland, 7Department of Surgery, Uppsala University Hospital, Akademiska sjukhuset, 75185 Uppsala, Sweden, 8Schwerpunktpraxis fu¨r Angiologie, Wilsnacker Strabe 14, 10559, Berlin, Germany, 9Department of Vascular Surgery, Academic Medical Centre, Meibergdreef 9, NL-1105 AC DE Amsterdam, The Netherlands, 10Servizio di Radiologia, Ospedale Valduce, Via Dante Alighieri, 11, 22100 Como, Italy, 11Bracco Diagnostics Inc., 107 College Road East Princeton, Princeton, NJ 08540, USA and 12Bracco Imaging SpA, Via E. Folli 50, 20134 Milan, Italy

Abstract. The purpose of this study was to demonstrate the improvement in diagnostic quality and diagnostic accuracy of SonoVueH microbubble contrast-enhanced ultrasound (CE-US) versus unenhanced ultrasound imaging during the investigation of extracranial carotid or peripheral arteries. 82 patients with suspected extracranial carotid or peripheral arterial disease received four SonoVue doses (0.3 ml, 0.6 ml, 1.2 ml and 2.4 ml) with Doppler ultrasound performed before and following each dose. Diagnostic quality of the CE-US examinations was evaluated off-site for duration of clinically useful contrast enhancement, artefact effects and percentage of examinations converted from non-diagnostic to diagnostic. Accuracy, sensitivity and specificity were assessed as agreement of CE-US diagnosis evaluated by an independent panel of experts with reference standard modality. The median duration of clinically useful signal enhancement significantly increased with increasing SonoVue doses (p¡0.002). At the dose of 2.4 ml of SonoVue, diagnostic quality evaluated as number of inconclusive examinations significantly improved, falling from 40.7% at baseline down to 5.1%. Furthermore, SonoVue significantly (p,0.01) increased the accuracy, sensitivity and specificity of assessment of disease compared with baseline ultrasound. SonoVue increases the diagnostic quality of Doppler images and improves the accuracy of both spectral and colour Doppler examinations of extracranial carotid or peripheral arterial disease.

Colour and spectral Doppler ultrasound examination of the peripheral [1, 2] and extracranial carotid [3, 4] arterial systems is a well established non-invasive method of assessment of arterial disease. Frequently, Doppler ultrasound replaces conventional angiography with the associated cost savings and reduction in patient morbidity [5–8]. However, Doppler ultrasound does not always provide a full diagnostic assessment and there is no alternative but to seek confirmatory evidence of arterial disease with conventional angiography and increasingly with helical CT angiography and MR angiography [9–12]. A number of factors preclude a full Doppler ultrasound examination of a vessel: heavily calcified plaque causes acoustic shadowing, a deep-seated artery returns a poor echo-signal and vessel tortuosity precludes a satisfactory Doppler angle for accurate velocity measurements. In Received 6 May 2005 and accepted 9 June 2005.

44

order to improve the diagnostic capability of a Doppler ultrasound examination of the peripheral and carotid arteries, introducing an echo-enhancing agent would be expected to facilitate visualization of difficult arteries, thus overcoming inherent problems associated with ultrasound, and ultimately reducing unnecessary invasive and expensive diagnostic procedures. SonoVueH is the trademark name of a new ultrasound contrast agent (BR1, Bracco, Italy) [13]. SonoVue is a suspension of phospholipid stabilized sulphur hexafluoride (SF6) microbubbles. When reconstituted with normal saline the product is stable at room temperature for several days, but should be used after reconstitution within 6 h as the product contains no preservative [14]. Reconstitution produces a high microbubble concentration (up to 5 6 108 microbubbles ml21), a favourable size (90% of microbubbles smaller than 8.0 mm, mean diameter 2.5 mm) and strong echogenicity over the range of The British Journal of Radiology, January 2006

Diagnostic efficacy of SonoVue

frequencies used in medical ultrasound examinations [15]. The microbubbles produced are not trapped in the capillary vasculature, and the use of SF6 (an innocuous gas) renders the microbubbles more resistant to pressure increases from the left ventricle of the heart, increasing microbubble survival. SonoVue demonstrates a maximum backscatter coefficient at about 3 MHz and an elimination half-life of approximately 6 min. More than 80% of the compound is exhaled via the lungs in 11 min [16]. The efficacy of SonoVue in extracranial carotid or peripheral arterial disease was evaluated in a multicentre study where the quality end-points were as follows: to ascertain the optimal dosage with regard to global quality of images, to assess the duration of microbubble contrast effect, to aid the interpretation of diagnostically difficult colour and spectral Doppler ultrasound examinations and to evaluate the potential of contrast-enhanced ultrasound (CE-US) to change a non-diagnostic ultrasound examination into a diagnostic examination. In a subset study population, the diagnostic accuracy, sensitivity and specificity of SonoVue enhanced Doppler investigations were evaluated in terms of agreement of CE-US examinations in comparison with other recognized diagnostic imaging modalities.

Material and methods The overall study consisting of two (study A and B) parallel multicentre studies was aimed at investigating, with Doppler ultrasound and SonoVue, different vascular territories: renal, abdominal, cerebral, extracranial carotid or peripheral arteries, and the portal circulation. We refer here to the results pertaining to extracranial carotid or peripheral arteries. Local Medical Ethics Committees granted approval for the study at each hospital site according to local legal requirements and the study was conducted in accordance with the Declaration of Helsinki and European Good Clinical Practice. All patients recruited gave written informed consent.

Study population The study population comprised 82 male and female patients, over 18 years of age, with a suspected vascular pathology, referred for Doppler ultrasound investigations of carotid, iliac, femoral, popliteal or tibial arteries and for whom the observers could not make an interpretation with confidence at baseline unenhanced colour and spectral Doppler ultrasound examination. The main criteria for patient exclusion from the study were: severe congestive heart failure (New York Heart Association Class IV); unstable angina; severe cardiac arrhythmia; recent myocardial infarction; recent organ transplant or unstable neurological disease. Lactating women or women known or suspected to be pregnant were excluded. Patients were also excluded if they were critically ill, medically unstable or were in an intensive care setting. Patients receiving another investigational drug within 30 days prior to the study were not recruited. The British Journal of Radiology, January 2006

Study design The study was a multicentre, open-label (on-site), blinded (off-site), randomized, dose-ranging, cross-over study to compare four different doses of SonoVue (0.3 ml, 0.6 ml, 1.2 ml, and 2.4 ml) in Doppler ultrasound investigations of extracranial carotid (22 patients) or peripheral arteries (60 patients). Patients were randomized to one of four dose sequences, according to a randomization schedule with block size 4. The SonoVue doses were administered as an intravenous bolus injection over 20 s via a 20 gauge cannula (Introcan-W; Braun Melsungen AG, Germany) placed in the forearm. All centres employed the standard ultrasound machine used for routine colour Doppler ultrasound vascular examinations within the department. A variety of ultrasound machines were used, depending on the centre. Once the optimum colour and spectral Doppler ultrasound parameters were set for each patient at the baseline examination to extract maximum information (with the gain turned down to the lowest informative level), the parameters were unaltered for the remainder of the CE-US examination. For each patient, a vessel of interest was designated for further investigation, based on the vessel that would most likely drive the patient’s diagnosis. Doppler ultrasound investigations of the designated vessel were performed at baseline and after each injection of SonoVue, with SuperVideo Home System (S-VHS) videotape recording of images beginning 30 s prior to injection and continuing until the end of the microbubble contrast effect. At each time point, the designated vessel was studied first with either colour Doppler ultrasound or power Doppler ultrasound (only one mode was used for each patient depending on investigator choice), and then with spectral Doppler imaging. All SonoVue administration and imaging procedures were completed on the same day. The interval between administrations of the different doses was at least 10 min or until disappearance of the microbubble contrast effect from the previous administration.

Assessments Four independent experienced readers, paired for each of the studies (A and B) and unaffiliated with the study sites performed an off-site assessment of the recorded ultrasound images. These readers were blinded to study agent dose (whether baseline or post-dose), and patient information, including results of other imaging procedures. The S-VHS videotape recordings were divided into sets consisting of four post-injection images, one for each of the four SonoVue injection doses, plus the baseline images. Within each image set, an assigned random code number determined the order of presentation of patient images to the off-site readers. The off-site readers were provided with the identification of the vessel under investigation for each video sequence. Following completion of these unpaired assessments, the baseline and corresponding post-injection images for each dose of SonoVue in each patient were then assessed in matched pairs. For patients with an available reference diagnostic modality (conventional angiography, MR angiography or CT angiography) from which a diagnosis could be ascertained (on-site), a committee of three experienced physicians (Accuracy Review Committee) unaffiliated with the study sites 45

P S Sidhu, P Allan, F Cattin et al

compared the diagnosis obtained by the off-site assessments of Doppler ultrasound images with the diagnosis obtained with the reference modality.

Diagnostic quality Duration of clinically useful signal enhancement, defined as the time from appearance until disappearance of a microbubble contrast effect of sufficient intensity to be diagnostically or clinically useful, was assessed and documented by each off-site observer subjectively during review of the video recordings of the individual examinations. Incidence and duration of artefactual microbubble contrast effects (shadowing, blooming and saturation effects) were assessed following each dose of SonoVue. Artefacts were defined as follows: a shadowing effect appeared as an obscured image and/or Doppler spectrum, blooming appeared as the presence of colour in an area without flow, while a saturation effect appeared as a noisy Doppler spectrum with artificially high velocities [17]. Duration was evaluated from the actual time of appearance to the disappearance of shadowing and/or blooming and/or saturation effects. Each of the artefacts was evaluated at their maximal effect in accordance with the following three-point scale: 05no artefactual effect; 15artefactual effect not compromising the image analysis; 25artefactual effect compromising the image analysis. Assessment of inconclusive Doppler examinations was performed on each baseline or post-injection video clip where off-site readers had to assess if a diagnosis was possible or not and, in patients where it was possible, make a diagnosis based on a pre-defined check list.

Diagnostic accuracy Diagnostic accuracy was assessed for baseline and for the clinically recommended dose of SonoVue only (2.4 ml). Assessment of agreement was carried out by an Accuracy Review Committee, based on a comparison of the diagnosis recorded by each of the off-site blinded readers from the Doppler ultrasound investigations with the diagnosis from the reference imaging modality. The following four-point scale was used: 15full agreement; 25basic agreement (differences in details but leading to the same diagnostic conclusion); 35partial agreement (differences in details possibly leading to a different diagnostic conclusion); 45disagreement.

Sensitivity and specificity For the evaluation of the diagnostic performance of SonoVue CE-US in terms of sensitivity and specificity, it was necessary to further define agreement with the reference modality in terms of detection/exclusion (presence/absence) of particular lesions in the investigated vessels for each study patient. An independent experienced radiologist, not previously involved in these studies, was asked to classify the Doppler ultrasound off-site diagnoses and the reference modality diagnoses according to the following predetermined list of possible diagnoses for the designated vessel of interest: (1) no abnormality; 46

(2) abnormality present: (a) stenosis . 50% or occlusion; (b) atheromatous plaque; (c) arteriovenous malformation; (d) aneurysm; (e) vessel displacement/compression due to extrinsic space-occupying mass; (f) collaterals or collateralization of normal vessels; (g) arterial wall dissection; (h) other. Sensitivity was defined as the proportion of patients with a matching abnormality in the vessel of interest using Doppler ultrasound and patients with an abnormality in the vessel of interest using the reference standard. Specificity was defined as the proportion of patients with no abnormality in the vessel of interest using Doppler ultrasound and patients with no abnormality in the vessel of interest using the reference standard.

Statistical methods Demography Demographics and other baseline characteristics were summarized using descriptive statistics.

Efficacy analysis An analysis of variance (ANOVA) using ranked durations of clinically useful signal enhancement was performed to investigate overall differences between doses. Summary statistics, frequency distributions and crosstabulations were elaborated for efficacy parameters, but no formal statistical analyses were performed. For the purposes of statistical summaries from the assessment by the Accuracy Review Committee, these data were further categorized as follows: agreement denoted full agreement or basic agreement, and disagreement denoted partial agreement or disagreement. Individual study results for diagnostic accuracy were analysed using McNemar’s test of association between baseline and post-dose in the proportion of patients for whom agreement was recorded. For all analyses, a two-sided p-value was used to test for significance.

Results Eighty-two patients (study A, n543; study B, n539), 49 male and 33 female subjects, median age of 71 years (range 41–87 years), with suspected extracranial carotid artery or peripheral vascular disease were recruited and received SonoVue for the assessment of diagnostic quality parameters. The diagnostic accuracy assessment was performed on 59 patients where final diagnosis made from a reference imaging modality (conventional angiography/CT angiography n558 and MR angiography n51) was available for assessment by the Accuracy Review Committee. At Doppler examination 32 of 59 were found positive and 27 of 59 were negative for the presence of pathology. Sensitivity and specificity were calculated in the subset of patients (n546) where the reference standard with the pathology or no pathology in the vessel of interest was available for assessment (n522 and 24 patients, respectively, for the two studies). The British Journal of Radiology, January 2006


Diagnostic quality

Assessment of inconclusive Doppler ultrasound examinations

Duration of signal enhancement A statistically significant dose response was observed in the duration of clinically useful signal enhancement with a significant increase in the median duration across the doses (p,0.001 for 3 readers and p50.002 for 1 reader). At the highest SonoVue dose of 2.4 ml, the average median duration of clinically useful signal enhancement was of 3.9 min, range 0.0–14.3 (Table 1). The pair of off-site reader assessments for each patient was combined by calculating the average duration of clinically useful signal enhancement.

Despite the limitations of the methodology used in this study, the results of the statistical analysis performed showed that, at the dose 2.4 ml, which is recommended for Doppler ultrasound of macrovasculature, SonoVue markedly decreased the number of baseline inconclusive Doppler ultrasound examinations (rated as ‘‘no diagnosis possible’’). Considering the entire population, the percentage decreased from 40.7% to 7.4% (decrease533.3%) while, in the patient population with a reference gold standard control, the percentage decreased from 45.8% to 5.1% (decrease540.7%) (Figure 2).

Assessment of diagnostic accuracy

Artefactual effects Due to methodology which did not permit gain adjustment, a dose response was observed in the incidence and duration of artefactual contrast effects, with median values increasing up to a maximum of 3.6 min at the 2.4 ml dose. The most common artefactual effects were blooming in colour or power Doppler ultrasound (up to 92.9% with the 2.4 ml dose) followed by a saturation effect on spectral Doppler ultrasound. Both of these artefacts are related to the increase in Doppler signal intensity caused by the microbubble contrast agent (Figure 1). Shadowing was not reported to be a significant microbubble contrast artefactual effect by any of the four off-site readers.

The percentage of agreement between diagnosis from Doppler ultrasound investigations and diagnosis from the reference imaging modality in the entire patient population increased from 30.7% at baseline to 68.9% post-contrast (Table 2).

Table 1. Duration of clinically useful signal enhancement in the two studies combined (A and B) from the off-site evaluation SonoVue dose n582

Baseline

0.3 ml

0.6 ml

1.2 ml

2.4 ml

Median (min) Range

0.00 0.0– 2.0

2.5 0.0–9.2

2.9 0.0–10.1

3.4 0.0–20.5

3.9 0.0–14.3

(a)

Figure 2. Bar chart diagram demonstrating the alteration in the number of inconclusive off-site Doppler ultrasound assessments at the dose of 2.4 ml of SonoVue (entire population: all patients in the study, n582; population with ref. std: patients with the reference gold standard, n546).

(b)

Figure 1. Illustration of the ‘‘blooming’’ artefact. (a) Following the administration of SonoVue 1.2 ml, extensive blooming at 36 s obscures the arterial anatomy precluding diagnostic interpretation. (b) Without adjustment of the ultrasound machine imaging parameters, at 63 s blooming has subsided and there is better delineation of the arterial anatomy. The British Journal of Radiology, January 2006

47

P S Sidhu, P Allan, F Cattin et al

The change in agreement rates from baseline was statistically significant for three of the four off-site readers (range p,0.05–0.001). Furthermore, considering the subset of patients whose investigation was diagnostic and the reference standard imaging modality available, agreement between the Doppler ultrasound diagnosis and the diagnosis from the reference imaging modality further increased to 72.3% after microbubble contrast administration (Figure 3). Diagnostic performance, in terms of sensitivity and specificity, was assessed in the subset of patients (n546) with an abnormality or no abnormality in the vessel of interest on the available reference gold standard examinations. In study A, in the eight patients with an abnormality in the carotid/peripheral vessel of interest on the reference standard modality, the sensitivity increased from 13% precontrast to 75% at the 2.4 ml dose for reader 1 and from 50% to 75% for reader 2 (Table 3). In the 16 patients with no abnormality on the reference modality, specificity increased from a pre-contrast value of 0% to 85 % at 2.4 ml for reader 1 and from 31% to 75% for reader 2. In study B, in the nine patients with an abnormality on the reference modality, the sensitivity increased from 89% pre-contrast to 100% at the 2.4 ml dose for reader 3 and increased from 22% to 67% for reader 4. In the 13 patients with no abnormality on the reference modality, specificity increased from a precontrast value of 23% to 46% at 2.4 ml for reader 3 and from 8% to 85% for reader 4.

Discussion Failure to obtain a diagnostic colour Doppler ultrasound examination of the extracranial carotid and peripheral arteries is typically a consequence of patient factors. Rather than abandon the colour Doppler ultrasound examination instituting another examination, with the implications of higher cost, the introduction of a microbubble contrast agent would enable the examiner to attempt to establish a conclusive diagnosis and reduce the amount of time necessary to perform a peripheral arterial examination. In the present study, a significant dose effect was observed for the duration of clinically useful signal enhancement for all four off-site readers; the average median duration of useful enhancement was of 3.91 min for the 2.4 ml dose. Due to a conservative approach in the study design, the incidence and duration of artefactual microbubble contrast effects also tended to increase with increasing dose. These artefacts can normally be limited by reducing the effective sensitivity of the system, by decreasing the colour or power and spectral Doppler ultrasound gains. Indeed, in order to maximize Doppler quality, gain settings for both colour or power and spectral Doppler ultrasound should be continuously adjusted as enhancement returns to baseline. However, in this study Table 2. Diagnostic accuracy. Percentage agreement between diagnosis from Doppler ultrasound investigations (both unenhanced and SonoVue microbubble contrast-enhanced) and diagnosis from the reference standard in the entire population

Agreement with gold standard

48

Unenhanced

SonoVue 2.4 ml

30.7%

68.9%

the protocol required that the level of the gain for both colour or power and spectral Doppler be set before the first injection of microbubble contrast and could not be modified after that. With gain adjustment, the duration of artefacts would have been greatly reduced if not completely eliminated and, as a result, the duration of clinically useful signal enhancement would have been increased. Moreover, now there is a tendency to use infusions rather than bolus injection of microbubble contrast during the investigation of vascular disease. This has been demonstrated to further improve the duration of useful enhancement and reduce artefactual effects in the extracranial carotid and peripheral arteries [18], in transcranial Doppler ultrasound [19] and in the portal vein [20, 21]. Administration of microbubble contrast resulted in an increase in agreement between colour Doppler ultrasound diagnosis and diagnosis from a reference imaging modality. This is of importance where full reliance can be placed on the results of a colour Doppler ultrasound examination in order to bypass angiography prior to any surgical procedure, particularly in carotid end-arterectomy surgery [5] where there is a small but significant morbidity attached to diagnostic angiography [22]. The effect of introducing a microbubble contrast agent, the ‘‘Doppler rescue’’ effect, has been successful in the imaging of the renal arteries [23], the hepatic artery in the liver transplant patient [24, 25] and the portal vein [20, 26, 27]. In the assessment of renal artery disease, using a galactose based microbubble contrast agent (LevovistTM; Schering AG, Berlin, Germany), visualization of the renal arteries improved from 65.7% to 78.3% (p,0.01) following the administration of microbubble contrast [23]. The use of SonoVue in the present study improved the ability of all the off-site readers, presented with a minimum amount of information, to make a confident interpretation of the underlying vascular disorder on the CE-US examination. On the baseline ultrasound examinations, a correct diagnosis confirmed by the standard of reference was achieved in 29.7% of studies, improving to 67.6% with the 2.4 ml dose of SonoVue. Moreover, if accuracy is evaluated in the subset of patient population with diagnostic examinations, a further increase in the percentage of agreement with reference gold standard is observed after SonoVue (72.3%) compared with unenhanced examinations. There was an overall improvement in sensitivity and specificity for all the off-site readers. The ability shown by SonoVue to improve the diagnostic information from a recorded ultrasound examination, having knowledge of the vessel of interest only, is remarkable since in the clinical practice the nature of any ultrasound examination is one of examiner-patient interaction, where the physician is allowed to develop an overall concept of the diagnosis. This would suggest that the use of a microbubble contrast agent as part of an onsite ultrasound assessment would improve the diagnostic ability to an even greater degree than what appeared under the investigational conditions of the present study. One limitation of the current study is the level of sophistication of the ultrasound machines used. When this multicentre study was commenced, each centre was equipped with a ‘‘top-of-the-range’’ ultrasound machine, but during the course of the study introduction on the market of newer machines with digital capability, more The British Journal of Radiology, January 2006


(a)

(b)

(d)

(c) Figure 3. (a) Baseline unenhanced colour and spectral Doppler ultrasound examination of a patient right lower limb. Insufficient information for a firm conclusion about patency of the anterior tibial artery. (b) Following the administration of 2.4 ml of SonoVue, clear depiction of a patent anterior tibial artery is seen (long arrow) with a large collateral artery seen in a superior position (short arrow). (c) Spectral Doppler contrast ultrasound confirms a monophasic abnormal arterial trace. (d) Corresponding arteriogram confirms the patent anterior tibial artery (long arrow) and the collateral artery (short arrow). Collateral arteries have formed around an occluded popliteal artery. (Courtesy of Dr J Langholz).

sensitive to blood flow, was seen to improve vascular ultrasound diagnosis. Nevertheless, even with the improved capabilities of these newer ultrasound machines, problem patients will still exist and the need for ‘‘Doppler-rescue’’ with The British Journal of Radiology, January 2006

microbubble contrast will still be advantageous to reduce the need for further imaging. The quality and standard of the on-site colour Doppler ultrasound examinations were dependent on the experience of the examining 49

P S Sidhu, P Allan, F Cattin et al Table 3. Sensitivity and specificity of Doppler ultrasound investigations based on diagnosis from the reference gold standard Study A (n524)

Baseline

SonoVue 2.4 ml

Sensitivity (n58) CI Specificity (n516) CI

(1/8) 0.13 (20.10, 0.36) 0.00 (0.00, 0.00)

(6/8) 0.75 (0.45, 1.05) (14/16) 0.88 (0.72, 1.04)

Sensitivity (n58) CI Specificity (n516) CI Study B (n522)

(4/8) 0.50 (0.15, 0.85) (5/16) 0.31 (0.083, 0.54) Baseline

(6/8) 0.75 (0.45, 1.05) (12/16) 0.75 (0.54, 0.96) SonoVue 2.4 ml


(8/9) 0.89 (0.68, 1.09) (3/13) 0.23 (0.0012, 0.46)

(9/9) 1.00 (1.00, 1.00) (6/13) 0.46 (0.19, 0.73)


(2/9) 0.22 (20.051, 0.49) (1/13) 0.08 (20.067, 0.23)

(6/9) 0.67 (0.36, 0.98) (11/13) 0.85 (0.65, 1.04)

Acknowledgments We wish to thank Dr Franca Heiman for her statistical assistance.

Reader 1

Reader 2

Reader 3

Reader 4

CI, 95% confidence interval.

sonographer, with likely variation between the centres involved in the study. Accepting centres with an established reputation for vascular ultrasound and ensuring that only the most experienced sonographers performed the examination minimized this variation. Not all of the patients had an acceptable standard of reference imaging examination, but even in the smaller number where this was available, addition of microbubble contrast improved the diagnostic capability of the colour Doppler ultrasound examination. The variation in the assessment of the baseline colour Doppler ultrasound examinations by the off-site investigators highlights the difficulties of ultrasound interpretation and the subjective nature of conclusions reached. However, addition of microbubble contrast, although not completely eliminating this subjectivity, dramatically improved the confidence in interpretation allowing the off-site investigator to establish the correct diagnosis more consistently. In conclusion, at the dose of 2.4 ml of SonoVue, the duration of useful enhancement achieved, which may be further extended by adjustment of the ultrasound machine settings, allowed a sufficiently prolonged period to establish a definitive diagnosis avoiding further imaging. Further studies to evaluate the potential use of infusion methods of SonoVue administration are needed; these may have advantages over bolus methods of administration. The use of SonoVue allowed an improvement in diagnostic accuracy to be achieved in comparison with an accepted reference examination. The administration of 2.4 ml dose significantly produces an overall improvement in terms of diagnostic performance. Microbubble ultrasound contrast represents the next stage of development, following on from the introduction of duplex Doppler and colour Doppler ultrasound, in the improving the overall diagnostic capability of ultrasound in the vascular system. 50

References 1. Kohler TR, Nance DR, Cramer MM, Vandenburghe N, Strandness DEJ. Duplex scanning for diagnosis of aortailiac and femoropopliteal disease: a prospective study. Circulation 1987;76:1074–80. 2. Whelan JF, Barry MH, Moir JD. Color flow Doppler ultrasonography: comparison with peripheral arteriography for the investigation of peripheral arterial disease. J Clin Ultrasound 1992;20:369–74. 3. Sidhu PS, Allan PL. The extended role of carotid artery ultrasound. Clin Radiol 1997;52:643–53. 4. Sidhu PS, Allan PL. Ultrasound assessment of internal carotid artery stenosis. Clin Radiol 1997;52:654–8. 5. Khaw KT. Does carotid duplex imaging render angiography redundant before carotid endarterectomy? Br J Radiol 1997;70:235–8. 6. Garrard CL. Cost savings associated with the non-routine use of carotid angiography. Am J Surg 2000;174:650–3. 7. Koelemay MJ, Legemate DA, de Vos H, van Gurp AJ, Balm R, Reekers JA, Jacobs MJ. Duplex scanning allows selective use of arteriography in the management of patients with severe lower leg arterial disease. J Vasc Surg 2001;34:661–7. 8. van der Zaag ES, Legemate DA, Nguyen T, Balm R, Jacobs MJ. Aortoiliac reconstructive surgery based upon the results of duplex scanning. Eur J Vasc Endovasc Surg 1998;16:383–9. 9. Rankin SC. CT angiography. Eur Radiol 1999;9:297–310. 10. Dupuy DE, Boland GW. Non-invasive angiography with magnetic resonance imaging and computed tomography. Imaging 1995;7:134–47. 11. Koelemay MJ, Lijmer JG, Stoker J, Legemate DA, Bossuyt PM. Magnetic resonance angiography for the evaluation of lower extremity arterial disease: a meta-analysis. JAMA 2001;285:1338–45. 12. Nelemans PJ, Leiner T, de Vet HC, van Engelshoven JM. Peripheral arterial disease: meta-analysis of the diagnostic performance of MR angiography. Radiology 2000;217:105–14. 13. Schneider M, Arditi M, Barrau MB, et al. BR 1: a new ultrasonographic contrast agent based on sulfur hexafluoridefilled microbubbles. Invest Radiol 1995;30:451–7. 14. Schneider M. Characteristics of SonoVue. Echocardiography 1999;16:743–6. 15. Schneider M. SonoVue, a new ultrasound contrast agent. Eur Radiol 1999;9:S347–8. 16. Morel DR, Schwieger I, Hohn L, et al. Human pharmacokinetics and safety evaluation of SonoVue, a new contrast agent for ultrasound imaging. Invest Radiol 2000;35:80–5. 17. Forsberg F, Liu JB, Burns PN, Merton DA, Goldberg BB. Artifacts in ultrasonic contrast agents studies. J Ultrasound Med 1994;13:357–65. 18. Albrecht T, Urbank A, Mahler M, et al. Prolongation and optimization of Doppler enhancement with a microbubble US contrast agent by using continuous infusion: preliminary experience. Radiology 1998;207:339–47. 19. Goertler M, Kross R, Baeumer M, et al. Diagnostic impact and prognostic relevance of early contrast-enhanced transcranial color-coded duplex sonography in acute stroke. Stroke 1998;29:955–62. 20. Schiedermaier P, Layer G, Sauerbruch T. Impact of the continuous infusion of Levovist on color Doppler sonography in portal hypertension. AJR Am J Roentgenol 2002;178:61–5.


Diagnostic efficacy of SonoVue 21. Sidhu PS, Sellars ME, Heneghan M, Blomley MJK, Bauer A. Visualization of the portal vein in normal subjects and patients with cirrhosis: comparison of a bolus versus different infusion injections of ultrasound contrast. Eur Radiol 2000;10:119. 22. Hankey GJ, Warlow CP, Sellar RJ. Cerebral angiographic risk in mild cerebrovascular disease. Stroke 1990;21:209–22. 23. Claudon M, Plouin PF, Baxter GM, Devos DM. Renal arteries in patients at risk of renal arterial stenosis: multicentre evaluation of the Echo-enhancer SH U 508A at color and spectral Doppler US. Radiology 2000;214:737–46. 24. Sidhu PS, Shaw AS, Ellis SM, Karani JB, Ryan SM. Microbubble ultrasound contrast in the assessment of hepatic artery patency following liver transplantation: role in reducing frequency of hepatic artery arteriography. Eur Radiol 2004;14:21–30.


25. Sidhu PS, Ellis SM, Karani JB, Ryan SM. Hepatic artery stenosis following transplantation: significance of the tardus parvus waveform and the role of microbubble contrast media in the detection of a focal stenosis. Clin Radiol 2002;57:789– 99. 26. Gebel M, Caselitz M, Bowen-Davies PE, Weber S. A multicenter, prospective, open label, randomized, controlled phase IIIb study of SH U 508A (Levovist) for Doppler signal enhancement in the portal vascular system. Ultraschall in der Medizin 1998;19:148–56. 27. Marshall MM, Beese RC, Muiesan P, Sarma DI, O’Grady J, Sidhu PS. Assessment of portal venous system patency in the liver transplant candidate: a prospective study comparing ultrasound, microbubble contrast enhanced colour Doppler ultrasound, with arteriography and surgery. Clin Radiol 2002;57:377–83.

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E


Lymphoepithelioma-like carcinoma of salivary glands: treatment results and failure patterns 1

C-Y HSIUNG, MD, 2C-C HUANG, MD, 1C-J WANG, MD, 1E-Y HUANG, MD and 2H-Y HUANG, MD

Departments of 1Radiation Oncology and 2Pathology, Chang Gung Memorial Hospital-Kaohsiung, Taiwan, R.O.C.

Abstract. The purpose of this study was to evaluate the treatment results and failure patterns of lymphoepithelioma-like carcinoma (LELC) of salivary glands. From June 1987 to May 2001, nine patients with LELC of salivary glands were treated at our hospital. One patient was excluded due to the loss of clinical follow-up after surgery. For the remaining eight patients, the primary tumour sites were parotid glands (4 patients), submandibular glands (3), and the minor salivary glands in right cheek (1), respectively. Seven patients underwent surgical treatment and post-operative radiotherapy, while the other one patient was treated with surgery only. The total radiation dose to the salivary tumour bed ranged from 39.6 Gy to 67.6 Gy (mean dose: 58.3 Gy and median dose: 59 Gy). The treatment results and failure patterns were analysed. The survival time ranged from 21.4 months to 145.2 months (mean: 69.1 months, median: 54.5 months). At the end of follow-up, six patients were still alive and two died. One patient died of distant metastases 21.5 months after the surgical treatment of LELC. The other case died of intercurrent disease (pontine haemorrhage) 53 months after surgery. No patient had local or regional failure after the treatments. Distant failure was noted in two patients. The patients with LELC of salivary glands were shown to have favourable prognoses. No local or regional failure was noted. However, distant failure developed in two patients. The risk of distant metastasis should be carefully monitored, especially for those patients with more advanced neck node involvement.

Lymphoepithelioma [1] consisted of poorly differentiated cells with large nuclei and nucleoli within the lymphoid stroma. Lymphoepithelioma occurs mainly in the nasopharynx [2, 3]. Also, lymphoepithelioma-like carcinoma (LELC) has been found in salivary glands [4– 6]. Because LELC is a rare histological type of cancer of salivary glands [7, 8], the clinical data concerning LELC of salivary glands is inadequate compared with other common histological types. Also, the clinical course and prognosis of this disease after the treatments have not been thoroughly studied in the medical literature. As a result, a retrospective study based on our patient database was undertaken to analyse the treatment results and failure patterns of LELC of salivary glands.

Patients and methods From June 1987 to May 2001, nine patients with LELC of salivary glands were treated at our hospital. One patient was excluded due to the loss of clinical follow-up. The remaining eight patients are followed up regularly after the treatments and included in the current study. The general characteristics of these patients were shown in Table 1. Three out of eight patients were male and five were female. The primary tumour sites were parotid glands (4 patients), submandibular glands (3), and the minor salivary glands in right cheek (1), respectively. These patients with LELC were staged according to TNM classification of the American Joint Committee on Cancer [9] (Table 1). The treatment data of these patients are presented in Table 2. All these eight patients underwent the excision of Received 5 January 2005 and accepted 7 June 2005. Address correspondence to: Hsuan-Ying Huang, Department of Pathology, Chang Gung Memorial Hospital-Kaohsiung, 123, Ta-Pei Road, Niao Sung Hsian, Kaohsiung Hsien, Taiwan, R.O.C.

52

primary salivary gland tumours. The dissection of enlarged neck lymph nodes was also performed for the five patients (patients 3, 4, 5, 6, and 8 in Table 2) with neck node metastases noted by physical examination or CT scans. After surgery, seven cases received post-operative radiotherapy with a 60Co machine or 6–10 MV linear accelerator. Six (patients 1, 3, 4, 5, 6, and 8 in Table 2) out of these seven patients were irradiated with two bilateral portals covering the salivary tumour bed and upper neck and an anterior–posterior portal covering the bilateral lower neck. The remaining one patient (patient 2 in Table 2) received small-field radiotherapy covering only salivary tumour bed without elective nodal irradiation to bilateral low neck. In the seven patients treated with post-operative radiotherapy, the total radiation dose to the salivary tumour bed ranged from 39.6 Gy to 67.6 Gy (mean dose: 58.3 Gy and median dose: 59 Gy). The dose to spinal cord was no more than 45 Gy. For the six patients receiving elective nodal irradiation to bilateral low neck, the low-neck dose ranged from 34.2 Gy to 45 Gy (Table 2). After the treatments, all the patients were followed regularly at the clinics. The treatment results and failure patterns were retrospectively reviewed. The survival time was measured from the date of the first surgical treatment to the date of last follow-up or death. The survival curves were calculated by the Kaplan-Meier product-limit method [10]. Local failure was defined as tumour recurrence in the salivary tumour bed. Regional failure was defined as tumour recurrence in the head and neck outside the salivary tumour bed.

Results The histology of LELC of one patient is shown in Figure 1. The treatment results and failure patterns are summarized in Table 3. The survival time ranged from The British Journal of Radiology, January 2006

Lymphoepithelioma-like carcinoma of salivary glands Table 1. The general characteristics of the eight patients with lymphoepithelioma-like carcinoma (LELC) of salivary glands

Patient Patient Patient Patient Patient Patient Patient Patient

1 2 3 4 5 6 7 8

Age (years)

Sex

Primary site

Stage [9]

42 50 40 39 43 40 42 46

Male Female Male Female Female Male Female Female

Right submandibular gland Minor salivary gland in right buccal area Left submandibular gland Right parotid gland Right parotid gland Left submandibular gland Right parotid gland Left parotid gland

T3 T1 T2 T4 T3 T3 T2 T3

21.4 months to 145.2 months (mean: 69.1 months, median: 54.5 months). The survival curve of these patients is shown in Figure 2. At the last follow-up, six patients were still alive and two had died. One patient died of distant metastases 21.5 months after the surgical treatment of LELC (patient 6 in Table 3). The other case died of intercurrent disease (pontine haemorrhage) 53 months after surgery (patient 1 in Table 3). No patient had local or regional failure after the treatments. However, distant metastases were noted in two patients (patients 6 and 8 in Table 3). The interval between surgery and distant failure was 6.3 months and 6.5 months for patient 6 and 8, respectively. After the occurrence of distant metastases, these two patients received chemotherapy with CDDP and 5-FU. At last follow-up, five patients were alive without cancer, one was alive with distant metastases, another one had died of distant metastases, and the remaining one had died of intercurrent disease (Table 3). During radiotherapy, oral mucositis and skin reaction over radiation field were experienced in all the seven patients irradiated. The major long-term complications after the treatments were xerostomia (8 patients), neck fibrosis (6 patients), and facial palsy (3 patients). The complication of facial palsy was due to tumour encasement of facial nerve and the surgical treatment.

Discussion Lymphoepithelioma in nasopharynx is known as a radiosensitive tumour and radiotherapy is the standard treatment for nasopharyngeal lymphoepithelioma [2, 3]. Non-nasopharyngeal lymphoepithelioma of the head and neck is also reported to be radiosensitive with high rates of locoregional tumour control [5]. In the study of salivary gland carcinoma by Teo et al [6], seven patients had LELC from the parotid glands and only two of them experienced locoregional relapses; one had isolated regional relapse

N0 M0 N0 M0 N2b M0 N1 M0 N2b M0 N2b M0 N0 M0 N2b M0

Figure 1. The histology of lymphoepithelioma-like carcinoma (LELC) of one patient is shown here. Microscopically, diffuse lymphoid infiltration is noted around the atrophic acini of salivary gland and occasionally forms lymphoid follicles. Islands of neoplastic epithelial cells bearing pleomorphic, vesicular nuclei and indistinct cell border are present within the lymphoid tissue.

outside the post-operative radiation field 6 years after treatments and the other had in-field failure in the parotid tumour bed 3.5 years after total parotidectomy and postoperative radiation (50 Gy). In the current study, seven of these eight patients with LELC of salivary glands received surgery and post-operative radiotherapy and the other one was treated with surgery only. No local or regional failure was noted. From the results of this study and the above literature [5, 6], surgery and post-operative radiotherapy may be the appropriate treatment combination with satisfactory locoregional control for patients with LELC

Table 2. The treatment data of the eight patients with lymphoepithelioma-like carcinoma (LELC) of salivary glands Treatments


1 2 3 4 5 6 7 8

Operation Operation Operation Operation Operation Operation Operation Operation

& radiotherapy & radiotherapy & radiotherapy & radiotherapy & radiotherapy & radiotherapy alone & radiotherapy


Radiation dose (Gy) Salivary tumour bed

Bilateral low neck

67.6 65.2 57.6 39.6 64.8 59

45 0 45 39.6 45 45

54.2

34.2

53

C-Y Hsuing, C-C Huang, C-J Wang et al Table 3. The treatment results and failure patterns of the eight patients with lymphoepithelioma-like carcinoma (LELC) of salivary glands


1 2 3 4 5 6 7 8

Survival months

Distant metastases

Status at last follow-up

53 116.5 56 21.4 103.6 21.5 145.2 35.4

No No No No No Lung, bone & liver No Lung, bone & liver

Died of intercurrent disease Alive without cancer Alive without cancer Alive without cancer Alive without cancer Died of distant metastases Alive without cancer Alive with distant metastases

Figure 2. The survival curve of the eight patients with lymphoepithelioma-like carcinoma (LELC) of salivary glands.

of salivary glands. In the future, further study with more patients is needed to find the appropriate radiation field and radiation dose for LELC of salivary glands. Distant metastases to lung, bone, and liver were noted in two patients (patient 6 and 8, Table 3). The duration from the date of operation to distant metastases was 6.4 months and 6.6 months for patients 6 and 8, respectively. Among these eight patients, there were four patients with N0 or N1 stage (Table 1), and none of them experienced distant metastases. The other four patients were all staged as N2b and two of them had distant metastases after the treatments. From this finding, the neck node status might be associated with the risk of distant metastases. In the study of non-nasopharyngeal lymphoepithelioma of the head and neck [5], the main cause of treatment failure was distant metastasis, which occurred more frequently in patients with lymph node involvement. As a result, the risk of distant metastasis should not be overlooked for those patients with more advanced neck node involvement.

Conclusion In the current study, patients with LELC of salivary glands were shown to have favourable prognoses. No local or regional failure occurred in these patients. However, 54

distant failure developed in two patients. The risk of distant metastasis should be carefully monitored, especially for those patients with more advanced neck node involvement.

Acknowledgments The authors thank Yu-Ling Wu, M.S. for the kind assistance with manuscript preparation.

References 1. Schmincke A. Uber lymphoepitheliale Geschevulste. Beitr Pathol Anat 1921;68:161. 2. Perez CA. Nasopharynx. In: Perez CA, Brady LW, editors. Principles and practice of radiation oncology. 2nd edn. Philadelphia, PA: JB Lippincott; 1992:617–43. 3. Moss WT. The nasopharynx. In: Cox JD, editor. Moss’ radiation oncology: rationale, technique, results. 7th edn. St. Louis, MO: Mosby, 1994:149–68. 4. Cleary KR, Batsakis JG. Undifferentiated carcinoma with lymphoid stroma of the major salivary glands. Ann Otol Rhinol Laryngol 1990;99:236–8. 5. Dubey P, Ha CS, Ang KK, El-Naggar AK, Knapp C, Byers RM, et al. Nonnasopharyngeal lymphoepithelioma of the head and neck. Cancer 1998;82:1556–62.


Lymphoepithelioma-like carcinoma of salivary glands 6. Teo PM, Chan AT, Lee WY, Leung SF, Chan ES, Mok CO. Failure patterns and factors affecting prognosis of salivary gland carcinoma: retrospective study. Hong Kong Med J 2000;6:29–36. 7. Simpson JR. Salivary glands. In: Perez CA, Brady LW, editors. Principles and practice of radiation oncology. 2nd edn. Philadelphia, PA: JB Lippincott, 1992:657–71. 8. Moss WT. The salivary glands. In: Cox JD, editor. Moss’ radiation oncology: rationale, technique, results. 7th edn. St. Louis, MO: Mosby, 1994:121–31.


9. Major salivary glands (parotid, submandibular, and sublingual). In: American Joint Committee on Cancer: AJCC Cancer Staging Manual. Philadelphia, PA: Lippincott-Raven Publishers, 5th edn, 1997:53–8. 10. Kaplan EL, Meier P. Nonparametric estimation from incomplete observations. J Am Stat Assoc 1958;53:457–81.

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E


Comparison of patient doses in 256-slice CT and 16-slice CT scanners 1,2

S MORI, MS, RT, MPR, 1M ENDO, PhD, MPH, 1K NISHIZAWA, PhD, MPH, 2K MURASE, PhD, MPH, H FUJIWARA, PhD and 3S TANADA, MD

2 1

Department of Medical Physics, National Institute of Radiological Sciences, Chiba 263-8555, Japan, 2School of Allied Health Sciences, Faculty of Medicine, Osaka University, Osaka 565-0871, Japan and 3Department of Medical Imaging, National Institute of Radiological Sciences, Chiba 263-8555, Japan

Abstract. The 256-slice CT-scanner has been developed at the National Institute of Radiological Sciences. Nominal beam width was 128 mm in the longitudinal direction. When scanning continuously at the same position to obtain four-dimensional (4D) images, the effective dose is increased in proportion to the scan time. Our purpose in this work was to measure the dose for the 256-slice CT, to compare it with that of the 16-slice CT-scanner, and to make a preliminary assessment of dose for dynamic 3D imaging (volumetric cine imaging). Our group reported previously that the phantom length and integration range for dosimetry needed to be at least 300 mm to represent more than 90% of the line integral dose with the beam width between 20 mm and 138 mm. In order to obtain good estimates of the dose, we measured the line-integral dose over a 300 mm range in PMMA (polymethylmethacrylate) phantoms of 160 mm or 320 mm diameter and 300 mm length. Doses for both CT systems were compared for a clinical protocol. The results showed that the 256-slice CT generates a smaller dose than the 16-slice CT in all examinations. For volumetric cine imaging, we found an acceptable scan time would be 6 s to 11 s, depending on examinations, if dose must be limited to the same values as routine examinations with a conventional multidetector CT. Finally, we discussed the studies necessary to make full use of volumetric cine imaging.

In 2001 the introduction of a 16-slice CT-scanner raised some new topics in CT technology development. 16-slice CT allows applications of three-dimensional (3D) images in clinical fields such as diagnosis, surgical simulation, planning of radiation therapy and monitoring of interventional therapy. However, it is still difficult to take dynamic 3D images of moving organs such as the heart or lung to enlarge the application fields. In order to take these images, we have developed a prototype 256-slice CT at NIRS (National Institute of Radiological Sciences) which employs continuous rotations of a cone-beam [1]. Clinical applications of CT techniques have continued to increase the dose to patients during recent decades, as CT examinations have come to provide higher quality X-ray imaging with substantial benefits in clinical diagnosis [2]. Notwithstanding the potential benefits to the healthcare of patients using CT, the fundamental concern in radiological protection is the optimization of radiation exposure. The maximum nominal beam width of the 256-slice CT is 128 mm and is four times larger than the thirdgeneration 16-slice CT-scanner (Toshiba Aquilion; Toshiba Medical Systems, Japan). A wider beam width is more efficient for imaging in a wider coverage. However, doses to patients with 256-slice CT are of considerable concern if it is to be used for obtaining dynamic 3D images (volumetric cine images). When scanning continuously at the same position, the effective dose is increased in Received 6 August 2004 and in revised form 8 April 2005, accepted 13 June 2005. Address correspondence to: Shinichiro Mori, 4-9-1 Anagawa, Inageku, Chiba-shi, Chiba, 263-8555, Japan.

56

proportion to the scan time and a wider coverage brings larger doses to patients. Therefore, it is very important to assess the dose of the 256-slice CT before volumetric cine imaging for patients. This work was carried out to compare doses, including scattered radiation, of the 256-slice CT and 16-slice CT and to make a preliminary assessment of dose for volumetric cine imaging.

Materials and methods Acquisition systems of 256-slice CT and 16-slice CT scanners The prototype 256-slice CT-scanner uses a wide-area 2D detector designed on the basis of the present CT technology and is mounted on the gantry frame of a state-of-the-art CT-scanner (Figure 1) [3]. The number of elements is 912 channels6256 segments; element size is approximately 1 mm61 mm, corresponding to a 0.5 mm (transverse)60.5 mm (longitudinal) beam width at the centre of rotation. Gantry rotation time is 1.0 s. Data sampling rate is 900 views/s, and the dynamic range of the A/D converter is 16 bits. As shown in Appendix 1, the reconstructed regions are cylinders of 240 mm diameter and 102.4 mm length for the head scan and 320 mm diameter and 93.9 mm length for the body scan. The detector element consists of a scintillator and photodiode, which are the same as for the scintillator of multidetector CT (MDCT) (Toshiba Aquilion). Three wedge designs (large, small, and flat) on the 256-slice CT are intended to extend the conventional wedge designs of the thirdgeneration 16-slice CT-scanner (Toshiba Aquilion) in the The British Journal of Radiology, January 2006

Comparison of patient dose in multislice CT

(b)

(a)

Figure 1. (a) Front view of 256-slice CT-scanner. (b) A wide-area 2D detector is designed on the basis of the present CT technology and mounted on the gantry frame of the state-of-the-art CT-scanner.

longitudinal direction. The large and small wedges are shaped to compensate for the variable path length of the patient across the scan field of view (FOV). The small wedge is used for an object under 240 mm FOV, and the large wedge is used for over 240 mm FOV (e.g. chest and abdomen). The flat wedge is thicker at the centre than the other wedges. A Feldkamp-Davis-Kress (FDK) algorithm [4] is used for reconstruction. All further data processing and interpretation are done with a high-speed image processor with field programmable gate-array based-(FPGA) architecture. It takes less than 1 s to reconstruct volume data of a 51265126256 matrix. The 16-slice CT detector consists of 40 segments, which can be electronically grouped to provide different image slice configurations. The longitudinal FOV is 32 mm at the maximum. Other major components are the same as those of the 256-slice CT. In addition to the axial scan, the helical scan mode can be selected to cover volumes beyond the detector width.

Phantoms The length of the IEC-recommended dosimetry phantom [5] is at least 140 mm. This conventional phantom contains holes just large enough to accept the pencilshaped ionization chamber. For dose measurement in cone-beam CT, the length of the phantom should be longer, because of the wider scatter distribution. According to our previous results [6], the phantom length and integration range for dosimetry needed to be at least 300 mm to represent more than 90% of line integral dose with the beam width between 20 mm and 138 mm. Therefore, in the present study we used 300 mm long phantoms of PMMA (polymethylmethacrylate). The diameters of the phantoms are 160 mm for head and 320 mm for body examination. These phantoms were provided by joining unit cylinders 150 mm long. The details of the phantoms were described by Mori et al [6]. The British Journal of Radiology, January 2006

Detectors A pencil-shaped ionization chamber (CT-30; Oyogiken, Japan) of active length 300 mm was connected to a dosemeter (AE-132; Oyogiken, Japan) and used to measure dose. The dosemeter was calibrated (National Institute of Advanced Industrial Science and Technology, Japan) for the appropriate radiation qualities.

Clinical scan conditions We compared the doses of the 256-slice CT and the 16-slice CT for clinical scan conditions. These conditions were mainly derived from those recommended by the manufacture for the 16-slice CT. The X-ray tube current was set such that the effective mAs should be the same for both CTs, as given by (current)6(rotation time)/ (helical pitch) for the 16-slice CT and by (current)6(rotation time) for the 256-slice CT. For the 256-slice CT, slice collimation was 224 mm60.5 mm for the head, 128 mm61.0 mm for the pelvis, and 256 mm60.5 mm for other sites. For the 16-slice CT, the slice collimation was set to 16 mm61.0 mm for pelvis and 16 mm60.5 mm for other sites, helical pitch was 0.69 for the head, and 0.94 for other sites, because the scan conditions were chosen to obtain the same spatial resolution as for the 256-slice CT. The whole scan ranges were 93.9 mm for chest, 187.8 mm for abdomen, and 281.7 mm for pelvis. These scan ranges, except chest examination, were beyond the detector width of the 256-slice CT in the longitudinal direction, therefore they were set as multiples of 93.9 mm, the maximum longitudinal FOV of the 256-slice CT (Appendix 1). For the head examination, because the recommended value for the 16-slice CT was shorter than the maximum FOV of 256-slice CT, the FOV was adjusted to narrow the collimator width for the 256-slice CT. The clinical scan conditions thus obtained are summarized in Table 1. 57

S Mori, M Endo, K Nishizawa et al Table 1. Scan conditions for 256-slice CT and 16-slice CT-scanners Examination Scanner

Voltage Current Rotation Scan Beam collimation FOV Scan range Scan mode Helical pitch (kV) (mA) time (s) time (s) (mm6mm) (mm) (mm)

Head

120

Chest Abdomen Pelvis

256-slice CT 16-slice CT 256-slice CT 16-slice CT 256-slice CT 16-slice CT 256-slice CT 16-slice CT

120 120 120

326 300 160 300 213 400 213 400

1.0 0.75 1.0 0.5 1.0 0.5 1.0 0.5

1.0 17.0 1.0 8.3 1s62 14.5 1s63 11.4

22460.5 1660.5 25660.5 1660.5 25660.5 1660.5 12861.0 1661.0

240

90.0

320

93.9

320

187.8

320

281.7

Axial Helical Axial Helical Axial Helical Axial Helical

N/A 0.69 N/A 0.94 N/A 0.94 N/A 0.94

FOV, field of view.

Dose measurements The dose for both CT systems was measured with the 300 mm long pencil-shaped ionization chamber and 300 mm long phantoms (160 mm and 320 mm diameter) in one rotation scan. The measurement range in the longitudinal direction was 300 mm (z5¡150 mm). The phantom was placed on the patient table and its centre was aligned at the isocentre. The ionization chamber was inserted into either the central or one of the peripheral cavities of the phantom (other cavities were filled with PMMA rods). The exposure (expressed as Roentgens) was obtained with the ionization chamber dosemeter and converted to the values of absorbed dose to air measured in PMMA with the f-factor 0.898 cGy R21.

Dose assessment The dose was assessed using the dose profile integral (DPI) over 300 mm (z5¡150 mm) (Appendix 2), which was given by the output of the pencil ionization chamber of 300 mm length [6]. The weighted average of DPI at the centre and peripheries of the phantoms is given by 1 2 DPIw ~ DPIc z DPIp 3 3

ð1Þ

if we assume a linear decrease (or increase) of DPI in the radial direction, where DPIc is the DPI at the centre and DPIp the average DPI on the peripheries.

Clinical image quality

was carried out as follows: (i) head, (ii) chest, (iii) abdomen, and (iv) pelvis for one subject at each anatomical site. The subjects held their breath at endinhale for the chest examination and end-exhale for the abdomen and pelvis examinations during scanning. Scan conditions were the same as the clinical conditions (Table 1) except the scan ranges, which were 102.4 mm for head (one scan), 375.6 mm for chest (four contiguous scans), 93.9 mm for abdomen and pelvis (one scan). The matrix size was 51265126111251265126205, and the convolution kernel was the standard head kernel (FC43) for the head examination and the standard body kernel (FC10) for the others. Image quality was evaluated by three board-certified radiologists who had more than 10 years experience in clinical diagnosis. They compared quality of the images taken with the prototype scanner to their quality standard formed by experience. It took about 1.5 h to read the images obtained in multiple planes in all four cases.

Results For both CTs, DPIc, DPIp, and DPIw in an axial scan are summarized in Table 2. These values are normalized to 100 mAs. For the 256-slice CT, DPIw is 1966 mGy?mm/100 mAs for the head phantom and 1109 mGy?mm/100 mAs for the body phantom. For the 16-slice CT, DPIw is 181.6 mGy?mm/100 mAs with 8 mm Table 3. Dose profile integral weighted average (DPIw) for clinical protocols for 256-slice CT and 16-slice CT Examination

We imaged four healthy male volunteers (mean age 30.0 years¡7.6 (standard deviation) (SD); age range 23–53 years) using the 256-slice CT. The study was approved by the Institutional Review Board, and written informed consent was obtained from all subjects before starting. A non-enhanced examination with a step-and-shoot approval

Head Chest Abdomen Pelvis

DPIw

(mGy?mm)

256-slice CT

16-slice CT

6410 1775 4725 7088

12127 2462 5773 7981

DPIw percentage (%) 52.9 72.1 81.9 88.8

Table 2. Dose profile integral (DPI) for the 256-slice CT and 16-slice CT CT scanner

Phantom

Beam width (mm)

DPIc (mGy mm/100 mAs)

DPIp (mGy mm/100 mAs)

DPIw (mGy mm/100 mAs)

256-slice CT

Head Body Head Body Body

112 128 8 8 16

1829 781 174.2 67.7 117.6

2034 1273 185.3 98.7 175.0

1966 1109 181.6 88.4 155.9

16-slice CT

58



beam width for the head phantom, 88.4 mGy?mm/ 100 mAs with 8 mm beam width and 155.9 mGy?mm/ 100 mAs with 16 mm beam width for the body phantom. In Table 3 DPIw values are calculated for the clinical protocols. Values for the 256-slice CT are smaller than those for the 16-slice CT in all examinations. We note that especially in the head examination, the DPIw for the 256-slice CT is approximately 47% smaller than that for the 16-slice CT. With regard to the clinical image quality, Figure 2 shows normal anatomical images from the 256-slice CT. Auditory ossicles are observed clearly in the sagittal section with the same image quality as the state-of-the-art CT-scanner (Figure 2a). For the chest examination, 3D visualization of the lung from four contiguous axial scans is shown in Figure 2b. For the abdomen examination, the coronal image has an image quality as good as that of conventional CT (Figure 2c). For the pelvis examination, three contiguous coronal images are shown in Figure 2d. These images also show the same image quality as conventional CT.

Discussion In the present study, we compared doses in the 256-slice CT and the 16-slice CT for clinical conditions. The results showed that the dose for the 256-slice CT was smaller than that of the 16-slice CT in all examinations (Table 3).

(a)

(b)

The percentages of DPIw for the 256-slice CT to that for the 16-slice CT were 52.9%, 72.1%, 81.9% and 88.8% in the examinations of head, chest, abdomen and pelvis, respectively. The dose for the 256-slice CT was less than that of the 16-slice CT in all examinations for the following reason. In a MDCT-scanner the actual beam width is set as the nominal beam width (slice thickness6slice number) plus a certain margin, where the margin is added to cover penumbra and mechanical errors. X-ray photons incident on a marginal portion do not contribute to image formation, but they do contribute to increased dose. If the nominal beam width becomes large, the contribution of this portion becomes smaller. Thus, the 256-slice CT with larger beam width provides smaller DPIw values than the 16-slice CT. For the 16-slice CT the pelvis examination with 16 mm nominal beam width is more effective than the others with 8 mm beam width. In general, helical scans with pitch less than one caused overlap regions. Therefore in the present study, we set the effective mAs value to be the same to obtain the same signal-to-noise ratio in both CT systems. Notwithstanding the dose for the 256-slice CT being smaller than that of the 16-slice CT, the 256-slice CT provides sufficient image quality for diagnosis (Figure 2) [7]. In these clinical conditions, the 256-slice CT achieved a 0.5–0.8 mm isotropic resolution and large volumes of data were taken in a one-rotation scan [8]. Therefore coronal

(c)

(d)

Figure 2. Clinical images. (a) The 0.5 mm isotropic normal anatomy images of auditory ossicles in sagittal section. (b) 3D visualization of the chest with four contiguous scans. (c) Normal anatomy images of abdomen (0.63 mm reconstruction increment). (d) Coronal image (0.63 mm reconstruction increment) of pelvis with three contiguous scans.


59

S Mori, M Endo, K Nishizawa et al

Figure A1. Reconstruction geometry of cone-beam CT. An X-ray source and a 2D detector rotate around the z-axis. The volume that can be reconstructed with the Feldkamp algorithm is shown by the shaded region and is a double conical region within a cylinder of radius Rmax, which is determined by the detector size in the x-direction and shows the maximum field of view in the transverse plane. R and H show diameter and height, respectively, of a cylindrical reconstructed volume as it varied with an object. N6T show the nominal beam width where N is the number of slice and T is the slice collimation. Table A1. Calculated weighted CT dose index (CTDIw), dose– length product (DLP) and effective dose E for the 256-slice CT

Chest Abdomen Pelvis

DPIw (mGy?mm)

CTDIw (mGy)

DLP (mGy?cm)

E (mSv)

1775 2363 2363

13.87 18.46 18.46

130.2 173.3 173.3

2.21 2.60 3.29

and sagittal images were obtained at sufficient spatial resolution without secondary reconstruction. Regarding the diagnostic reference level, the effective dose [9] for the MDCT was approximately 15 mSv for routine chest examinations and 30 mSv for routine abdomen or pelvis examinations [10]. If these values are taken as upper limits and X-ray conditions are the same as those in Table 1, the acceptable scan time in volumetric cine imaging might be estimated in the following way. From Appendix 2, the estimated effective dose for a 1 s scan was 2.21 mSv, 2.60 mSv and 3.29 mSv for chest, abdomen and pelvis, respectively. Therefore, the acceptable scan time should be 6 s (5 15 [mSv]/2.21 [mSv]), 11 s (530 [mSv]/ 2.60 [mSv]) and 9 s (5 30 [mSv]/3.20 [mSv]) for chest, abdomen, and pelvis, respectively. As these scan times may not be sufficient for a dynamic study in some cases, further efforts are necessary to develop dose reduction methods such as automatic dose control [11–13], as well as to justify increasing the dose in dynamic studies consistent with riskbenefit. Resolution of these issues will allow full use of volumetric cine images which will significantly increase the amount of diagnostic information available to radiologists. In particular, we expect new applications such as computed tomographic angiography (CTA) of coronary arteries or perfusion studies of the whole brain.

Appendix 1. Field of view for the 256-slice CT In the 256-slice CT, the reconstructed images with the Feldkamp algorithm is the region that is passed 60

through during scanning by the tetra-angular pyramid whose apex and base are the X-ray source and the 2D detector, respectively (Figure A1). The reconstructed region is a double conical shape within a maximum FOV (Rmax) in the transverse plane that is determined by the detector size in the transverse direction. Reconstruction is not made in the entire Rmax except at the midplane and depends on a reconstructed FOV (R). In the case of the 256 mm60.5 mm (5 N6T) beam collimation, the length of the reconstruction region (H) is 102.4 mm for R5 240 mm and 93.9 mm for R5 320 mm. As seen in this example, the reconstructed region is generally smaller than the nominal beam width in cone beam CT.

Appendix 2. Effective dose estimation CT dose index (CTDI), dose–length product (DLP), and effective dose (E) are usually used for CT dosimetry [2], and they are derived from DPI described in the present report. CTDI is given as follows. ð 1 l=2 d(z)dz ½mGy CTDI~ (A1) NT {l=2 where N is the number of slices, T (mm) is the nominal slice thickness, and d(z) is the dose profile for an axial scan, l indicates the integration range. The International Electrotechnical Comission (IEC) recommended an integration range of 100 mm. However we used the integration range of 300 mm for the reason described. DPI is given with these notations as follows. ð l=2 DPI~ d(z)dz ½mGy mm (A2) {l=2

From Equations (A1) and (A2), CTDI~

1 DPI NT

(A3)

Weighted CTDI (CTDIw) is defined with CTDIs measured at the centre and peripheries of the phantoms as follows. 1 2 CTDIw ~ CTDIC z CTDIP ½mGy 3 3

(A4)

CTDIc and CTDIp represent the CTDI measured at the centre and the average CTDIs measured on the periphery of the phantom, respectively. CTDIw is given by DPIw as follows. CTDIw ~

1 DPIw ½mGy NT

(A5)

Dose–length product (DLP) for a complete examination is given as: DLP~CTDIw |L ½mGy cm

(A6)

where L (cm) is the scan range in the longitudinal direction. Estimation of effective dose (E) may be derived from values of DLP for an examination using appropriately normalized coefficients: E~EDLP . DLP ½mSv

(A7)



EDLP is the region-specific normalized effective dose (mSv mGy21 mm21) [9]. From these equations CTDIw, DLP and E can be calculated from measured DPIw. Table A1 gives calculated DPIw, CTDIw, DLP and E with one second scan of the 256-slice CT in the clinical conditions for chest, abdomen and pelvis examinations, respectively.

References 1. Endo M, Mori S, Tsunoo T, Kandatsu S, Tanada S, Aradate H, et al. Development and performance evaluation of the first model of 256-slice CT-scanner. IEEE Trans Nucl Sci 2003;50:1667–71. 2. International Commission on Radiological Protection. Managing patient dose in computed tomography. ICRP Publication 87. Annals of the ICRP 30 (4), Pergamon Press, Oxford, 2000. 3. Saito Y, Aradate H, Igarashi K, Ide H. ‘‘Large area 2dimensional detector for real-time 3-dimensional CT (256slice CT),’’ Proc SPIE 2001;4320:775–82. 4. Feldkamp LA, Davis LC, Kress JW. Practical cone-beam algorithm. J Opt Soc Am 1984;A 1:612–9. 5. International Electrotechnical Commission. Evaluation and routine testing in medical imaging departments constancy tests – X-ray equipment for computed tomography. Publication IEC 1994:1223-2-6. 6. Mori S, Endo M, Nishizawa K, Tsunoo T, Aoyama T, Fujiwara H, et al. Enlarged longitudinal dose profiles in cone-beam CT and the need for modified dosimetry. Med Phys 2005;32:1061–9.


7. Mori S, Endo M, Obata T, Murase K, Fujiwara H, Kandatsu S, et al. Clinical potentials of the prototype 256-detector roe CT-scanner. Acad Radiol 2005;22:149–55. 8. Mori S, Endo M, Tsunoo T, Kandatsu S, Tanada S, Aradate H, et al. Physical performance evaluation of a 256-slice CTscanner for 4-dimensional imaging. Med Phys 2004;31:1348– 56. 9. European Guidelines. Quality Criteria for Computed Tomography. EUR 16262, CEC Luxembourg, 1997. 10. Aoki C, Nishizawa K, Tonari A, Hachiya J. Effective dosing for multi-detector CT scanning. Jpn J Med Imaging 2001;20:101–9. 11. Nagel HS, Galanski M, Hidajat N, Maier W, Schmidt T. Radiation exposure in computed tomography-fundamentals, influencing parameters, dose assessment, optimization, scanner data, terminology. Hamburg: CTB Publications, 2001. 12. Thomas LT, Neil BB, Tin-Su P, Jerry R, Steven JW, Jianying Li, et al. A dose reduction X-ray beam positioning system for high-speed multislice CT scanners. Med Phys 2000;27: 2659–68. 13. Kachelrieß M, Kalender WA. Dose reduction by generalized 3D adaptive filtering for conventional and spiral single-, multirow and conebeam CT. Radiology 1999;213:283–4.

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E


Assessment of tube current modulation in pelvic CT G R IBALL, MSc, DipIPEM, D S BRETTLE, PhD and A C MOORE, MSc, DipIPEM Department of Medical Physics & Engineering, Leeds Teaching Hospitals NHS Trust, Leeds General Infirmary, Leeds LS1 3EX, UK

Abstract. An anatomically shaped polymethylmethacrylate (PMMA) phantom was used to assess the effect of the Siemens CARE Dose mA modulation system on pelvic CT scans. The effect of the system on absorbed dose to air, image percentage noise and the signal to noise ratio of clinically relevant details was assessed. The signal to noise ratio was calculated using Polytetrafluoroethylene (PTFE) and distilled water inserts; PTFE was used to represent bony structure and distilled water was used to represent soft tissue abscess. Pelvis protocols identified from local hospitals and the UK CT Dose Survey (2002), were assessed and compared with those provided by Siemens Medical (UK). These protocols were tested on a Siemens Sensation 4 CT scanner, both with and without CARE Dose. Results were obtained which showed that dose savings were possible with no significant increase in image noise. Dose reductions were 8% in the lateral positions in the phantom and 42% in the centre, top and bottom. The calculated ‘‘CTDIvol’’ was 32% lower with CARE Dose than without CARE Dose. This is slightly greater than the 25% change in the effective mAs values that was found. This implies that the reduction in the effective mAs values is a reasonable predictor of the total reduction in absorbed dose to air, whilst slightly underestimating the actual change. The results also showed a non-significant trend towards decreased signal to noise ratios for clinically relevant CT numbers when CARE Dose was activated. This suggests that tube current modulation may detrimentally affect signal detection due to changes in image noise.

CT examinations account for a large proportion of the collective dose from medical X-ray examinations in the UK. In 2000 this was reported as being 40% [1] but may now be even higher due to an increase in the range and volume of routine examinations and the uptake of CT fluoroscopy and cardiac CT scanning. There is a requirement for all X-ray examinations to be optimized such that the patient dose is ‘‘As Low As Reasonably Practicable’’ (ALARP) [2]. However, it is often difficult to implement procedures which significantly lower the radiation dose without decreasing the image quality to a non-diagnostic level. One recent technological advance from CT manufacturers in terms of dose reduction has been to introduce tube current modulation systems for CT scanning. The approach taken by Siemens Medical Systems (Erlangen, Germany) is a system called CARE Dose which claims to reduce patient doses whilst having no significant adverse effects on the image quality. This system has been described extensively in the literature [3–5]. The human body varies in composition both along its length and in the transverse plane at any given point along the body. This produces variations in X-ray attenuation due to both the external dimensions of the body and its internal composition. In CT scanning, as the X-ray tube and detectors rotate around the body, the attenuation can change by two orders of magnitude [4]. These differences in attenuation are most significant in the regions of the shoulder and pelvis, where large thicknesses of bone are found in the lateral projections, but a much smaller thickness of bone is present in the anterior–posterior projections. It is these examinations which provide the greatest challenges, in terms of the dose–image quality balance. As a result, using a constant tube current (mA) for each scan angle within a given rotation may result in Received 18 October 2005 and in final form 24 May 2005, accepted 31 May 2005.

62

either photon starvation artefacts on the high attenuation projections or overdosing in the lower attenuation projections. In the CARE Dose system, during each rotation of the tube and detector assembly around the patient, a small number of the central detector channels provide attenuation information, which is dependent upon the patient cross section and scan angle, to the X-ray generating system [3]. The information provided by these detector channels is used to determine to what extent the mA can be modulated, with respect to an initial tube current setting, without adversely affecting the image quality. As a result the tube current is modulated dynamically with a delay of one rotation relative to the attenuation measurement. The first patient based assessment by Greess et al [6] showed that, when CARE Dose is used, a dose reduction of approximately 25% (in terms of total mAs reduction) is possible in pelvic scanning ‘‘with no significant decrease’’ in subjective assessments of image quality. Similar percentage dose reductions have been demonstrated in other clinical work [7] and these showed good agreement with phantom based data [2, 5]. Most of the published work has used image noise and/or subjective image assessment to quantify image quality. A small number of papers [8, 9] have used standard deviations from regions of interest (ROIs) to yield a more objective assessment of image noise. Claims that the image quality was not affected by the CARE Dose system were queried by local users. Having used CARE Dose for a period of time, they perceived that the quality of the images for pelvis scans was subjectively worse when CARE Dose was used and this raised concerns that it may have a detrimental effect on the accuracy of diagnosis. This is despite the manufacturer’s recommendation that CARE Dose is used for all clinical situations other than for extremely large patients. This discrepancy between the reported claims and The British Journal of Radiology, January 2006

Assessment of tube current modulation

local experience prompted this investigation into the relationships between patient dose, image percentage noise and the signal to noise ratio (SNR) as an indicator of diagnostic detectability. The objective was to clarify whether the CARE Dose system can yield significant dose reduction for no loss of image quality in pelvic scanning.

Materials and methods A series of measurements were made using an anatomically shaped polymethylmethacrylate (PMMA) phantom which has been described in the literature [3] as a ‘‘hip’’ phantom. A schematic diagram of the phantom is shown in Figure 1. The thickness of the phantom is 14.5 cm in the z-axis. The hip phantom does not contain any bony structure and therefore the similarity of the phantom to the pelvic region is geometric only. As such it may be expected that the magnitude of the tube current modulation in clinical practice may be different from that found for this phantom.

Protocol selection Routine protocols for soft tissue assessment of the pelvis on Siemens 4 slice CT scanners (Somatom Volume Zoom and Somatom Sensation 4) were obtained from three local hospitals, the UK CT Dose survey 2002 and from Siemens Medical (UK) (Bracknell, UK). These protocols fell into two main groups, those that used a pitch of 1.00 and those that used a pitch of 1.25. There was some variation in the mA/mAs setting that was used, but all of the protocols used 120 kV and a rotation time of 0.5 s. In light of these findings all measurements were performed at the standard exposure factors given in Table 1. Reconstruction kernel B40s was used. Effective mAs is defined as the tube mAs per rotation divided by the helical pitch, where the pitch is the ratio of

Figure 1. A schematic diagram of the hip phantom. Solid cylinders within the phantom (1) represent the five ion chamber positions, dashed circles within the phantom (2) represent the five polymethylmethacrylate (PMMA) CT number and noise measurement positions; the dashed circle outside the phantom (3) represents the air CT number measurement position.

the table feed per rotation and the total X-ray beam width [10]. The effective mAs value of 165 was chosen as this was representative of most of the protocols that were obtained. Pitch settings of 1.00, 1.13 and 1.25 were used, both with and without CARE Dose. All the measurements were performed using the same PMMA hip phantom on the same scanner, a Siemens Sensation 4. Measurements were made which investigated how four different parameters changed with the application of CARE Dose. The four parameters that were investigated were (i) absorbed dose to air, measured in the phantom, (ii) image percentage noise, (iii) CT number for water and polytetrafluoroethylene (PTFE), and (iv) the SNR for both of these materials. These parameters were investigated for each pitch setting, both with and without CARE Dose. Water was chosen to represent low density abscess and PTFE to represent bony structure.

Consistency tests Prior to the testing, all of the test equipment was placed in the scanning room for at least 4 h in order for the temperature of the phantom and test equipment to stabilize with the room temperature. At the start of each visit the scanner was air-calibrated using the software on the scanner. A short series of consistency tests were performed immediately after the air-calibration which, on all subsequent visits, enabled us to verify that the performance of the scanner had not changed from the previous visit. On each occasion the hip phantom was positioned 15 cm from the end of the couch on top of the mattress in order to maintain consistent scattering conditions. The phantom was aligned using the laser lights on the scanner and with a spirit level. The set up is shown in Figure 2. A Scan Projection Radiograph (Topogram) of the phantom was acquired and a helical acquisition was planned from this image. The scan length for the helical acquisition was the whole length of the phantom

Figure 2. The phantom as positioned for the dose and noise measurements.

Table 1. Standard exposure factors for all scans Tube voltage (kV)

Effective mAs

Rotation time (s)

Beam collimation (mm)

Image slice thickness (mm)

SFOV (mm)

120

165

0.5

462.5

5

380

SFOV, field of view.


63

G R Iball, D S Brettle and A C Moore

(145 mm), which gave a total scan time of 8.2 s, using a pitch of 1.25. All of the consistency tests were performed with CARE Dose on. The first test was a measurement of the absorbed dose to air. A scan was performed at the standard exposure factors at a pitch of 1.25 with a calibrated 3 cm3 pencil ionization chamber (Capintec Inc., Ramsey, NJ), having an active length of 100 mm, in the central position. The chamber was connected to a Keithley 35050A Dosimeter (Keithley Instruments Inc., Cleveland, OH). The absorbed dose to air was recorded and the mean PMMA CT number and standard deviation (p) were measured adjacent to each of the five possible chamber positions (see Figure 1), on the CT slice closest to the centre of the phantom, using the region of interest (ROI) tool on the scanner. The size of the ROI that was used was kept constant throughout all of the measurements. The mean CT number of air was also measured at a standard position outside the phantom using a ROI of the same size. This scan and measurement procedure was then repeated with the ion chamber in the right lateral measurement position. On each occasion the ambient air temperature and pressure were measured, in addition to the phantom temperature, so that an air density correction could be applied to the dose measurements. The ion chamber was then removed from the phantom and a PTFE rod was inserted into the central measurement position. The scan was repeated and the mean CT number and p of the PTFE rod were recorded in addition to the measurements described above. Again this was repeated with the PTFE rod in the right lateral position.

Absorbed dose to air and noise measurements Absorbed dose to air measurements were made for each of the five chamber positions both with and without CARE Dose at each of the three pitch settings. For each scan the measured dose and total mAs were recorded. On the central slice the mean PMMA CT number and p were recorded at each measurement position and the mean air CT number was also recorded. For each measurement position the image percentage noise was calculated using Equation (1) [11]: Image Percentage Noise~

pPMMA 100 CTPMMA {CTAir

that was used in this study. However, the CTDIvol method is an accepted way of accounting for the distribution of dose within a phantom. Since, in this case, it is the comparison between the CTDIvol values for two different scanning situations, rather than the absolute value that was of most importance, the CTDIvol was used simply as an indicator of the relative change in absorbed dose to air. As such, the term ‘‘CTDIvol’’ is used for all calculations that relate to the hip phantom. The effect of the CARE Dose system on the percentage dose reduction was also evaluated over a range of initial effective mAs settings (50–200 mAs).

Signal measurements Two sets of signal to noise measurements were made, for the water and PTFE inserts. For the water measurements thin rubber sheaths were inserted into each of the five holes in the phantom and distilled water was inserted into each of the sheaths and the ends were secured with plastic clips. The sheaths were similar in diameter to the holes in the phantom which made it possible to almost completely fill the holes with water. The set up of the phantom for the water measurements is shown in Figure 3. For the PTFE measurements each individual rod was manufactured in house from a single PTFE rod (Barkston Plastics Ltd, Leeds, UK). All five rods were manufactured from the same batch of PTFE to ensure that there was no difference in composition between the individual rods. For each set of measurements the phantom was scanned five times at each pitch setting with and without CARE Dose. For each scan the mean CT number and p of the water/PTFE and PMMA were recorded at each measurement position on the central slice in addition to the mean CT number of air at the standard position. Measurements were also repeated 10 times on one scan of the PTFE rods in order to establish the repeatability of the measurements.

ð1Þ

where: p is the standard deviation and CT is the mean CT number (Hounsfield Unit) of the indicated material. The absorbed dose to air was corrected for ambient temperature and pressure and the ion chamber calibration factor was applied. The volume averaged CT dose index (CTDIvol) was then calculated for the scans with and without CARE Dose, using Equation (2) [11]. This was performed for each pitch setting: CTDIvol ~ 1=3 CTDIcentre z2=3 CTDIperiphery =pitch ð2Þ where: CTDIcentre is the CTDI measured in the centre of the phantom and CTDIperiphery is the average of the four CTDI values which were measured in the periphery of the phantom. Pitch is the ratio of the table feed per rotation and the total X-ray beam width. CTDIvol is actually defined for a cylindrical phantom and as such it is not strictly applicable to the hip phantom 64

Figure 3. The phantom as set up for the water signal to noise ratio (SNR) measurements. The British Journal of Radiology, January 2006


The SNR for the inserts was calculated Equation (3). CTsignal {CTPMMA SNR~ pPMMA

using

ð3Þ

Where: CTsignal is the mean CT number of water or PTFE and p is the standard deviation. The modulus was used as the mean CT number for water was sometimes below zero. SNR calculations were performed for each measurement point for each pitch setting. Error propagation was performed for all of the parameters of interest and the calculated values are shown with the results. The pooled standard deviation of the SNRs was calculated for each pitch setting for the water measurements and this result was used to power the study. The powering process showed that for a result to be statistically significant at the 95% level 25 measurements were required (both with and without CARE Dose). As a result, a further set of SNR measurements were made for both water and PTFE. The phantom was set up as described earlier and 25 scans were performed both with and without CARE Dose. For each scan the mean CT number and p of the insert (PTFE/water) was measured in the central position in addition to the mean CT number and p of the PMMA adjacent to the central insert. SNRs were calculated from these measurements and errors were calculated as for the previous measurements. Statistical analysis was performed on these results (Kruskal–Wallis non parametric test) to determine whether the SNRs of water and PTFE changed significantly for the scans with CARE Dose.

Dose measurements For the scans without CARE Dose (i.e. constant mA) the absorbed doses to air were significantly higher in the top and bottom positions than in the lateral positions. For the scans with CARE Dose there was a significant decrease in the absorbed dose to air in each position. The reductions were approximately 42% in the central position, 42% in the top and bottom positions and 8% in the lateral positions. These results were as expected and are shown in Figure 4. The calculated value of ‘‘CTDIvol’’ with CARE Dose was 32% lower than the value for the scans without CARE Dose. The scanner indicated reduction in effective mAs for the scans with CARE Dose was 25% (relative to the constant tube current case). The percentage reduction in ‘‘CTDIvol’’ was independent of pitch to within 0.5% over the pitch range of 1–1.25, as shown in Figure 5. The error bars that are shown in Figure 5 represent one standard deviation about the mean. The percentage reduction in absorbed dose to air in the central position for varying initial effective mAs settings is shown in Figure 6. The reduction in absorbed dose to air is approximately 40% for mAs settings between 50 mAs and 165 mAs. However, this reduction in absorbed dose to

Results The results of the consistency tests that were performed showed that on each occasion the performance of the scanner had not changed since the first visit. For clarity all the results for the 1.25 pitch setting are shown with summary results for the other pitches.

Figure 5. Percentage reduction in ‘‘CTDIvol’’ against pitch setting.

Figure 4. Variation of absorbed dose to air with position in phantom.


65


Noise measurements

Figure 6. Reduction in absorbed dose to air in the central position against initial effective mAs setting.

air rises to 50% at 200 mAs. It was not possible to obtain results for scans with mAs settings above 200 mAs as this would have exceeded the maximum tube loading at this pitch setting.

The initial set of image percentage noise values, calculated using Equation (1), across the five positions showed that, in general, there was not a large difference between the measured values with and without CARE Dose, other than for the top and bottom positions, as shown in Figure 7. These discrepancies between the noise values in the top and bottom positions were not found for the other pitch settings and are thought to be anomalous results, relative to the other pitch settings. As expected the noise values in the top and bottom were slightly different from those found in the lateral positions. For all pitches the highest noise values were found in the centre of the phantom. There was a general reduction in the image percentage noise as the pitch setting was increased (Figure 8), for both CARE Dose on and off, although this was not greater than the experimental uncertainties. For the 25 additional scans the image percentage noise was assessed in the centre of the phantom as for the initial tests. The difference between the noise values for CARE Dose on and off, which was approximately 10%, was tested for significance using the Kruskal–Wallis test. The mean and standard deviations of the noise values for the

Figure 7. Variation of image percentage noise phantom.

with

position

in

the

Figure 8. Image percentage noise variation with pitch setting.

66



25 additional scans are calculated p-values. These results show difference, at the 95% scans with and without

shown in Table 2 along with the that there was no significant level, in the noise levels for the CARE Dose.

SNR measurements The measurements for the water and PTFE inserts were used to calculate the SNR for each material (Equation (3)) and these results, for a pitch of 1.25, are shown in Figures 9 and 10. These results from the initial tests show that the SNRs, for both water and PTFE, are lowest in the centre

Table 2. Calculated mean, standard deviation and p-values for the image percentage noise tests

PTFE Water

Mean noise (SD) for CARE Dose off

Mean noise (SD) for CARE Dose on

p-value

19.4 (3.7) 17.3 (3.1)

19.0 (3.6) 19.1 (3.8)

0.727 0.099

SD, standard deviation; PTFE, polytetrafluoroethylene.

of the phantom both with and without CARE Dose. This is as expected as the noise values were highest in the centre of the phantom. The differences seen between the PTFE SNR values with and without CARE Dose, were generally within the experimental uncertainties. There appears to be a general decrease in the water SNR at each position for the scans with CARE Dose, which is an undesirable trend. However, the differences in water SNR were also within the experimental uncertainties. Figures 9 and 10 show that there are positional variations in the SNR within the phantom. As such it is not valid to average the SNR for the five different positions as this will mask the positional variations and will result in large uncertainties in the results. For a pitch of 1.25, the water SNR values measured with CARE Dose were lower than those without CARE Dose. This trend was observed for the other pitch values for the water scans but was not observed for the PTFE scans. For the 25 additional scans the signal and noise values were measured for PTFE and water in the central position in the phantom. From these results the SNR for both inserts were calculated as for the original scans. The differences between the values of CT number and SNR for the scans with CARE Dose on and off were tested for Figure 9.

Signal to noise ratios (SNR) for polytetrafluoroethylene (PTFE) for each phantom position.

Figure 10. Signal to noise ratios (SNR) for water for each phantom position.


67


significance using the Kruskal–Wallis test. The mean and standard deviations of the CT numbers and SNRs and the resulting p-values are shown on Table 3. These results show that there were no significant differences, at the 95% level, in CT number or SNR between the scans with and without CARE Dose despite the SNRs generally being decreased when CARE Dose was used.

Discussion Significant reductions in absorbed dose to air were found in all five positions in the phantom for the scans with CARE Dose relative to the constant tube current situation (Figure 4). The largest reductions, up to 42%, were found in the top, bottom and central positions as these positions lie on the lowest attenuation paths through the phantom and therefore experience the largest tube current modulation and reduction in absorbed dose to air. The dose reductions in the lateral positions, around 8%, are much smaller in magnitude as the attenuation is at its highest in these positions which means the tube current will be at its maximum value. These dose reductions were smaller than those shown by Kalender et al [5] who found a 45% ‘‘average’’ dose reduction via direct dose measurement. Kalender used a scan time of 1 s (compared with a 0.5 s scan time in our work) which allowed for a larger modulation amplitude and therefore a greater dose reduction than in this study. We also found that the reduction in the effective mAs values were lower than those found by Kalender et al, at approximately 25% compared with 40%. However, Kalender’s work showed an associated increase in image noise of approximately 10%, which was not found in the first part of this study. Kalender’s work was performed with a prototype version of the CARE Dose system which may also explain some of the differences between those initial results and the results of this study. Kalender et al [5] measured an average dose reduction of 45% in the hip phantom with a 3 cm3 ionization chamber similar to that used in this study. This 45% reduction in dose, however, was a straightforward average of the five measurement points rather than a volume average (‘‘CTDIvol’’) which was calculated here. A straightforward average of our results yields a dose reduction of 30%. The reduction in ‘‘CTDIvol’’ of approximately 32% was in good agreement with the relative dose reduction found by Gies et al [4], who found dose reductions of approximately 38%, for computer simulations using the hip phantom. The large reduction in absorbed dose to air in the central position is of importance as most of the more radiosensitive organs lie centrally. These results imply that the reduction in an individual organ dose (with an associated change in the effective dose) may be larger than the reduction in the values of ‘‘CTDIvol’’ shown

here. These results have implications for calculating effective doses in CT as the current Monte Carlo data sets that are used do not reflect the distribution of dose within the patient when a tube current modulation system is used. The large dose reduction in the centre of the phantom also has significant implications for pelvic scans of pregnant patients. If CARE Dose was used for these patients the risk to the fetus may be significantly reduced relative to scans performed with a constant tube current. Tack et al [10] showed that when using CARE Dose, the percentage dose reduction was independent of the initial effective mAs setting. They used six different mAs settings between 20 mAs and 100 mAs for chest and abdomen CT scans. Our results (Figure 6) show that the percentage dose reduction is approximately constant at a value of around 40% for initial effective mAs values up to 165 mAs. Above this value the percentage dose reduction increases, to approximately 50% at 200 mAs. This occurred as the mAs setting approached the maximum tube current rating for the tube. The Manufacturers recommend that for extremely large patients, where the mAs setting may be close to the tube limit, CARE Dose is not used. No measurements were made to determine whether or not the tube output varied linearly with mAs so we cannot exclude poor output linearity with mAs as a possible cause of the results shown in Figure 6. The image percentage noise level was not significantly affected by the application of CARE Dose, as shown in Figure 7, for the initial set of noise measurements. The reduction in dose of approximately 8% in the left and right positions occurs as a result of the integration of the reduction in tube current over all scan angles as there is no reduction in the tube current setting in the lateral projections. Given that there has been a general reduction in dose across the phantom there should have been an associated increase in the image percentage noise. No such increase in image percentage noise was found. Combining these results and those for the dose measurements shows that the reductions in absorbed dose to air that were calculated are net dose savings, i.e. they come with no significant noise penalty. Previous work [3, 5–7] showed that dose reductions of 23–45% were possible in the pelvis region with no significant difference in subjective assessments of image quality. The slight decrease in the image percentage noise with pitch setting, for both CARE Dose on and off is thought to be due to the combined effect of setting a constant effective mAs value and the magnitude of the over-scan which is necessary in helical scanning. For the additional scans with the water inserts there was an increase in the noise level of approximately 10% for the scans with CARE Dose on relative to the scans with CARE Dose off. This was not found to be significant at the 95% level (p50.099). This 10% increase in noise agrees

Table 3. Calculated mean, standard deviation and p-values for CT number and SNR for water and PTFE

PTFE Water

Mean CT number (SD) for CARE Dose off

Mean CT number (SD) for CARE Dose on

Mean SNR (SD) for CARE Dose off

Mean SNR (SD) for CARE Dose on

p-value (CT number)

p-value (SNR)

962.4 (4.6) 7.0 (3.6)

961.2 (4.5) 6.5 (3.3)

44.3 (8.8) 7.6 (1.7)

44.9 (8.8) 7.0 (1.6)

0.393 0.421

0.764 0.197

SD, standard deviation, SNR, signal to noise ratio, PTFE, polytetrafluoroethylene.

68



well Kalender’s work [5]. A similar change in noise was not found for the scans of the PTFE inserts (p50.727). As the PTFE provides much greater X-ray attenuation than water there is less scope for modulation of the tube current when the PTFE inserts are scanned. As a result the slightly larger reduction in the reported tube current that was found when the water inserts were scanned results in a larger percentage change in noise relative to the scans with CARE Dose off. Figures 9 and 10 show that there were differences between the SNRs calculated for the scans with CARE Dose on and off. These figures also show that the SNR varied with position within the phantom. The highest values of image percentage noise and the lowest values of SNR were found in the central position which is as expected from photon path length and reconstruction theories. For PTFE the SNRs for the scans with CARE Dose on showed no distinct trend relative to the SNRs for the scans without CARE Dose. This is in contrast to the situation for water where the SNRs for the scans with CARE Dose on were lower than those for the scans with CARE Dose off for 80% of the total number of scans. This shows that there is a trend towards decreased SNR for water when CARE Dose is used. The larger set of SNR measurements showed a difference in the SNRs of approximately 10% for water whilst there was no difference for the PTFE measurements. This is attributable to the similar percentage change in the noise which was found (Table 2). Statistical analysis showed that there was no statistically significant difference in the SNRs for PTFE and water between the situations with and without CARE Dose (p50.197 for water, p50.764 for PTFE). Table 3 shows that, at the 95% level, there was also no significant change in the CT numbers for water and PTFE for the scans with and without CARE Dose. Since the SNR depends on both the signal and noise, neither of which showed a significant change at the 95% level, there was no associated significant change in the calculated SNRs for both water and PTFE. This does not provide an explanation for the users’ subjective opinions that the images acquired with CARE Dose, for imaging pelvic abscess, were unsatisfactory. When the SNR values for water are error corrected (mean value minus uncertainty), the average SNR for the scans with CARE Dose is only just above the detectability threshold of 5 as defined by Rose [12]. Water has an inherently low SNR relative to the PMMA background, but this is further reduced by 10% when CARE Dose is activated. The worse case SNR (i.e. the lowest value of SNR taking into account the calculated uncertainties) was below the threshold value of 5 for 25% of the measurements with CARE Dose off and for 40% of the measurements with CARE Dose on. Although these differences may not be statistically significant they may be detectable by the person viewing the image and are therefore important differences. The X-ray attenuation path in clinical scanning is nonhomogeneous and the human pelvis may have an even more asymmetric attenuation pattern than this phantom. This may introduce a larger modulation in the tube current which would affect the noise and serve to further worsen the SNR situation. This may therefore reduce the confidence with which the viewer of the image can detect tissues which have subtle differences in SNRs. This The British Journal of Radiology, January 2006

combination of the decrease in the water SNR and the non-homogeneous attenuation path may therefore explain why subjectively the images that were acquired with CARE Dose had been reported as unsatisfactory for pelvic abscess imaging. There were large uncertainties in the results of this study. However, the reductions in the SNRs that were found were repeatable over a large number of scans and are therefore considered to be a true representation of the performance of the system. The main explanation for the large uncertainties was that the ROIs that were used for the water and PTFE measurements were small – these were limited by the size of the inserts which were, in turn, limited by the construction of the phantom. If measurements were made too close to the edge of the insert then the mean CT number would have been skewed by the presence of any air around the insert or by the background material itself. It was not possible to make any changes to the phantom design. If it had been possible to use larger inserts (and therefore larger ROIs) it may have been possible to obtain results which were less error dominated. We would recommend that any future studies should consider using larger inserts and ROIs to improve the noise statistics and to ensure homogeneity in the measurements taken within the signal areas. However, it should be noted that at 12 mm in diameter the size of the water inserts were representative of abscesses which are found in the pelvis. Some differences were found between the results for the left and right lateral positions in the phantom, in terms of absorbed dose to air, noise and SNR for both inserts. Further tests showed that the central alignment laser was inaccurate by approximately 3 mm which resulted in a relative difference between the left and right measurement positions of around 6 mm and that the differences were not due to the performance of the CARE Dose system.

Conclusions The CARE Dose system on Siemens 4 slice CT scanners results in significant dose savings for scans of the pelvic region. This yielded a reduction of approximately 32% in the value of ‘‘CTDIvol’’ which agreed well with the 25% reduction in the displayed effective mAs. This implies that the reduction in the effective mAs value can be used as an approximate indicator of the true dose reduction. This reduction is a real, net dose saving as there was no statistically significant increase in the noise. There appears to be a trend towards decreased SNRs for both water and PTFE when CARE Dose was used although no significant differences were found at the 95% level. These changes in SNR were mostly due to changes in the image percentage noise values. The largest decreases in SNR were found for water and were as large as 14%. Since the water inserts were representative of low-density abscess this suggests that the use of CARE Dose may decrease the visibility of low-density structures relative to the background. Therefore using CARE Dose in situations where subtle differences in low CT number tissue pathology are of interest may not be advisable. 69


Acknowledgments The authors wish to thank the following for their invaluable assistance in the work: Leeds Nuffield Hospital, especially Joanna Hartley for use of their scanner and for involvement in the measurement procedures; ImPACT, St George’s Hospital, London for loan of the phantom and general advice; Harrogate District Hospital, York District Hospital, UK CT Dose survey (2002) and Dr Paul Shrimpton of the National Radiological Protection Board (NRPB) for provision of protocol data; Siemens Medical (UK), Bracknell for protocol data.

References 1. Hart D, Wall BF. Radiation exposure of the UK population from medical and dental x-ray examinations (NRPB-W14). Chilton: National Radiological Protection Board, 2002. 2. Annals of the ICRP 26 (2). Radiological protection and safety in medicine. ICRP Publication 73, ICRP, 1996. 3. Kalender WA, Wolf H, Seuss C, Gies M, Greess H, Bautz WA. Dose reduction in CT by online tube current control: principles and validation on phantoms and cadavers. Eur Radiol 1999;9:323–328. 4. Gies M, Kalender WA, Wolf H, Suess C. Dose reduction in CT by anatomically adapted tube current modulation. I. Simulation studies. Med Phys 1999;26:2235–47.

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5. Kalender WA, Wolf H, Seuss C. Dose reduction in CT by anatomically adapted tube current modulation. II. Phantom measurements. Med Phys 1999;26:2248–53. 6. Greess H, Wolf H, Baum U, Lell M, Pirkl M, Kalender WA, et al. Dose reduction in computed tomography by attenuation based online modulation of tube current: evaluation of six anatomical regions. Eur Radiol 2000;10:391–4. 7. Greess H, Nomayr A, Wolf H, Baum U, Lell M, Bowing B, et al. Dose reduction in CT examination of children by an attenuation based online modulation of tube current (CARE Dose). Eur Radiol 2002;12:1571–6. 8. Mastora I, Remy-Jardin M, Seuss C, Scherf C, Guillot JP, Remy J. Dose reduction in spiral CT angiography of thoracic outlet syndrome by anatomically adapted tube current modulation. Eur Radiol 2001;11:590–6. 9. Jakobs TF, Becker CR, Ohnesorge B, Flohr T, Seuss C, Schoepf UJ, et al. Multislice helical CT of the heart with retrospective ECG gating: reduction of radiation exposure by ECG-controlled tube current modulation. Eur Radiol 2002;12:1081–6. 10. Tack D, De Maertelaer V, Gevenois PA. Dose reduction in multidetector CT using attenuation based online tube current modulation. AJR Am J Roentgenol 2003;181:331–4. 11. Institute of Physics and Engineering in Medicine Report 32, Measurement of the Performance Characteristics of Diagnostic X-ray Systems used in Medicine Part III Computed Tomography X-ray Scanners, 2nd Edition, IPEM, 2003. 12. Rose A. Vision: human and electronic. New York: Plenum, 1974:21–3.



E


Short communication

Radiosurgical palliation of aggressive murine SCCVII squamous cell carcinomas using synchrotron-generated X-ray microbeams 1

M MIURA, PhD, 2H BLATTMANN, PhD, 3E BRAÜER-KRISCH, BEng, 3A BRAVIN, PhD, 1 A L HANSON, PhD, 1M M NAWROCKY, BA, 1P L MICCA, BS, 1,4D N SLATKIN, MD and 4 J A LAISSUE, MD 1

Medical Department, Brookhaven National Laboratory, Upton, NY 11973-5000, USA, 2Niederwiesstrasse 13C, Untersiggenthal, Switzerland, 3European Synchrotron Radiation Facility, 6 Rue Jules Horowitz, BP 220, Grenoble, France and 4Pathologisches Institut der Universita¨t Bern, Murtenstrasse 31, Bern, Switzerland

Abstract. Microbeam radiosurgery (MBRS), also referred to as microbeam radiation therapy (MRT), was tested at the European Synchrotron Radiation Facility (ESRF). The left tibiofibular thigh of a mouse bearing a subcutaneously (sc) implanted mouse model (SCCVII) of aggressive human squamous-cell carcinoma was irradiated in two orthogonal exposures with or without a 16 mm aluminium filter through a multislit collimator (MSC) by arrays of nearly parallel microbeams spaced 200 mm on centre (oc). The peak skin-entrance dose from each exposure was 442 Gy, 625 Gy, or 884 Gy from 35 mm wide beams or 442 Gy from 70 mm wide beams. The 442/35, 625/35, 884/35 and 442/70 MBRSs yielded 25 day, 29 day, 37 day and 35 day median survival times (MST) (post-irradiation), respectively, exceeding the 20 day MST from 35 Gy-irradiation of SCCVIIs with a seamless 100 kVp X-ray beam.

A century ago, radiotoxic doses of X-rays delivered through a flexible grid of 1 mm thick strands of iron woven 3.5 mm on centre and a thin, continuous underlay of leather (a low-Z filter), pressed hard against the skin to blanch it, were able to palliate deep malignancies safely; iron-shielded epidermal cells healed the resultant punctate skin burns within 2 weeks [1]. After half a century, such millimetre-scale grid therapy (GT) was generally superseded by skin-sparing megavoltage radiotherapy, although at least one centre is currently pursuing a version of GT clinically [2]. It was, however, the radiobiological studies in mice, which used a deuteron microbeam to simulate cosmic radiation in space [3] that led to microbeam radiosurgery (MBRS) investigations, GT’s micrometre-scale analogue. The MBRS studies have continued since ,1990 using ,200–800 Gy doses of ,30–200 keV X-rays delivered almost instantaneously through an array of multiple nearly parallel microslices of tissues [4–13]. Putatively, MBRS irreparably damages microsegments of neoplastic but not of normal endothelium; surviving clonogenic tumour cells may be insufficiently perfused and too sparse to re-grow. Received 7 January 2005 and in revised form 2 June 2005, accepted 16 June 2005. This manuscript has been sponsored by Brookhaven Science Associates, LLC under Contract No. DE-AC02-98CH10886 with the United States Department of Energy. The US Government retains, and the publisher, by accepting the article for publication, acknowledges, a world-wide license to publish or reproduce the published form of this manuscript, or allow others to do so, for the US Government purposes. Funding was provided by the DOE Office of Biological and Environmental Research, the Institute of Pathology of the University of Bern, and the European Synchrotron Radiation Facility.


Imminently lethal intracerebral rat 9L gliosarcomas have been palliated with 25 mm wide microbeams, 100 mm on centre (oc). About 4 months later when untreated controls had long been euthanized for tumour overgrowth, 50%, 18%, or 36% of rats remained alive after crossfired 625 Gy, crossfired 312 Gy, or unidirectional 625 Gy skinentrance doses, respectively [7]. Despite its weak immunogenicity [14] and robust radioresistance [15], the deadly aggressive squamous-cell carcinoma (SCCVII) can be ablated either by immunotherapy [16] or by X-irradiation using a radiosensitizer [17]. However, the outcome of an experimental therapy for the murine SCCVII carcinoma is generally informative in terms of growth delay rather than ablation [18, 19]. Accordingly, we compared SCCVII growth delays and their normal-tissue radiotoxicities following different MBRS strategies to enable future ranking of various proposed clinical MBRS treatment plans.

Material and methods Radiation source MBRS was performed at the ID17 beamline of the European Synchrotron Radiation Facility (ESRF), a 6 GeV electron storage ring with an operating current of 180–200 mA. Beamline ID17 is equipped with a 1.6 T wiggler, which produces a beam of X-rays [20, 21] with a median energy of 38.1 keV. The beam is filtered with 1.5 m each of C and Al followed by 1.0 mm Cu. This filtration hardened the spectrum to 93 keV at maximum intensity, suitable for MBRS. The beam emerged from the beam pipe through a beryllium window to air in the 71

M Miura, H Blattmann, E Braüer-Krisch et al

radiation-shielded ID17 irradiation hutch, where it was collimated to 18 mm60.5 mm.

Collimator and irradiations The microbeams were created with a variable width tungsten multislit collimator (MSC) (Tecomet, Woburn, MA) before impinging on the animal [13]. For MBRS, the anaesthetized mouse was placed prone, lengthwise, on the 15 cm61.5 cm surface of a 15 cm66.5 cm61.5 cm PlexiglasH block, each foreleg and the left, tumour-bearing hind leg gently taped to the sides of the block (Figure 1). The first exposure (of the entire tumour-bearing left tibiofibular thigh) was nearly anteroposterior, with the mouse saddle rotated 5 ˚ clockwise (from the horizontal 0 ˚ reference direction of the oncoming beam) about a vertical axis (as seen by an observer looking downward toward the mouse, Figure 2) to avoid irradiating the left foreleg; the second (orthogonal) exposure was implemented after the block was rotated 95 ˚ clockwise from the 0 ˚ reference direction about the same vertical axis. Although each of the two 16 mm broad, 15 mm high anatomical (skinentrance) targets had its estimated vertical and horizontal midplane at the estimated level of the centre of the tumour, the actual upper horizontal limit of each target was parallel to and 1 mm below the long edge of the block’s upper surface. The right hind leg had been taped slightly backward to avoid exposure to microbeams during the second exposure. For each irradiation, a computer-guided platform moved the mouse directly upward (at several cm s21) past the microbeam array emerging from the MSC. The shutteractivated exposure time was selected to conform to the slowly decaying ring current and the pre-programmed upward acceleration and speed of the platform.

Animal tumour model SCCVII murine squamous cell carcinoma cells (Prof. J Martin Brown, Stanford University) were cultured in D-MEM enriched with 10% fetal bovine serum, 1%

Figure 2. A PlexiglasH polymethylmethacrylate block (thick black outline) served as a ‘‘saddle’’ for the mouse, viewed as it would be by an observer directly above it. The mouse was anaesthetized and placed prone on the block for its first tumour irradiation. In this figure, the outline of the mouse is represented by an ellipse. Two black dots represent its eyes. To avoid irradiating its left foreleg, the 150 mm long axis of the block was rotated 5 ˚ (about a vertical axis through the centre of the block) clockwise from the reference 0 ˚ microbeam direction. The microbeam array, symbolized by thin arrows, was propagated in a thin, wide, slightly divergent fan-beam, substantially in a horizontal plane, represented here as the plane of this page. The second irradiation was implemented after the block was rotated 95 ˚ clockwise from the 0 ˚ reference direction about the same vertical axis.

penicillin/streptomycin, and 1% L-glutamine. Only passages 1–3 were used to initiate tumours. Cells (26105 in 0.05 ml of medium) were then implanted subcutaneously (sc) into the left thighs of 20–25 g female C3H mice (Taconic Farms, Germantown, NY or Charles River Laboratories, Wilmington, MA). Alternatively, freshly removed ,1 mm3 fragments of mouse tumours that had been initiated sc on the dorsal thorax with 56105 cells in 0.1 ml of medium [22] were minced in saline, then implanted sc in the left thighs through a 16-gauge trocar. All MBRS-irradiated tumours and 16 of the 40 untreated control tumours grew from cell suspensions. Our preliminary studies had shown that growth rates using cell suspensions were the same as those using tumour fragments; the former are preferred because the suspended cells do not seem to form satellite tumours along the trocar track when implanted. Mice bearing ,80–100 mg tumours (as estimated from volume < x2y/2, where x , y) were anaesthetized (0.01 ml per gram of body weight (gbw) of an aqueous 6 mg ml21 sodium pentobarbital solution) by intraperitoneal injection, (,60 mg gbw21) for MBRS. Mice were irradiated 10 days after tumours were implanted. Median survival times (MST) are defined as the time interval between the day the treated groups were irradiated, which is equivalent to 10 days after tumour implantation, and the day they were euthanized unless otherwise stated.

Therapy studies Single-exposure irradiations were used throughout. Tumour dimensions were measured 2–3 times per week and mice were euthanized either when estimated tumour volumes exceeded 500 mm3 or when skin ulceration or severe oedema (foot diameter . 5 mm) was observed. Mice were weighed whenever the tumours were measured, except during the first week after irradiation, when they were weighed daily. Figure 1. Photograph of anaesthetized female C3H mouse bearing a leg squamous-cell carcinoma (SCCVII) carcinoma taped to PlexiglasH block, readied for microbeam radiosurgery (MBRS) at the European Synchrotron Radiation Facility (ESRF). 72

100 kVp seamless X-rays In three groups of anaesthetized mice placed prone on a horizontal surface, tumours were X-irradiated at 2.10 Gy min21 vertically downward, delivering 25 Gy or The British Journal of Radiology, January 2006

Short communication: MBRS for murine SCCVII carcinomas

35 Gy. A Philips RT-100 generator was operated at 100 kVp and 8 mA with a 0.4 mm thick Cu filter, a 10 cm focus-to-skin distance, and a 2.5 cm collimator aperture in contact with the thigh. Radiation dosimetry was carried out using an air-equivalent thimble ionization chamber, adhering to the 1996 IPEB code of practice for 10– 300 kVp, Cu-filtered X-rays [23].

Irradiation groups The rapidly growing SCCVII cancers were treated in a clinically analogous way, i.e. after the tumours became palpable, which took 7 days after implantation (volumes ¢50 mm3). They were then sorted into groups bearing tumours of comparable size and were irradiated 3 days later, 1 day after they arrived at the ESRF. Microbeam widths were either 35 mm or 70 mm and the on centre (oc) distances for each of the treatment groups were 200 mm. Groups 1, 2 and 3 were irradiated at skinentrance doses of 442 Gy, 625 Gy, and 884 Gy, respectively, using 35 mm microbeam widths in each direction. Group 4 was similarly irradiated to Group 3 (884 Gy) but with a 16 mm aluminium filter upstream from the collimator. Group 5 was irradiated at a skin-entrance dose of 442 Gy with 70 mm microbeam widths in each direction and Group 6 was similarly irradiated but with the 16 mm aluminium filter. The control group comprised 40 untreated SCCVII tumour-bearing mice from five separate experiments.

Results 100 kVp seamless X-irradiation at 25 Gy and 35 Gy yielded MSTs of 14 days and 20 days, but long-term survivals were only 0/10 and 1/9, respectively (Figure 3a). Untreated controls had a MST of only 6 days or a median post-implantation survival time of 16 days. MBRS survival data are shown in Figure 3b. Figure 4 shows average growth rates of various irradiated and control SCCVIIs. In Groups 1 and 2, euthanasia was

Figure 4. Average relative tumour volumes of the various microbeam radiosurgery (MBRS)-irradiated and control mice. The lower tumour volumes noted in groups 3 to 6 relate to the fact that those tumours had regressed to relatively small or undetectable volumes when most of the mice had to be euthanized due to severe radiodermatitis of the inner thigh.

usually for tumour overgrowth (volume ¢500 mm3); in Groups 3–6, it was mainly for foot/leg damage (severe oedema; diameter of the foot .5 mm) (Table 1). Figures 3 and 4 do not distinguish those reasons for euthanasia. Euthanasia necessitated by skin radiotoxicity probably prevented much longer survivals of the third of 884/35 and 442/70 MBRS mice that showed no residual tumour at necropsy. MBRS yielded long-term survival rates (up to 153 days) of 0/12 in Groups 1 and 2, 1/10 in Group 3, and 0/10 in Groups 4, 5, and 6. Group 3 (884/35 without aluminium) showed the highest median survival time, and only 1/10 was euthanized for tumour overgrowth; but 8/10 were euthanized for foot/leg damage and only 1 of those 8

Figure 3. Kaplan-Meier graphs of C3H mice bearing aggressive squamous-cell carcinoma (SCCVII) leg carcinomas irradiated with various radiation modalities. The on-centre distances for microbeam radiosurgery (MBRS)-irradiations were 200 mm. Mice euthanized due to foot/leg damage were not distinguished from those euthanized due to tumour overgrowth: (a) Survival graphs of mice bearing SCCVII carcinomas treated with seamless 25 Gy or 35 Gy skin-entrance doses of X-rays in comparison with unirradiated controls. (b) Survival graphs of similar mice in MBRS groups (1–6) with skin entrance doses of 442 Gy, 625 Gy, and 884 Gy at 35 mm and 442 at 70 mm beam width. ‘‘Al’’ designates a 16 mm-thick aluminium filter placed upstream from the collimator. The British Journal of Radiology, January 2006

73

M Miura, H Blattmann, E Braüer-Krisch et al Table 1. Number of ablated tumours, median survival times and explanations for euthanasia in mice treated with microbeam radiosurgery (MBRS) tracked up to 153 days after irradiation compared with those treated with seamless X-rays and with untreated controls Group Dose/beam width (MBRS) 1 2 3 4 5 6

442 Gy/35 mm 625 Gy/35 mm 884 Gy/35 mm 884 Gy/35 mm + aluminium filter 442 Gy/70 mm 442 Gy/70 mm + aluminium filter Unirradiated control 25 Gy seamless 35 Gy seamless

Euthanized Number 153-d tumour Post-irradiation Euthanized (tumour Euthanized (foot/leg damage) (foot/leg damage) of mice control median survival overgrowth) mice with tumours time (days) 12 12 10 10

0 0 1 0

25 29 41 33

11 9 1 3

1 3 8 7

1 2 7 5

10 10

0 0

38 31

1 1

9 9

5 6

40 10 9

0 0 1

6 14 20

40 10 8

0 0 0

– 0 0

energy per beam (442/70) than in the group with the narrower microbeam imparting greater energy per beam (884/35). The radiation field of the SCCVII carcinoma on the mouse leg encompassed the entire thigh, but not the foot (Figure 1). Radiodermatitis was most marked in the inner thigh and oedema was most severe in the left hind foot below the irradiation field. We attribute the latter to ablation of overirradiated lymphatics proximal to the foot. At the higher radiation doses, such damage limited survival time more than did tumour overgrowth. The radiodermatitis of the inner thigh was explained with microdosimetry simulations using the MCNPX code [25]. The simulations were performed assuming a water phantom of the left mouse thigh, shaped as an inverted, truncated cone (16 mm high with a 13 mm diameter top and a 3 mm diameter bottom) in which a 0.4 mm diameter sphere of water, the phantom tumour, was embedded. Computations showed that doses between the microbeams (‘‘valley doses’’) in the epidermis adjoining the PlexiglasH would have been ,25% less without contributions from back-scattered X-rays. Even at 1.5 mm from the PlexiglasH, the dose would have been reduced ,15% if the PlexiglasH was not present.

showed no tumour at necropsy (Table 1). In contrast, in Group 5 (442/70 without aluminium), 9/10 mice were euthanized due to severe foot/leg damage, of which 4/9 euthanized mice (two each on days 27 and 31) showed no residual tumour at necropsy. On those same days (27 and 31) in Group 3, seven and four mice, respectively, had no tumours, indicating that the lower incidence of tumours in Group 5 compared with Group 3 was due to the earlier time of euthanasia. Some non-parametric Wilcoxon Two-Sample analyses to rank palliation rates, using a morbidity/mortality index technique [24] on days 9 and 23, before foot/leg damage became apparent, are shown in Table 2. Without regard to radiodermatitis, Group 3 (884 Gy) followed by Group 5 (442/70, showed the most effective tumour palliation, which was expected from the survival graphs (Figure 3b) and because they received the highest tumour ionization energies.

Discussion Figures 3 and 4, and Table 1, demonstrate that MBRS delayed tumour growth more than did seamless 100 kVp X-rays and that the former effect is dose-dependent. However, in the higher dose groups, the plotted growth rates after day 40 are based on only a few animals, as many mice had to be euthanized due to radiodermatitis. The MSTs of each MBRS-treated group were longer than were the 14 day and 20 day MSTs observed for the seamless 25 Gy or 35 Gy groups, respectively. Normal tissue damage occurred more quickly in mice irradiated with the broader microbeams imparting less

Conclusions Palliation of the exceptionally radioresistant murine SCCVII carcinoma was better from MBRS than from seamless 35 Gy irradiation with no more risk to normal tissue in the radiation field. Normal-tissue damage in the higher-dose MBRS groups, especially to the left foot

Table 2. p-values from the non-parametric Wilcoxon Two-Sample Test on tumour volumes using morbidity/mortality indices [24] on days 9 and 23 after irradiationa Group

Days after irradiation

625/35 (2)

884/35 (3)

884/35+Al (4)

442/70 (5)

442/70+Al (6)

442/35 (1) 442/35 (1) 625/35 (2)

9 23 23

0.011 0.001

0.001 0.001

0.017

0.003 0.001 0.017

0.036 0.004

a

Differences between one group (numbered in parentheses) in the top row and another numbered in the left column were deemed significant if p ¡ 0.05, in which case that column shows the group with the better tumour palliation. No other pairs of groups showed an advantage in palliation that was significant at the p ¡ 0.05 level.

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Short communication: MBRS for murine SCCVII carcinomas

below the radiation field, could be deemed clinically irrelevant as most of that damage was anatomically remote from the cancer in structures that would have been spared high doses under clinical circumstances. Left foot oedema probably resulted from radiation-induced strictures of proximal blood vessels and lymphatics. Thus our computations suggest that MBRS of such SCCVII tumours using similar skin-entrance doses without the irradiated skin in contact with the PlexiglasH may enable a greater proportion of mice to survive long-term.

Acknowledgments The authors thank Mr Seymour Brittman of Brittman & Son, East Northport, New York, for constructing the ventilated hardwood cases to enclose mouse cages for intercontinental air transportation. We also thank Mr Larry McMillan of Swiss International Airlines and his coworkers for facilitating our air travel with mice at the JFK Airport.

References 1. Ko¨hler A. Zur Ro¨ntgentiefentherapie mit Massendosen. Mu¨nchener medizinische Wochenschrift 1909;56:2314–6. 2. Zwicker RD, Meigooni A, Mohiuddin M. Therapeutic advantage of grid irradiation for large single fractions. Int J Radiat Oncol Biol Phys 2004;58:1309–15. 3. Curtis HJ. The microbeam as a tool in radiobiology. Adv Biol Med Phys 1963;175:207–24. 4. Slatkin DN, Spanne P, Dilmanian FA, Sandborg M. Microbeam radiation therapy. Med Phys 1992;19:1395–400. 5. Slatkin DN, Dilmanian FA, Nawrocky MM, Spanne P, Gebbers J-O, Archer DW, et al. Design of a multislit, variable-width collimator for microplanar beam radiotherapy. Rev Sci Instr 1995;66:1346–7. 6. Slatkin DN, Spanne P, Dilmanian FA, Gebbers J-O, Laissue JA. Subacute neuropathological effects of microplanar beams of x-rays from a synchrotron wiggler. Proc Natl Acad Sci USA 1995;92:8783–7. 7. Laissue JA, Geiser G, Spanne PO, Dilmanian FA, Gebbers J-O, Geiser M, et al. Neuropathology of ablation of rat gliosarcomas and contiguous brain tissues using a microplanar beam of synchrotron-wiggler-generated X rays. Int J Cancer 1998;78:654–60. 8. Dilmanian FA, Morris GM, Zhong N, Bacarian T, Hainfeld JF, Kalef-Ezra J, et al. Murine EMT-6 carcinoma: high therapeutic efficacy of microbeam radiation therapy. Radiat Res 2003;159:632–41. 9. Schweizer PM, Spanne P, Di Michiel M, Jauch U, Blattmann H, Laissue JA. Tissue lesions caused by microplanar beams of synchrotron-generated X-rays in Drosophila melanogaster. Int J Radiat Biol 2000;76:567–74. 10. Laissue JA, Lyubimova N, Wagner H-P, Archer DW, Slatkin DN, Di Michiel M, et al. Microbeam radiation therapy, in medical applications of penetrating radiation. In: Bradford Barber H, Roehrig H, editors. Proceedings of SPIE 1999;3770:38–45.


11. Laissue JA, Blattmann H, Di Michiel M, Slatkin DN, Lyubimova N, Guzman R, et al. The weanling piglet cerebellum: a surrogate for tolerance to MRT (microbeam radiation therapy) in pediatric neuro-oncology. In: Bradford Barber H, Roehrig H, Doty FP, Schirato RC, Morton EJ, editors. Penetrating radiation systems and applications III. Proceedings of SPIE 2001;4508:65–73. 12. Blattmann H, Burkard W, Djonov V, Di Michiel M, Brauer E, Stepanek J, et al. Microbeam irradiation of the chorioallantoic membrane (CAM) of chicken embryo. Strahlenther Onkol 2002;178:118. 13. Braüer-Krisch E, Bravin A, Zhang L, Siegbahn E, Stepanek J, Blattmann H, et al. Characterization of a tungsten/gas multislit collimator [TMSC] for microbeam radiation therapy at the European Synchrotron Radiation Facility. Rev Sci Instr 2005;76:0643031–7. 14. Kanazawa H, Rapacchietta D, Kallman RF. Scheduledependent therapeutic gain from the combination of fractionated irradiation and cis-diamminedichloroplatinum(II) in C3H/Km mouse model systems. Cancer Res 1988;48:3158–64. 15. Suit H, Allam A, Allalunis-Turner J, Brock W, Girinsky T, Hill S, et al. Is tumor cell radiation resistance correlated with metastatic ability? Cancer Res 1994;54:1736–41. 16. Mandpe AH, Tsung K, Norton JA. Cure of an established nonimmunogenic tumor, SCC VII, with a novel interleukin 12-based immunotherapy regimen in C3H mice. Arch Otolaryngol Head Neck Surg 2003;129:786–92. 17. Nakajima S, Fujii T, Murakami N, Aburano T, Sakata I, Nakae Y, et al. Therapeutic and imaging capacity of tumor-localizing radiosensitive Mn-porphyrin KADT-F10 for SCCVII tumors in C3H/He mice. Cancer Lett 2002;181:173–8. 18. Ning S, Yu N, Brown DM, Kanekal S, Knox SJ. Radiosensitization by intratumoral administration of cisplatin in a sustained-release drug delivery system. Radiother Oncol 1999;50:215–23. 19. Katori K, Baba Y, Imagawa Y, Nishimura G, Kagesato, Takagi E, et al. Reduction of in vivo tumor growth by MMI166, a selective matrix metalloproteinase inhibitor, through inhibition of tumor angiogenesis in squamous cell carcinoma cell lines of head and neck. Cancer Lett 2002;178:151–9. 20. Thomlinson W, Berkvens P, Berruyer G, Bertrand B, Blattmann H, Braüer-Krisch E, et al. Research at the European Synchrotron Radiation Facility medical beamline. Cell Mol Biol 2000;46:1053–63. 21. Braüer-Krisch E, Bravin A, Lerch M, Rosenfeld A, Stepanek J, Di Michiel M, et al. MOSFET dosimetry for microbeam radiation therapy at the European Synchrotron Radiation Facility. Med Phys 2003;30:583–9. 22. Fu KK, Rayner PA, Lam KN. Modification of the effects of continuous low dose rate irradiation by concurrent chemotherapy infusion. Int J Radiat Oncol Biol Phys 1984;10:1473–8. 23. Klevenhagen SC, Aukett RJ, Harrison RM, Moretti C, Nahum AE, Rosser KE. The IPEMB code of practice for the determination of absorbed dose for x-rays below 300 kV generating potentials (0.035 mm Al-4 mm Cu HVL; 10-300 kV generating potential). Phys Med Biol 1996;41:2605–25. 24. Coderre JA, Slatkin DN, Micca PL, Ciallella JR. Boron neutron capture therapy of a murine melanoma with pboronphenylalanine: dose-response analysis using a morbidity index. Radiat Res 1991;128:177–85. 25. Los Alamos National Laboratory Report LA-UR-03-5916; August, 2003.

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Solitary pulmonary nodule with growth and contrast enhancement at CT: inflammatory pseudotumour as an unusual benign cause 1

S DIEDERICH, Prof. Dr. med., 2D THEEGARTEN, Priv. Doz. Dr. med., 3G STAMATIS, Prof. Dr. med. and 4 ¨ THEN, Priv. Doz. Dr. med. R LU 1

Department of Diagnostic and Interventional Radiology, Marien Hospital, Academic Teaching Hospital, Rochusstr. 2, D-40479 Du¨sseldorf, 2Institute of Pathology, BG-Kliniken Bergmannsheil, Ruhr University Bochum, 3Department of Thoracic Surgery and Endoscopy, Ruhrland Hospital Essen and 4Department of Medicine, Marien Hospital Du¨sseldorf, Germany

Abstract. Small (¡10 mm) pulmonary nodules are frequently detected at modern chest CT. As most of these nodules are benign, non-invasive classification is required – usually based on assessment of growth and perfusion. Absence of growth and no evidence of perfusion, as demonstrated by lack of enhancement at contrast-enhanced CT or MRI, strongly suggest a benign nodule. On the other hand, growth with a doubling of the nodule’s volume between 20 days and 400 days or enhancement suggest a malignant nature of the lesion. We present an example of a nodule with strong contrast enhancement and a doubling time of approximately 260 days, which histologically represented a benign inflammatory pseudotumour.

Case report A 56-year-old asymptomatic male underwent chest radiography in two views as part of a general health survey. This revealed a small non-calcified nodule projected over his right mid lung field not demonstrated on a chest radiograph obtained 3 years previously. CT of the chest (collimation 5 mm) confirmed a non-calcified nodule in the lateral segment of the right middle lobe adjacent to a subsegmental artery and bronchus with a diameter of approximately 9 mm. No other abnormality was demonstrated, in particular no hilar or mediastinal lymphadenopathy was observed. The patient presented to our hospital for a second opinion 3 months later. Evaluation of size and contrastenhancement was performed obtaining limited spiral CT data sets with a collimation of 1 mm (Somatom Plus 4; Siemens, Erlangen, Germany) before and 1 min, 2 min, 3 min and 4 min after administration of 1.4 cm3 kg21 body weight iomeprerol (Imeron 300H; Altana Pharma, Konstanz, Germany) with an injection rate of 2 cm3 s21. Images were displayed at lung and mediastinal windows (Figure 1a, b, c). Nodule density was measured in regions of interest representing 70% of the nodule’s cross section at anatomically identical levels. Density was 27 Hounsfield Units (HU) before contrast injection and increased to 80 HU, 95 HU, 63 HU and 62 HU after 1 min, 2 min, 3 min, and 4 min. Thus, maximum enhancement after 2 min was 68 HU. The diameter of the nodule was again measured to be 9 mm (Figure 1a). It has been shown that lack of contrast-enhancement almost excludes malignancy with a negative predictive value of 96%, whereas demonstration of contrast-enhancement allows no differentiation between benign and Received 24 February 2005 and in revised form 18 April 2005, accepted 29 April 2005.

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malignant nodules [1]. Thus, the patient was informed that malignancy could not be excluded and biopsy was recommended. Due to the central location of the relatively small nodule adjacent to a subsegmental artery, it was felt that percutanous biopsy was not appropriate and surgical biopsy was suggested. The patient, however, did not agree to immediate biopsy. Therefore, follow-up thin section CT with 1 mm slice thickness was performed at 6 months and 10 months. There was questionable growth at 6 months and definite growth at 10 months (Figure 2). The nodule’s volume was calculated from measurements digitally on the monitor of a workstation in the axial plane and also by counting the number of contiguous 1 mm slices for estimation of the diameter in craniocaudad direction. As the nodule appeared almost ideally spherical at the baseline measurement (9 mm) and at 10 month followup (12 mm) its volume was calculated (V54/3 p r3) as 381 mm3 (baseline measurement) and 904 mm3 (10 months) resulting in a doubling time of 8.6 months. Again surgical biopsy was recommended and now the patient agreed. As the nature of the nodule could not be established prior to surgery it was decided to proceed to minimally invasive thoracotomy. During surgery a tumour measuring 18 mm624 mm adjacent to the medial segmental bronchus of the middle lobe was palpated rendering wedge resection impossible. Due to the small volume of the middle lobe, primary middle lobectomy was performed including resection of regional lymph nodes. Final histological assessment of the well-circumscribed lesion (Figure 3) including immunostaining for CD-1a, CD-3, CD-20, CD-68, and EMA showed a mixed inflammatory infiltrate and connective tissue typically for a benign inflammatory pseudotumour (Figure 4a). Diagnosis of a malignant tumour could not be confirmed. Several vessels were demonstrated within the nodule (Figure 4b). The The British Journal of Radiology, January 2006

Case report: Pulmonary nodule with contrast enhancement and growth at CT

(a)

(b)

(c)

Figure 1. Dynamic thin-section CT scan before ((a) lung window, (b) mediastinal window) and 2 min after ((c) mediastinal window) contrast enhancement: The nodule shows an enhancement of 68 Hounsfield units.

post-operative course was unremarkable and the patient was discharged from hospital after 8 days.

Discussion Pulmonary nodules are common findings at chest radiography and even more at chest CT. With the introduction of spiral CT, and particularly multirow-detector spiral CT, an increasing number of small nodules (¡10 mm) is detected. The ratio of benign and malignant nodules strongly depends on nodule size. In nodules .10 mm the proportion of malignant nodules is high requiring biopsy in many cases [2]. In nodules ¡10 mm more than 90% of nodules are benign [3]. Therefore, biopsy is not routinely performed in these lesions and non-invasive diagnostic

Figure 2. Follow-up thin-section CT scan after 10 months revealing growth of the nodule to 12 mm.


tests are required to differentiate between benign and malignant nodules. The two techniques used routinely are analysis of nodule growth and perfusion. It has been shown that most malignant tumours exhibit doubling times between 20 days and 400 days, whereas faster or slower doubling times suggest benign lesions [4, 5]. Also, assessment of tumour perfusion is helpful in predicting a nodule’s nature. As malignant tumours .1 mm require neoangiogenesis for further growth, all malignant tumours visible at chest CT should exhibit enhancement at contrast-enhanced CT or MRI [1]. It has been shown that absence of contrast enhancement strongly predicts the benign nature of a nodule (e.g. granuloma); on the other hand, not all enhancing nodules are malignant due to enhancing benign lesions such as inflammatory nodules or intrapulmonary lymph nodes [1, 6, 7].

Figure 3. Histological slide showing the nodule with focal sclerosis (Haematoxylin and eosin, magnification62). 77

S Diederich, D Theegarten, G Stamatis and R Lu¨then

(a)

(b)

Figure 4. Histological slides demonstrating a mixed inflammatory infiltrate with lymphocytes and plasma cells ((a) Haematoxylin and eosin, magnification6200) and involvement of medium-sized vessels with occlusion ((b) CD34 staining, ABC method, magnification6100).

Our case is another example of a benign nodule with strong enhancement that also exhibited growth with a volume doubling time suspicious for malignancy. Inflammatory pseudotumour (other diagnostic terms: fibroxanthoma, xanthogranuloma, xanthofibroma, histiocytoma) is a rare entity histologically composed of a mixture of inflammatory cells, including plasma cells, lymphocytes, macrophages, a few eosinophils, fibroblasts and connective tissue. In cases with dominance of plasma cells, the term plasma cell granuloma is used. The lesions are typically solitary, round and well circumscribed. The diameter varies from 0.8 cm to 36 cm. Pulmonary inflammatory pseudotumours clinically present in 60% of patients with symptoms such as cough, dyspnoea and haemoptysis; 40% are asymptomatic. The lesion presents in patients ranging from 1 year to 77 years, but approximately 60% are under the age of 40 years [8, 9]. Radiologically, most inflammatory pseudotumours present as well-defined nodules or masses measuring between 1 cm and 10 cm. The large difference in the maximum size reported in the literature probably depends on the mode of detection as well as the presence or absence of symptoms. Inflammatory pseudotumours are slightly more common in the lower lobes. If followed radiographically, growth has been documented. Contrast studies usually demonstrate significant enhancement of the lesions. Cavitation or calcification is rare. Infiltration of adjacent organs may be observed and misinterpreted as evidence of malignancy [9, 10]. In symptomatic patients surgical resection is the therapy of choice. In conclusion, inflammatory pseudotumour has to be included in the differential diagnosis of enhancing pulmonary nodules with growth particularly in children

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and young adults. As there is no specific imaging feature, biopsy is required for the diagnosis.

References 1. Swensen SJ, Viggiano RW, Midthun DE, et al. Lung nodule enhancement at CT: multicenter study. Radiology 2000;214:73–80. 2. Tan BB, Flaherty KR, Kazerooni EA, Iannettoni MD. The solitary pulmonary nodule. Chest 2003;123:89S–96S. 3. Kim YH, Lee KS, Primack SL, et al. Small pulmonary nodules on CT accompanying surgically resectable lung cancer: likelihood of malignancy. J Thorac Imaging 2002;17:40–6. 4. Yankelevitz DF, Henschke CI. Does 2-year stability imply that pulmonary nodules are benign? AJR Am J Roentgenol 1997;168:325–8. 5. Yankelevitz DF, Gupta R, Zhao B, Henschke CI. Small pulmonary nodules: evaluation with repeat CT – preliminary experience. Radiology 1999;212:561–6. 6. Bankoff MS, McEniff NJ, Bhadelia RA, Garcia-Moliner M, Daly BDT. Prevalence of pathologically proven intrapulmonary lymph nodes and their appearance on CT. AJR Am J Roentgenol 1996;167:629–30. 7. Matsuki M, Noma S, Kuroda Y, Oida K, Shindo T, Kobashi Y. Thin-section CT features of intrapulmonary lymph nodes. J Comput Assist Tomogr 2001;25:753–6. 8. Colby TV, Koss MN, Travis WD. Inflammatory pseudotumor. In: Tumors of the lower respiratory tract. Atlas of tumor pathology (3rd edn), Fascicle 13. Washington, DC: Armed Forces Institute of Pathology, 1995:327–38. 9. Agrons GA, Rosado de Christensen ML, Kirejczyk WM, Conran RM, Stocker JT. Pulmonary inflammatory pseudotumor: radiologic features. Radiology 1998;206:511–8. 10. McCall IW, Woo-ming M. The radiological appearances of plasma cell granuloma of the lung. Clin Radiol 1978;29:145–50.



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Case report

Non-haemorrhagic subdural collection complicating rupture of a middle cranial fossa arachnoid cyst C OFFIAH, BSc, FRCS, FRCR, W ST CLAIR FORBES, MA, DMRD, FRCR and J THORNE, FRCS Departments of Neuroradiology, Hope Hospital, Salford Royal Hospitals NHS Trust, Stott Lane, Salford, Manchester M6 8HD and Royal Manchester Children’s Hospital, Central Manchester and Manchester Children’s University Hospitals NHS Trust, Manchester, UK

Abstract. Arachnoid cysts are a common incidental finding on routine brain imaging and, for the most part, their presence is uneventful. Occasionally they may be associated with haemorrhage into the subdural compartment. Rarer still is simple rupture of the contents of the arachnoid cyst into the extra-axial space. MRI can help distinguish between these two rare occurrences – an important distinction to make as this may assist in directing the treating clinician toward the most appropriate management plan.

Arachnoid cysts are a well-recognized benign intracranial lesion occuring most commonly in the middle cranial fossa. Although most are small and asymptomatic, they may be associated with a complicated course most typically causing mass effect or hydrocephalus. Spontaneous and post-traumatic intracystic and subdural haemorrhage has also been reported. We describe a case of the very rare complication of symptomatic rupture of a middle cranial fossa cyst into the subdural compartment without haemorrhage. Despite extensive literature review, there has been no previous description of this.

Case report An 8-year-old boy presented with a history of intermittent headaches, vomiting and double vision over a period of several weeks, the onset of which was related to a fall playing football when he struck his head on concrete. No loss of consciousness occurred at the time of the injury. On examination, his Glasgow coma scale (GCS) was 15 and there was no focal neurology, cranial nerve deficit or papilloedema. CT performed on admission demonstrated a low-attenuation right-sided subdural collection – isodense to cerebrospinal fluid (CSF) – causing moderate mass effect, compression of the ipsilateral ventricle and effacement of the cortical sulci. No focal intraparenchymal abnormality was present. A left middle cranial fossa arachnoid cyst was also noted (Figure 1). The patient underwent burr-hole drainage of the right subdural collection and at surgery fluid consistent with CSF was seen to escape under pressure. He was discharged a few days later following improvement but re-admitted shortly after that with recurrence of his symptoms and a notable right-sided scalp swelling. A repeat CT scan (not shown) demonstrated re-accumulation with a slight increase in size of the right subdural collection, which remained low-attenuation. The left arachnoid cyst remained unchanged and evidence of a Received 3 August 2004 and in revised form 9 January 2005, accepted 6 May 2005.


small right Sylvian fissure arachnoid cyst persisted. Right-sided extracranial soft-tissue swelling was also demonstrated. Ultimately, a subdural-peritoneal shunt was placed on the right and his subsequent recovery was unremarkable. Prior to his discharge, a CT scan was performed which showed only a thin residual right-sided subdural collection with some associated subdural air, but no residual mass effect. The left-sided arachnoid cyst was noted as previously. The small right-sided arachnoid cyst has become more readily appreciated with resolution of the ipsilateral subdural fluid (Figure 2). Since his discharge, a MR scan of the brain has been performed (7 months since the index admission). The subdural collection has completely resolved. The rightsided middle cranial fossa arachnoid cyst has increased significantly in size since the preceding CT examination obtained during the admission some 6 months previously. The left-sided arachnoid cyst has remained unchanged in size (Figure 3). The patient remains clinically well and the subduralperitoneal shunt in situ.

Discussion Arachnoid cysts derive from the meninx primitive, embryologically, which is the primitive membrane ensheathing the developing central nervous system (CNS). As subarachnoid CSF accumulates, this meninx cavitates and resorbs under normal circumstances leaving only the subarachnoid space and the arachnoid membrane. During this process, the arachnoid membrane may split with secretion of fluid by the arachnoid cells into the resulting cleft ultimately yielding a cyst – the so called arachnoid cyst – which is truly intra-arachnoid anatomically [1–3]. Both intracystic haemorrhage and rupture of middle cranial fossa arachnoid cysts into the subdural space resulting in acute or chronic subdural haematoma – either spontaneous or post-traumatic – have been well documented in the medical literature [1–14]. Bleeding occurs due to tearing of an unsupported bridging vein or veins that are stretched by the cyst and susceptible to rupture by 79

C Offiah, W St Clair Forbes and J Thorne

(a)

(b)

Figure 1. (a) Unenhanced CT brain demonstrating a right subdural effusion causing mass effect and (b) a left middle cranial fossa arachnoid cyst. The right Sylvian fissure demonstrates notable prominence of low (cerebrospinal fluid) density consistent with an underlying right-sided middle cranial fossa arachnoid cyst.

a rise in intracystic pressure [5, 7, 15]. What has not been highlighted in the radiological literature is the occurrence of arachnoid cyst rupture into the subdural compartment resulting in progressive symptoms of raised intracranial pressure, despite the lack of haemorrhage. Appreciation of this complication does appear to have implications in relation to the management of these patients and is a valuable differential to highlight to the referring clinician. This is because there potentially remains a communication between the arachnoid cyst and the subdural compartment following rupture so that, despite burr hole drainage of the collection, there remains predisposition to re-accumulation of cyst fluid in the subdural compartment and therefore the increased probability of drain insertion being required as an immediate definitive treatment. An important imaging manoeuvre to assist in this differentiation would be early MRI as the signal characteristics of the subdural collection would aid distinction between acute or subacute haemorrhage, as opposed to rupture of arachnoid cyst contents into the subdural compartment. In the case of the latter, the signal characteristics of the subdural fluid would present as isointense to cyst contents (and to CSF). It would appear that very minor trauma, if any, is required for arachnoid cyst rupture to occur [6]. In our case study the head injury that preceded the onset of symptoms was not associated with any loss of consciousness at the time suggesting that the insult was indeed a minor one. Rupture has been reported to occur in cases following the Valsalva manoeuvre during various activities 80

such as swimming [4, 5]. There have been sporadic reports in the medical literature regarding spontaneous disappearance of middle cranial fossa arachnoid cysts following rupture or haemorrhage into the subdural space with eventual resorption [4, 5]. Various mechanisms have been proposed for such resolution [4–6]. However, in our case report the cyst was seen to increase in size consistent with re-accumulation after the subdural collection had been treated and had begun to resolve. Presumably, diversion of the subdural accumulation with shunt placement reduced the intracranial pressure enough for the arachnoid cyst to re-accumulate. The re-accumulation of the right-sided arachnoid cyst in this case may also have been aided by the widely conjectured ‘‘flap-valve’’ effect that may result after a tear in the inner cyst wall following rupture that allows passage of CSF from subarachnoid space into the cyst, but closure of the tear in the outer membrane that allowed cyst contents to egress from the cyst into the subdural compartment [4, 5, 9]. Forty-eight percent of arachnoid cysts occur in the middle cranial fossa. Only 20% occur in the posterior fossa [10–15]. It is reported that only middle cranial fossa cysts rupture [15] and this is supported by a review by Rogers et al that demonstrated six cases of subdural haematomas, which were all associated with middle cranial fossa arachnoid cysts [9]. In our experience arachnoid cysts are frequently bilateral; the presence of a middle cranial fossa arachnoid cyst and a contralateral subdural fluid collection should therefore raise the possibility of rupture The British Journal of Radiology, January 2006

Case report: Subdural collection complicating arachnoid cyst

Figure 2. Unenhanced CT brain following shunt drainage of the right subdural collection. Only a small residual effusion remains (with some air). The left middle cranial fossa cyst appears unchanged. The presence of the right middle cranial fossa cyst is more readily appreciated.

of a contralateral arachnoid cyst as a consideration, particularly if early MRI fails to confirm the presence of haemorrhagic subdural fluid contents.

Conclusion We have demonstrated the rare complication of rupture of a middle cranial fossa arachnoid cyst into the subdural space without haemorrhage following minimal trauma. Although it is well recognized that arachnoid cysts may be associated with acute and eventually chronic subdural blood following rupture due to tearing of the vessels that bridge the cyst wall, non-haemorrhagic rupture into the subdural compartment is an important radiological differential diagnosis to consider in order to direct the clinical/surgical management of such patients optimally as the imaging appearances of these two entities on CT examination can be identical. In such cases, early MRI would be a valuable adjunct.

References 1. Robertson R, Caruso PA, Truwit CL, Barkovich AJ. Disorders of brain development. In: Atlas SW, editor. Magnetic resonance imaging of the brain and spine. Philadelphia, PA: Lipincott Williams and Wilkins, 2002:316–9.


Figure 3. Axial T2 weighted MR brain performed 6 months later, confirming the re-accumulation of the right middle cranial fossa arachnoid cyst as indicated by the interval increase in size as well as the presence of the unaltered left middle cranial fossa arachnoid cyst. No subdural collection was present this time. 2. Barkovich AJ. Hydrocephalus. In: Barkovich AJ, editor. Pediatric neuroimaging. Philadelphia, PA: Lippincott Williams and Wilkins, 2000:592–5. 3. Intracranial arachnoid and ependymal cysts. In: Wilkins RH, Rengachary SS, editors. Neurosurgery. New York, NY: McGraw-Hill, 1985:2160–72. 4. Cullis PA, Gilroy J. Arachnoid cyst with rupture into the subdural space. J Neurol Neurosurg Psychiatry 1983;46:454–6. 5. Inoue T, Matsushima T, Tashima S, Fukui M, Hasuo K. Spontaneous disappearance of a middle cranial fossa arachnoid cyst associated with subdural hematoma. Surg Neurol 1987;28:447–50. 6. Rakier A, Feinsod M. Gradual resolution of an arachnoid cyst after spontaneous rupture into the subdural space. J Neurosurg 1995;83:1085–6. 7. Eustace S, Toland J, Stack J. CT and MRI of arachnoid cysts with complicating intracystic and subdural haemorrhage. J Comput Assist Tomogr 1992;16:995–7. 8. Ochi M, Morikawa M, Ogino A, Nagoki K, Hayashi K. Supratentorial arachnoid cyst and associated subdural hematoma: neuroradiological studies. Eur Radiol 1996;6:640–4. 9. Rogers M, Klug GL, Siu KH. Middle fossa arachnoid cysts in association with subdural haematomas. A review and recommendations for management. Br J Neurosurg 1990;4:497–501. 10. Gupta R, Vaishya S, Mehta VS. Arachnoid cyst presenting as subdural hygroma. J Clin Neurosci 2004;11:317–8. 11. Donaldson JW, Edwards-Brown M, Luerssen TG. Arachnoid cyst rupture with concurrent subdural hygroma. Pediatr Neurosurg 2000;32:137–9.

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C Offiah, W St Clair Forbes and J Thorne 12. Gelabert-Gonzalez M, Fernandez-Villa J, Cutrin-Prieto J, Allut AG, Martinez-Rumbo R. Arachnoid cyst rupture with subdural hygroma: report of three cases and literature review. Childs Nerv Syst 2002;18:609–13. 13. Poirrier AL, Ngosso-Tetanye I, Mouchamps M, Misson JP. Spontaneous arachnoid cyst rupture in a previously asymptomatic child: a case report. Eur J Paediatr Neurol 2004;8:247–51.

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14. Cayli SR. Arachnoid cyst with spontaneous rupture into the subdural space. Br J Neurosurg 2000;14:568–70. 15. Shapiro KN, Swift DM. Intracranial arachnoid cyst. In: Tindall GT, Cooper PR, Barrow DL, editors. The practice of neurosurgery. Baltimore, MD: Williams and Wilkins, 1996;2667–79.


The British Journal of Radiology, 79 (2006), 83

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Correspondence (The Editors do not hold themselves responsible for opinions expressed by correspondents)

Social factors perception

in

improving

radiological

The Editor—Sir, Manning, Gale and Krupinski are absolutely correct when they state ‘‘good displays and tools are clearly necessary ……but what we need to understand is how the Radiologist interacts with the displayed information during the reading process in order to determine how we can further improve decision making’’ [1]. They identify many of the perceptual, cognitive and ergonomic factors. Social factors also need to be addressed. There is considerable opportunity for improved knowledge sharing among radiologists. Performance can be improved by knowing where in particular to look and what exactly to look for in different clinical scenarios. While subspecialization is important in improving perception, targeted instruction and top up training can improve the performance of all [2]. We need to have much more extensive prompt feedback of our discrepancies and errors. The prevalence of eye strain among radiologists has been reported [3] as has the medicolegal implications of reporting at a significantly faster rate than average [4]. There is no magic solution that will produce a perfect imaging perceptual process. However, the social dimension of reporting needs to be included to optimize our performance. Yours etc., R FITZGERALD

Consultant Radiologist Radiology Department New Cross Wolverhampton WV10 0QP UK (Received 12 August 2005 and accepted 23 August 2005) DOI: 10.1259/bjr/18395574

References 1. Manning DJ, Gale A, Krupinski EA. Commentary: Perception research in medical imaging. Br J Radiol 2005;78:683–5. 2. FitzGerald R. Radiological error: analysis, standard setting, targeted instruction and teamworking. Eur Radiol 2005;15:1760–7. 3. Vertinsky T, Forster B. Prevalence of eye strain among radiologists. Influence of viewing variables on symptoms. AJR Am J Roentgenol 2005;184:681–6. 4. Berlin L. Liability of interpreting too many radiographs. AJR Am J Roentgenol 2000;175:17–22.

Sir Godfrey Hounsfield was well aware of the possibility of what at EMI was called volume scanning. He appreciated the difference between single slice and multiple slice data acquisition. At EMI, the former idea was based on movement of the patient through the scanner while the continuous (slip-ring) gantry rotated using a continuous power X-ray source. The latter technique was to have involved the collection of data from set of contiguous slices at the same time, a technique originally described in Sir Godfrey’s first CT patent. It was to achieve volume scanning that the TOPAZ geometry was invented. The patent makes clear the continuous rotation nature of the scanner. This system was conceived in the mid 1970s with discussion for implementation with the commercial Division later the same decade. This system used solid-state detectors and an X-ray tube with a directly oil cooled anode. A photograph of the prototype, built by the Research team in the Central Research Laboratories of EMI, was shown on the 25 May 2005. It is clear that this geometry was the first to make possible the matching of continuous rotation with a continuous power X-ray source. The apparent falter in the development of CT in the 1980s can be traced to other causes and not to a lack of technical innovation. The nature of the TOPAZ configuration also uniquely enabled focused layers to be obtained from the scannogram or pilot scan data (Zonogram). In the mid 1980s a 1 s, 1 mm slice thickness, version of the system, based on Sir Godfrey’s ideas, was successfully built and fully tested. From the scanner volume clinical scans were obtained from which 3D images were reconstructed. The implication that there were any constraints placed on the future of CT by Sir Godfrey is therefore wholly inaccurate. If not the father of volume scanning, Sir Godfrey Hounsfield must indeed be considered to be the grandfather. Yours etc., A BASKERVYLE STRONG Engineering Manager EMI Medical Ltd (retired) Broombank 267 Penistone Road Kirkburton Huddersfield HD8 0PF (Received 16 August 2005 and accepted 22 August 2005)

The Grandfather of volume scanning

DOI: 10.1259/bjr/15532036

The Editor—Sir, I would like to express my concern after some comments made at the memorial lectures held at the Royal Society on 25 May 2005.

We hope the article in this issue by E Beckmann (p. 5) and this 16 letter will2005 rectify any omission—Editor. (Received August and accepted 22 August 2005)


83


E


Case of the month

A deformed skull with enlarging hand and feet in a young female 1

B GUGLANI, MD, 1C J DAS, MD, DNB, 1A SEITH, MD, 2N TANDON, DM and 2B A LOWAY, MD

1

Radiodiagnosis and 2Department of Endocrinology, All India Institute of Medical Sciences, New Dehli, India

A 25-year-old woman presented with a 6 year history of gradually enlarging swelling at the back of her head. She had also noticed enlarging hands and feet with increased prominence of eyes for the last 3–4 years. She had been amenorrhoeic for the past 2 years. On physical examination, her height was 163 cm. There was facial deformity with a prominent right side of the face and bony swelling in the region of the external occipital protuberance. Her hands and feet were enlarged with a doughy consistency. In addition, mild scoliosis was found in the mid-thoracic region with convexity towards the right side. The left humerus was short and bowed and the left rib cage was deformed with multiple swellings. There was no evidence of abnormal skin pigmentation. Her thyroid was mildly enlarged. Galactorrhoea was also observed. Visual field examination revealed bitemporal hemianopsia on perimetry. Endocrine evaluation showed a non-suppressible growth hormone level (GH) of 60 ng ml21 and an increased prolactin level of 43 mg l21. Plain radiography of the skull was obtained (Figure 1). Subsequently, contrast enhanced MRI (CEMRI) of the sella was performed (Figure 2). What is the diagnosis in this case?

Figure 1. Plain radiograph skull (lateral view).

Received 25 February 2005 and in revised form 22 April 2005, accepted 31 May 2005.

(a)

(b)

Figure 2. (a) Pre- and (b) post-contrast enhanced MR coronal images through the sella turcica. 84


Case of the month: A deformed skull

Discussion The diagnosis of acromegaly was clinically and biochemically unequivocal in this case. Subsequently, we performed contrast enhanced MRI (CEMRI) of the sella for evaluation of the patient’s acromegaly. A T1 weighted image of the sella revealed a large sellar, suprasellar mass compressing the optic chiasm and causing expansion of the sella. The mass showed enhancement in post-contrast imaging suggestive of a pituitary macroadenoma, thus confirming the clinical diagnosis of acromegaly. Marked, diffuse expansion of the skull base was also noted enhancing on contrast administration (Figure 2). The patient sought medical advice primarily for her marked skull expansion seen in the occipital region, which had increased gradually over the last 6 years. Plain radiography of the skull showed gross expansion of the skull base and occiput with areas of sclerosis and ground glass density. Enlarged maxillary and frontal sinuses were also seen. A CT performed for detailing of the foraminal compression in the skull base showed the typical groundglass appearance of fibrous dysplasia (Figure 3). Narrowing of the bilateral optic canal and orbital apices were also seen. A subsequent skeletal survey revealed that the involvement was indeed multifocal with expansile lesions seen in the ribs, left humerus and radius (Figure 4). All of these features pointed to a pathology in addition to acromegaly due to a GH secreting pituitary adenoma. Based on the radiological appearance of the ground glass density and their characteristic distribution, a differential diagnosis of coexisting fibrous dysplasia was made. As both of these conditions are associated in only one genetic abnormality, a final diagnosis of McCune-Albright syndrome – polyostotic fibrous dysplasia with acromegaly due to pituitary macroadenoma – was made. The McCune-Albright syndrome (MCAS) is a sporadic disorder characterized by polyostotic fibrous dysplasia, cutaneous pigmentation and endocrine hyperfunction. The presence of any two of the three lesions (skin, bone and endocrine glands) is sufficient for the diagnosis of MCAS

Figure 3. Axial CT image (bone-window) showing groundglass expansion of skull base with narrowed basal foramina, bilateral optic canal and orbital apices.


Figure 4. Radiographs of the left humerus and radius show classical lesions of fibrous dysplasia.

[1]. The genetic basis of MCAS is now reasonably understood and is due to the post-zygotic activating mutations of arginine 201 in the guanine-nucleotidebinding protein (G protein) alpha-subunit (Gsalpha), leading to a mosaic distribution of cells bearing constitutively active adenylate cyclase [2]. The resultant disorder depends on when the mutation occurs; during embryonic development or post-natal life. The earlier it takes place, the more cells are affected. Somatic mutations in a small cell mass result in MCAS; whereas in a larger cell mass, mutation results in polyostotic fibrous dysplasia [3]. The distribution of affected cells follows embryological lines of ectodermal migration, which explains the unilateral and focal expression of MCAS in bones as well as in endocrine tissue. Various endocrinopathies reported in MCAS include precocious puberty, thyrotoxicosis, Cushing’s syndrome, acromegaly, hyperprolactinaemia and hypophosphataemic rickets [4, 5]. The association of polyostotic fibrous dysplasia and acromegaly, although rare, is a well described entity [5]. In MCAS, gigantism/acromegaly usually present at an earlier age (less than 30 years) than in classical acromegaly [6–8]. A pituitary adenoma may be found less often than in classical acromegaly [5, 8]. 85

B Guglani, C J Das, A Seith et al

Moreover, the macroadenomas in MCAS are smaller than those in classical acromegaly [5, 8]. The association of acromegaly and fibrous dysplasia may pose a diagnostic challenge to the clinician and the radiologist. The majority of patients with MCAS are short in stature because of precocious puberty, recurrent fractures and hypophosphataemic rickets, whereas those with associated GH excess/acromegaly usually reach a normal height [5, 9]. Also, since fibrous dysplasia has a predilection for skull base involvement, the facial dysmorphism may mask the usual features of acromegaly causing delay in diagnosis. CT and MRI play a pivotal role in the evaluation of these patients. MRI is better than CT in assessment of sella in the presence of bony skull base thickening due to fibrous dysplasia. The distinction between pituitary gland and abnormal fibrous bone tissue at skull base is better made on MRI. The combination of pre- and post-contrast images is useful in this regard. However, CT of skull base plays a useful role in some cases for detailing neural foraminal compression, especially if surgery is being contemplated [10]. Craniofacial fibrous dysplasia may mimic hyperostotic meningioma (meningioma en plaque) or even osteoma, especially in a monostotic lesion [9]. Association of acromegaly and meningioma has also been described [12, 13]. MCAS with acromegaly and skull base fibrous dysplasia is also a therapeutic challenge as transpituitary surgery is often not possible in the presence of fibrous dysplasia of skull base whereas radiation therapy can induce bone sarcomatous transformation [14]. Some authors have suggested a transfrontal route to approach the pituitary adenoma [15]. In a series by Akintoye et al, the authors described a distinct clinical phenotype of MCAS due to GH excess which is characterized by inappropriately normal stature, thyroid releasing hormone (TRH) responsiveness, prolactin cosecretion, small or absent pituitary tumours, a consistent but inadequate response to treatment with cabergoline and an intermediate response to long acting octreotide [5]. In our patient too, the medical treatment was chosen because surgery was not possible due to the fibrous dysplasia of the skull base and radiotherapy would increase the risk of sarcomatous transformation. The patient received octreotide LAR 40 mg (intramuscular) monthly. The plasma growth hormone levels (post-oral glucose) decreased from an initial 60 ng ml21 to 23 ng ml21 1 month after first injection. In conclusion, the association of acromegaly with MCAS may pose a diagnostic and therapeutic challenge. MRI is vital in the evaluation of such patients for the delineation of the pituitary adenoma separate from the

86

skull base abnormality. CT is a useful adjunct preoperatively to delineate the foraminal compression.

References 1. Schwindinger WF, Levine MA. McCune-Albright Syndrome. Trend Endocrinol Metab 1993;7:238–42. 2. Lumbroso S, Paris F, Sultan C; European Collaborative Study. Activating Gsalpha mutations: analysis of 113 patients with signs of McCune-Albright syndrome--a European Collaborative Study. J Clin Endocrinol Metab 2004;89:2107–13. 3. Feldman F. Tuberous sclerosis, neurofibromatosis, and fibrous dysplasia. In: Resnick D, editor. Diagnosis of bone and joint disease. 4th edn. Philadelphia, PA: WB Saunder’s Company 2002:4792–843. 4. Weinstein LS, Shenker A, Gejman PV, Merino MJ, Friedman E, Spiegel AM. Activating mutations of the stimulatory G protein in the McCune Albright Syndrome. N Engl J Med 1991;325:1688–95. 5. Akintoye SO, Chebli C, Booher S, Feuillan P, Kushner H, Leroith D, et al. Characterization of gsp-mediated growth hormone excess in the context of McCune-Albright syndrome. J Clin Endocrinol Metab 2002;87:5104–12. 6. Chanson P, Dib A, Visot A, Derome PJ. McCune-Albright syndrome and acromegaly: clinical studies and responses to treatment in five cases. Eur J Endocrinol 1994;131:229–34. 7. Lipson A, Hsu T. McCune Albright syndrome associated with acromegaly: report of a case and review of literature. Johns Hopkins Med J 1981;149:10–4. 8. Premawardhana LDKE, Vora JP, Mills R, Scanlon MF. Acromegaly and its treatment in the McCune-Albright syndrome. Clin Endocrinol (Oxf) 1992;36:605–8. 9. Leet AI, Chebli C, Kushner H, Chen CC, Kelly MH, Brillante BA, et al. Fracture incidence in polyostotic fibrous dysplasia and the McCune-Albright syndrome. J Bone Miner Res 2004;19:571–7. 10. Kim KS, Rogers LF, Goldblatt D. CT features of hyperostosing meningioma en plaque. AJR Am J Roentgenol 1987;149:1017–23. 11. Daly BD, Chow CC, Cockram CS. Unusual manifestations of craniofacial fibrous dysplasia: clinical, endocrinological and computed tomographic features. Postgrad Med J 1994;70:10–6. 12. Cannavo S, Curto L, Fazio R, Paterniti S, Blandino A, Marafioti T, et al. Coexistence of growth hormone-secreting pituitary adenoma and intracranial meningioma: a case report and review of the literature. J Endocrinol Invest 1993;16:703–8. 13. Bunick EM, Mills LC, Rose LI. Association of acromegaly and meningiomas. JAMA 1975;240:1267–8. 14. Immenkamp M. Malignant change in fibrous dysplasia (author’s transl). Z Orthop IhreGrenzgeb 1975;113:331–43. 15. Bhansali A, Sharma BS, Sreenivasulu P, Singh P, Vashisth RK, Dash RJ. Acromegaly with fibrous dysplasia: McCuneAlbright Syndrome -- clinical studies in 3 cases and brief review of literature. Endocr J 2003;50:793–9.


Acknowledgment to Referees Editors would to thank all their colleagues have contributed their valuable time and effortin inreviewing The The Editors would like like to thank all their colleagues whowho have contributed their valuable time and effort reviewing manuscripts submitted to The British Journal of Radiology. Listed below are the names of manuscripts submitted to The British Journal of Radiology. Listed below are the names of referees of papersreferees submitted to papers 1submitted BJRand between 1 December BJRof between Decemberto2004 1 December 2005.2004 and 1 December 2005.

A Abernethy, L Adam, E Adams, J A’Hern, R Aird, E Albrecht, T Allan, R Allen, G Allisy-Roberts, P Al-Qaisieh, B Amin, Z Anbarasu, A Andreyev, J Ansorge, O Armstrong, P Ashleigh, R Ashton-Key, M Atchley, J

B Balogun, M Barber, P Barker, C Barker, D Barrington, S Barron, D Bearcroft, P Beavis, A Beggs, I Bell, J Belli, A-M Benham, J Birchall, D Blake, P Blease, S Bownes, P Boyle, G Brada, M Bradley, A Bradley, D Brettle, D Broderick, N Brown, P Bryant, P Budgell, G Buffa, F Burch, A Burling, D Burn, P Burrell, H Butteriss, D Byrne, J

C Callaway, M Campbell, R

Campbell, S Carr, R Carrington, B Carroll, N Carruthers, D M Casey, M Castellano, I Chalmers, N Chambers, R Chandy, J Chapman, A Chapple, C-L Chawla, T Chinn, R Clarke, S Cleveland, T Cochlin, D Cole, D Colligan, S Collins, C Collins, M A Collins, M C Connolly, D Conway, J Cook, G Cook, P Cooper, P Copley, S Coral, A Corr, C Cosgrove, D Cosgrove, V Cousins, C Cowan, N Cowling, M Crawley, C Crawley, T Crellin, A Crowe, P Curtis, J Czajka, J

D Dance, D Darby, M Darroudi, F Davies, M Davies, N Davison, P Dawson, P Deane, C Deehan, C Derchi, L Dhingsa, R Ditchfield, A Dixon, A Domjan, J


Dowling, A Downes, M Doyle, P Driver, D M Drury, A Duck, F Durante, M

E Elabassy, M Elford, J Elias, D England, R Evans, D Evans, J Evans, S

F Fairbairn, K Faithfull, S Farrugia, M Faulkner, K Flynn, T Fogelman, I Forbes, K Forsyth, L Fowler, J Fox, B Francis, I Freeman, A Freeman, S Fukuda, S

G Gawne-Cain, M Gaze, M Geh, I Geleijns, J Gentle, D George, C George, J Gibson, M Gillams, A Gillespie, J Gilligan, P Gishen, P Given-Wilson, R Glynne-Jones, R Goddard, T Goh, V Goldstone, K Gordon, A Goss, D Gould, D Greaves, S Green, R Greener, A

Grier, D Griffiths, S Grubnic, S Grundy, A Guest, P Guthrie, J A

H Hale, M Hall, C Hall, E Hall-Craggs, M Halligan, S Halpin, S Hanlon, R Hanson, M Harbinson, M Harden, S Hardman, J Hare, C Harrison, R Hart, D Hart, G Hartley, A Harvey, C Haslam, P Hatton, M Healy, J Heenan, S Heinz-Peer, G Heneghan, M Henson, J Heron, C Hide, G Hillier, J Hiorns, M Hirst, D Hoggard, N Holemans, J Hollaway, P Hoole, A Hopewell, J Hopkins, K Hounsell, A Hubscher, S Huda, W Hufton, A Hughes, D Hutchinson, C IIrving, H

J Jackson, A Jackson, J Jackson, S Jackson, S A Jaspan, T 87

Jayakrishnan, V Jobling, J Johnson, K Julian, W Julyan, P

K Kaanders, J Karani, J Kaufmann, P Kay, C Keane, A Kearney, S Keat, N Kelly, C Kerslake, R Kessar, P Kessel, D Keston, P Khaw, K-T Khoo, L Kinsella, D Kirby, M Koh, D-M Koller, C Kuker, W Kuntzsch, M

L Laitt, R Langton, C M Larkin, E Lee, C Lee, S Lenthall, R Leslie, M Lewis, M Lewis, M A Lewis-Jones, H Lim, A Litherland, J Livsey, J Locks, S Logan, M Lomas, D Lopez, C Lowdell, C Lowe, S Lucraft, H

M Mackenzie, A Maliakal, P Malone, J Malone, L Manning, D Marples, B Marshall, M Marshall, N Marshall, T Martin, C Martin, D Maskell, G Matson, M 88

Matthews, S Maxwell, A Mayles, H Mayles, W McCafferty, I McCall, J McCallum, H McCavana, J McGee, S McHugh, K McHugo, J McKenzie, A McLean, A McNee, S Meeson, S Mikhaeel, G Miles, K Miller, S Mitchell, F Mitra, D Mohammed, S Mohan, H K Mooney, R Moore, C Moores, M Morcos, S Morgan, A Morgan-Fletcher, S Moriarty, M Morrison, R Moss, H Moss, J Mothersill, C Mott, J Mountford, P Moussa, S Mueller-Klieser, W Muirhead, C Munro, A Murphy, D Murphy, P Murray, D

N Nahum, A Nakielny, R Nanda Kumar, E Negus, S Ng, C Ng, K K-C Niven, S Nolte-Ernsting, C Nunan, T Nutting, C

O O’Connor, R O’Donnell, P O’Donovan, N Odurny, A Offiah, C Ogunremi, T Old, S Olliff, J

Olliff, S O’Neill, P O’Reilly, G Ormerod, O Ostlere, S Owens, S

P Padhani, A Padley, S Page, A Paisley, E Paley, M Parker, G Parry, J Patankar, T Patel, U Patsios, D Payne, G Pelling, M Pereira, P Phillips, A Phillips-Hughes, J Pickles, T Pilcher, J Pilling, D Plant, G Plowman, P N Pretorius, P Price, P Prise, K Pullinger, R

Q Quaghebeur, G

R Raby, N Ralleigh, G Rampling, R Ramsdale, M Rankine, J Rawlings, D Reading, P Reddy, M Reek, C Rees, M Reliene, R Reynolds, J Rezvani, M Richardson, D Richenberg, J Ridley, N Riley, P Rimoldi, O Ritchie, D Roach, H Robbins, M Roberts, D Robertson, G Robertson, I Robinson, P J A Robinson, P S Robson, K

Rockall, A Rodgers, P Rogers, A Ross, P Rottenberg, G Rowbottom, C Rowell, N Rowland Hill, C Rust, A Rutherford, E Ryan, P Ryan, S

S Saada, J Saidlear, C Saifuddin, A Salim, F Saran, F Saunders, M Saunders, T Schofield, K Scholz, M Senior, R Shah, P Sharma, B Shaw, A Shaw, A S Sheridan, M Shorvon, P Shrimpton, P Sidhu, P Sikdar, T Silver, D Slater, A Slevin, N Smart, J Sminia, P Smith, S Sohaib, A Sprigg, A Staffurth, J Stanley, S Steeds, R Stevenson, G Stratford, M Strickland, N Sulkin, T Sutton, D Sykes, J

T Tait, D Tait, N Tan, L T Tattersall, D Tawn, E Taylor, A Taylor, B Taylor, D Taylor, P Taylor, R Taylor, S Temperton, D


Thomas, S Thompson, P Thomsen, H Thurston, J Tibballs, J Tins, B Traill, Z Travis, S Troughton, A Tsalafoutas, I Tuck, J Tung, K Turnbull, I Twyman, N

U Uberoi, R Uthappa, V

V van der Molen, A van Zeeland, B Vao´, E Varghese, A Vennart, W Vijayanathan, S Vinjamuri, S Vinnicombe, S

W Wade, P Wakeley, C Waldman, A Walker, A Waller, M Ward, S Wardman, P


Warren, R Warrington, A Watson, J West, R Weston, M Whitby, E Whitehouse, R Wilkins, J Wilkinson, J Wilkinson, L Williams, H Williams, J Winder, J Wittkop, B Wood, A Wood, C Woolhouse, I Workman, A

Worthy, S Wu, H

Y Yates, S Ying, M

Z Zoetelief, J

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February 2006 Volume 79 Issue 938

February 2006, Volume 79, Issue 938 DNA repair: therapeutic implications ● Patients’ perception of tests in the assessment of faecal incontinence ● Enhanced visualization and quantification of magnetic resonance diffusion tensor imaging using the p:q tensor decomposition ● PET-CT findings in surgically transposed ovaries ● An audit of imaging test utilization for the management of lymphoma in an oncology hospital: implications for resource planning? ● Image quality and breast dose of 24 screen–film combinations for mammography ● The effect of phantom type, beam quality, field size and field position on X-ray scattering simulated using Monte Carlo techniques ● Techniques for measurement of dose width product in panoramic dental radiography ● A comparison of three-field and four-field techniques in different clinical target volumes in prostate cancer irradiation using dose volume histograms: a prospective three-dimensional analysis ● A comparative evaluation of two head and neck immobilization devices using electronic portal imaging ● Excessive leakage radiation measured on two mobile X-ray units due to the methodology used by the manufacturer to calculate and specify the required tube shielding ● Improvements in dose homogeneity for tangential breast fields from a selection of combinations of library compensators ● Ruptured spinal dermoid cyst with disseminated intracranial fat droplets ● Colobronchial fistula: a late complication of childhood radiotherapy ● Ingested foreign body mimicking an appendicolith in a child ● Misleading positioning of a Foley catheter balloon ● An unusual cause and presentation of a pelvic mass ●

The British Journal of Radiology, 79 (2006), 91–93

COMMENTARY

DNA repair: therapeutic implications 1

S R MCKEOWN,

MA, PhD

and 2B JONES,

MSc, MD

1

School of Biomedical Sciences, University of Ulster at Coleraine BT52 1SA, Northern Ireland and Queen Elizabeth University Hospital, Birmingham B15 2TH, UK

2

Received 24 March 2005 Revised 8 June 2005 Accepted 13 July 2005 DOI: 10.1259/bjr/22946335 ’ 2006 The British Institute of Radiology

The goal of all anti-cancer treatments is to design strategies that are specific for tumours and non-toxic to the patient. Molecular targeting is now becoming a reality with new treatments designed to target processes that are thought to be tumour specific, or where there are quantitative differences in target expression between cancer and normal cells. On 3 March 2005, the Radiation and Cancer Biology committee held a meeting to discuss the targeting of DNA repair pathways, which are often defective in tumours. Prof. Steve Jackson (Cambridge University) started the programme with a discussion of some of the main DNA damage response (DDR) pathways. Repair in normal cells is a hugely efficient process, with individuals requiring repair of an estimated 1018 DNA lesions per day caused by reactive oxygen species alone. Most of the inherited cancer predisposition syndromes involve DDR dysfunction and similar mutations are often found in sporadic cancers. The specificity of DDR targeting agents comes from the need for the faster dividing tumour cells to repair DNA damage more quickly and efficiently than the mostly quiescent, or more slowly cycling normal cells. This may also be compromised by an already defective DDR pathway, which further reduces the ability of the tumour cells to repair efficiently, while being less critical in normal cells. He pointed out the clearly integrated nature of stress response in cells since DNA-PK, ATM and ATR have overlapping roles in DDR, transcriptional regulation, cell cycle control and cell death pathways. These processes are relevant not only to cancer therapy, but also in immune deficiency syndromes, neurodegenerative disorders, infertility, premature ageing and impaired telomere function. Many laboratory approaches were discussed during the day. Of particular relevance was the talk by Dr Kai Rothkamm (Gray Cancer Institute, Northwood) who discussed the potential uses of the cH2AX assay to measure double strand breaks (DSB). This relatively new assay allows quantitation of DSBs with more accuracy and at much lower doses than was possible previously. The British Journal of Radiology, February 2006

However, as with many assays there are inherent pitfalls as well as advantages of this method, and further work is required to characterize the assay completely. Dr Rothkamm identified possibilities for its use as a low dose exposure assay, using blood lymphocytes, since it is sufficiently sensitive to quantify exposure after diagnostic CT scans. It can also be used in studies of DDR inhibitors to quantitate responses and several speakers during the day reported uses for this assay. Prof. Jackson proposed that if certain types of cancer possess inherent DNA repair disorders then, in principle, inhibition of the remaining DDR mechanisms should lead to cell death more efficiently that can be achieved in normal cells where the full complement of repair enzymes is available. This theme was exemplified by several of the symposium speakers. Dr Niall Martin (KuDOS Pharmaceuticals, Cambridge) described two approaches to this strategy. In colorectal tumour cell lines with mis-match repair (MMR) defects, the response to standard cytotoxic agents such as temozolamide is enhanced when combined with inhibitors of PARP-1 (Poly (ADP-ribose) polymerase-1) – an enzyme critical to the early response to single strand breaks (SSB). This combination increases the yield of both SSB and DSB; the latter have been shown using the DSB specific cH2AX assay. MMR defects are not found in normal bone marrow cells, so that enhanced acute marrow toxicity is not to be expected. BRCA1/2 are also known to be key proteins involved in the cellular response to DDR. A significant number of breast tumours contain defects in BRAC1/2, including almost all inherited breast tumours. Inhibition of repair with PARP-1 caused a profound sensitization of BRCA1/2 deficient cells affecting G2/M checkpoint arrest, increased chromosome aberrations and tumour regression in the absence of other cytotoxic agents. This offers an exciting opportunity to control this relatively large subset of breast tumours. A similar approach, targeting DNA-PK inhibitors, cause preferential cell kill in ATM -/- cells, again with dramatic effects in vitro. 91

S R McKeown and B Jones

Prof. Penny Jeggo (Sussex University) described the role in DNA repair of ATM, and a small ATM interacting nuclease, called Artemis. Using quiescent fibroblasts, so that cell cycle differences did not confound the interpretation, she showed that for full restoration of DNA damage caused by some agents, times in the region of 72 h are needed. This is significantly longer than most reported repair studies and interestingly is longer than the time allowed between fractions in conventional radiotherapy; in part this may offer an explanation for the poorer DNA repair capacity in tumour vs normal cells. Although most (,90%) of the DNA repair occurs rapidly, the residual damage is significantly more difficult to deal with. The slower repair process appears to be dependent on the integrity of cell cycle checkpoint control. In addition, more severe lesions, using alpha particles, show a longer time to complete resolution of the damage. She showed evidence that Artemis is required for this process suggesting another potential drug target. Prof. Hilary Calvert (University of Newcastle) gave a keynote lecture on the current clinical trials involving DNA repair inhibitors. Resistance to methylating agents in many cells is caused by the repair enzyme alkylguanine alkyltransferase (O6AT). This enzyme can be inhibited using 6-benzylguanine (6BG) and 4 bromothenylguanine (Patrin). Unfortunately the clinically tolerated dose of Carmustine must be reduced threefold in combination with 6BG, whereas temozolamide is less affected by combination with Patrin. This suggests that O6AT plays an important role in normal tissue recovery. Phase 2/3 trials are currently determining whether there is an overall therapeutic benefit with this combination. A Phase 1 trial combining a PARP-1 inhibitor with temozolamide is just about to report and further trials are in the planning stages. Prof. Calvert discussed briefly the difficulties of setting up clinical trials in the molecular targeting era, where precise control of sample collection, storage and evaluation must be in place to identify the molecular profile of the tumour and its likely susceptibility to the treatment under investigation. Dr Stephany Veuger (Newcastle University) presented further work on PARP-1 inhibition. NF-kB is a stress inducible transcription complex that induces genes that control proliferation responses and suppress apoptotic cascades. Aberrant activation of NF-kB is common in tumours and recently it has been noticed that its activation in PARP-1 deficient cells is reduced. The involvement of these two proteins in the presence or absence of a potent PARP-1 inhibitor (AG14361) was investigated when cells were also exposed to 20 Gy ionizing radiation (IR). The data provided evidence that PARP-1 function is required for NF-kB activity following exposure to IR. The results suggested that potentiation of IR-induced radiosensitivity may be through inhibition of NF-kB rather than as a direct consequence of PARP-1 mediated inhibition of DNA repair. This result clearly has implications for rationale design of PARP-1 inhibitors in the treatment of cancer. Dr Paul Mullan (Queen’s University, Belfast) showed the power of an initial microarray screen to identify differences in BRCA1 competent and deficient cells. Dr Mullan and colleagues have identified a family of calcium binding proteins that are novel BRCA-1 92

repressed targets. S100A7 (psoriasin) is dependent on functional c-Myc and is also inducible by DNA damage in a BRCA-1 dependent manner. They linked this to a novel pathway of p27kip1 down-regulation that has previously been seen to be consistently down-regulated in BRCA1 mutated cells. The data have allowed identification of a novel pathway that could provide a target for molecular targeting agents. Targeting of DNA base excision repair was discussed by Dr Srinivasan Madhusudan (CRUK, Weatherall Institute of Molecular Medicine, Oxford). The multifunctional protein endonuclease HAP-1/APE-1/Ref-1 is involved in base excision repair and is implicated in the pathogenesis of several human tumours. Its overexpression is linked to both chemoresistance and radioresistance. Using a high throughput chemical screen, the Oxford group has identified KM09181 as a lead inhibitor of HAP-1 with an IC50 value of 3.5 mM. At non-toxic concentrations it causes significant potentiation of the cytotoxicity of a number of agents. This report is the first biological evidence for the direct targeting of this DNA repair enzyme. A series of novel PARP-1 inhibitors were described by Dr Esther Woon (University of Bath). Previously, they had identified 5-aminoisoquinolin-1-one (5-AIQ), which shows a wide range of therapeutic activity in vivo. Using the PARP-1 crystal structure, they designed a series of compounds similar to 5-AIQ, with the aim of identifying novel compounds with more potent PARP-1 inhibitory activity while retaining the excellent biopharmaceutical properties of 5-AIQ. A compound, 5-amino-3-methylisoquinolin-1-one (3-Me-5AIQ) was identified, which was 7 times more potent than 5-AIQ. A rather surprising result was reported by Dr S C Sak (CRUK, Leeds) who used immunohistochemistry to assess expression of two DDR proteins, APE-1 and XRCC1, in biopsy samples from 90 muscle invasive bladder tumours. High levels of these proteins correlated with survival after radical radiotherapy. On first reflection, high expression should protect tumours from IR. However, others factors may be invoked to explain this. Since the median patient age was 75 years, it is possible that depletion of natural radioprotectors, e.g. glutathione and other sulphydryl compounds, might allow more DSB damage to occur per unit dose with enhanced repair responses in patients who are cured. Another potential explanation could be the occurrence of enhanced misrepair in these patients. This enigmatic finding needs further investigation. The response of DDR pathways following exposure to low dose radiation (0–2 Gy) was discussed by Dr Susan Short (Gray Cancer Institute, Northwood). She reported the response of a number of genes in two cell lines, +/for low dose hypersensitivity (HRS). ATM signalling to downstream targets such as P53, CHK1 and CHK2 is functional at doses as low as 0.2 Gy. The induction of DSB, measured using cH2AX, appear to be linear with dose, but inhibition of DNA repair produces an exaggerated effect when using ATM inhibitors. DNAPK inhibitors have a lesser effect on low dose responses, but Rad51/BRCA2 mediated repair events may increase at doses below 1 Gy, which may be applicable to normal tissue responses during radiotherapy. The British Journal of Radiology, February 2006

Commentary: DNA repair

Overall, the workshop provided an excellent update on the progress of molecular targeting of DNA repair as a strategy for enhancing anti-cancer treatments. Significant progress has been made in recent years. The processes are better understood and the development of the cH2AX assay has allowed the interrogation of effects in the low clinically relevant range. New and better drugs are currently being tested and there is an expectation that these strategies will be successful in controlling at least a subset of solid tumour treatment responses where DDR pathways are already significantly compromised. There are some important caveats and implications to radiotherapy, including a theoretical risk of enhanced carcinogenesis in normal tissues; malignant transformation assays should be performed to investigate the potential magnitude of this risk and whether there is a synergy with concomitant radiation and/or chemotherapy. The use of proton beam radiotherapy might allow these agents to be used more safely due to

The British Journal of Radiology, February 2006

the reduced collateral radiation of normal tissues; intensity-modulated radiotherapy (IMRT), associated with a dose bath effect of low to medium dose in surrounding normal tissues would need very careful assessment, although dose escalation may not be so necessary in the presence of DNA repair inhibitors. It is also self-evident that the extant mathematical models of repair used in radiotherapy might require specific changes to accommodate the mechanisms described in this paper. Changes in radiotherapy fractionation (dose per fraction and interfraction interval) might follow the determination of precise repair capacity in tumours relative to normal tissues. Robust laboratory, clinical and analytical methodology is necessary in order to determine whether enhanced cure rates and an improved therapeutic index can be achieved by exploitation of altered repair systems in some types of cancer. Based on the content of this meeting, the prospects seem good.

93


Patients’ perception of tests in the assessment of faecal incontinence 1

M DEUTEKOM, PhD, 2M P TERRA, MD, 1M G W DIJKGRAAF, PhD, 2A C DOBBEN, 3 1 PhD, G E BOECKXSTAENS, MD, PhD and P M M BOSSUYT, PhD

MSc,

2

J STOKER,

MD,

1

Department of Clinical Epidemiology and Biostatistics, 2Department of Radiology and 3 Department of Gastroenterology from the Academic Medical Center, Amsterdam, The Netherlands ABSTRACT. The objective of this study was to evaluate patient perception of endoanal MRI compared with defecography and anorectal functional testing in the workup of patients with faecal incontinence. Consenting consecutive patients underwent a standard testing protocol consisting of endoanal MRI, defecography and anorectal function combination. Patient experience was evaluated with a self-administered questionnaire, addressing anxiety, embarrassment, pain and discomfort, each measured on a 1 (none) to 5 (extreme) point-scale. Patients were also asked to rank the three tests from least to most inconvenient. Statistical analysis was performed with parametric tests. Data from 211 patients (23 men; mean age 59 years (SD¡12)) were available. MRI had the lowest average score for embarrassment and discomfort (1.6) and defecography the highest (1.9 and 2.0, respectively) (p,0.0001, tested with general linear model for related samples). The average pain score was lowest for MRI (1.4) and highest for the anorectal function combination (1.7) (p,0.0001). Level of anxiety was highest for MRI (1.6 versus 1.4; p50.03). MRI was scored as least inconvenient by 69% of patients. Endoanal MRI was scored as least inconvenient. However, the differences in patient burden between the three diagnostic tests were small and absolute values were low for all tests. Patient perception will not be a key feature in determining an optimal diagnostic strategy in faecal incontinence.

Faecal incontinence is defined as recurrent uncontrolled passage of faecal material at an inappropriate time or in an inappropriate place more than twice a month [1]. The reported prevalence values range from 1.4% in the general population [2] to 46% in institutionalized elderly [3, 4]. It is possible that the real prevalence is even higher than reported as faecal incontinence is associated with high social stigma and people do not easily seek help for this disorder out of embarrassment [5, 6]. Childbearing injuries (sphincter and/or pudendal nerve damage) and prior anorectal surgery (sphincter trauma) are the main causes of faecal incontinence [7, 8]. Diagnostic tools are used to determine the exact cause of the faecal incontinence complaints and aim to guide future therapy. In the evaluation of faecal incontinence clinicians can use a large variety of diagnostic tools, including anorectal function tests and anorectal imaging techniques after medical history and physical examination. Treatment guidance appears to be problematic as there exists debate in the value of diagnostic tests with respect to treatment outcomes [9]. It has been shown that the existence of sphincter atrophy has a negative predictive value on the success of sphincter repair [10]. About 10 years ago, endoluminal MRI of the rectum and the This research was supported by grant 945-01-013 of the Netherlands Organization for Health Research and Development.

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Received 26 May 2005 Revised 15 June 2005 Accepted 22 June 2005 DOI: 10.1259/bjr/63269033 ’ 2006 The British Institute of Radiology

anus was introduced [11]. It appeared that, although endoanal MRI was comparable with endoanal ultrasound for identifying defects of the sphincter, only endoanal MRI could reveal thinning of the external sphincter reflecting muscle atrophy [12]. Its ability to identify sphincter defects as well as external sphincter atrophy makes endoanal MRI a likely candidate for a diagnostic strategy to guide treatment decisions in patients with faecal incontinence. Yet an optimal diagnostic strategy should also try to minimize patient burden, as extensive testing may be taxing to patients. It has been shown in a review that anxiety-related reactions occur in approximately 4% to 30% of patients undergoing MRI, ranging from apprehension to severe reactions that interfere with the performance of the test [13]. These findings were from studies using noninvasive MRI techniques. It could be hypothesized that the use of an endoluminal coil could be even more bothersome for patients, but there exist no data to either refute or confirm this hypothesis. We designed a study to evaluate and compare the patient burden of diagnostic tests used in the work-up of patients with faecal incontinence as part of a clinical cohort study aiming to identify prognostic factors for treatment success by physiotherapy. We studied the perceived burden of endoanal MRI, defecography and an anorectal function test combination consisting of anorectal manometry, pudendal nerve The British Journal of Radiology, February 2006

Patients’ perception of tests

terminal motor latency, rectal capacity measurement, anal and rectal sensitivity measurement and endoanal ultrasound.

Materials and methods The clinical cohort study had started in December 2001. By February 2004, 240 consenting consecutive patients with faecal incontinence visiting one of 16 participating medical centres (Dutch) were included in the cohort study. The medical ethics committees of all participating hospitals approved the study. Patients were identified by surgeons, gastroenterologists and a gynaecologist participating in the large diagnostic cohort study. Patients were referred to these physicians by general practitioners or by physicians who were not participating in the cohort study. After receiving signed informed consent from patients, data concerning medical history were collected by physicians. All participating physicians used the same structured forms for medical history. Patients were questioned about the duration of their faecal incontinence complaints. The severity of faecal incontinence was assessed by means of an incontinence scale developed by Vaizey [14]. This scale contains items about the type (gas, fluid, solid) and frequency of incontinence and additional items addressing social invalidation, the need to wear a pad or plug, the use of constipating medication and the presence of urge incontinence. The total score on the Vaizey scale ranges from 0 (complete continence) to 24 (complete incontinence). Inclusion criteria for the cohort study were the existence of faecal incontinence complaints for 6 months or more, a Vaizey incontinence score of at least 12, and failure of conservative treatment, based on diet recommendations and/or antidiarrhetics. Excluded were patients aged below 18 years, patients diagnosed less than 2 years ago with an anorectal tumour and patients with a previous ileoanal or coloanal anastomosis. As the clinical cohort study investigated the treatment effect of physiotherapy, patients with chronic diarrhoea (always fluid stools, three or more times a day), overflow incontinence, proctitis, soiling (leakage of faecal material out of the anus after normal defecation leading to perineal eczema) and rectal prolapse were also excluded from participation. Some patient categories were excluded from one or more tests. Defecography was not performed in females younger than 45 years without sterilization, except on indication based on clinical symptoms (e.g. lower abdominal pain and/or false urge to defecate) and clinical findings (e.g. symptoms of prolapse) because of the cumulative radiation doses of somatic and genetic effects. Before the MRI examination, patients were questioned about claustrophobia and had to complete a questionnaire comprising exclusion criteria for a MRI examination, such as a pacemaker, claustrophobia and other contraindications. Patients with a pacemaker were excluded from the MRI examination, a pacemaker being an absolute contraindication for MRI. Only the data of patients who experienced all three test sessions were analysed in this study. The British Journal of Radiology, February 2006

Diagnostic tests Patients underwent three diagnostic sessions: one with endoanal MRI, a second with defecography and a third with a combination of anorectal function tests consisting of anorectal manometry, pudendal nerve terminal motor latency, rectal capacity measurement, anal and rectal sensitivity measurement, and endoanal ultrasound. None of the patients received sedation for any of the tests. The decision to evaluate the burden of the latter anorectal function test combination was made because these tests are usually performed in a single testing session and we expected that patients would find it difficult to differentiate the tests. Logistical considerations prevented us from randomizing the order of the tests. All diagnostic tests were performed according to a standard procedure that had been established during joint meetings of the research group members of all participating hospitals. Not all centres were equipped to perform all tests, therefore not all patients could be tested at a single site. Prior to testing, all patients received standard written information concerning the tests.

MRI Endoanal MRI visualizes the muscles of the pelvic floor. Endoanal imaging was performed with 1 T or 1.5 T MR (General Electric Horizon Echospeed; General Electric, Milwaukee, IL; Philips Gyroscan ACS-NT; Philips Medical Systems, Best, The Netherlands) clinical closed bore units and a dedicated endoanal coil with a diameter of 18 mm. All patients were asked to fast 4 h prior to the MR examinations to minimize artefacts from bowel peristalsis. In all hospitals except one, the patients were injected intramuscularly with an antiperistaltic drug to reduce bowel motion before the start of imaging. No intravenous contrast medium was used. The endoanal coil was covered with a condom and, after lubrication, inserted into the anal canal with the patient in a left lateral position. After positioning of the endoanal coil, the patients were turned to the supine position and moved into the magnet. The patient was instructed not to squeeze to prevent artefacts of movement. The scan period took on average 20 min. As this test was performed as part of a larger study, patients were also studied with a phased array coil in the same session after removal of the endoanal coil. No intravenous contrast medium was used and no dynamic sequences were performed with external phased array coil MRI. The burden expressed by the patients for MRI was the burden for the combination of endoanal and phasedarray MRI. The total duration for this combination was around 40 min.

Defecography Defecography allows an evaluation of the movements of the rectum, insufficiency of the sphincter, presence or absence of rectoceles, enteroceles and intussusceptions. Patients were instructed to drink contrast medium diluted in water prior to the examination. The test started with the patient in left decubital position. Through an injection, pistol barium paste (200–300 ml barium sulphate prepared 95

M Deutekom, M P Terra, M G W Dijkgraaf et al

by the hospital pharmacy or Evacu-Paste (E-Z-EMH Inc., Westbury, NY)) was injected manually into the rectum. In female patients, amidotrizoide acid 50% gel was also injected via a syringe into the vagina. The perineum was located with amidotrizoide acid 50% gel solution or located by a catheter with leadmark. Subsequently, the entire X-ray table was tilted upright 90˚ and the patient was seated on a specially developed radiolucent defecography chair. Defecography took approximately 15 min (room time). After the test was performed, the patient was instructed to drink extra to eliminate the contrast.

Anorectal function test combination All tests were performed in left lateral position with hips flexed to 90 ˚. Anal manometry evaluates the muscular contraction and relaxation of the anal sphincters by the measurement of pressures in the anal canal. Anal manometry took place according to the solid-state or water perfused technique, without or with sleeve. The catheter (Konigsberg Instrument Inc., Pasadena, CA; Medtronic, Skolvunde, Denmark; Dentsleeve Pty Ltd, Parkside, Australia) was introduced and stabilized in the anal sphincter complex. After positioning of the catheter, the basal sphincter pressure, maximum squeeze pressure and rectal anal inhibitory reflex were measured. Pudendal nerve terminal motor latency determines the integrity of the pudendal nerve. The finger with a glovemounted St Mark’s Hospital electrode (Dantec; Skovlunde, Denmark) was inserted into the rectum. The pudendal nerve was electrically stimulated (supra maximum stimulus of 0.05 ms) on each side near the ischial spine. With rectal and anal sensitivity measurements the threshold sensation of the rectum and anus was determined, respectively. The stimulation electrode (Dantec Keypoint, Skovlunde, Denmark) was mounted on a catheter and introduced into the rectum. A constant current was increased gradually to a maximum of 20 mA. The same procedure was performed in the anus to determine the threshold sensation of the anus. The capacity measurement of the rectum was performed by introducing a single use urinary catheter (female, 14 Ch) with a latex balloon tied to the end, covered with a lubricant and connected to a 50 ml syringe, into the rectum. The balloon catheter was inflated with air in gradual increments of 50 ml until the maximum tolerable volume was reached. The minimal rectal sensation perceived (sensory threshold), the volume associated with the initial urge to defecate (urge sensation) and the volume at which the patient experienced discomfort and an intense desire to defecate (the maximal tolerated volume) were determined. Endoanal ultrasound was performed with an ultrasound scanner (3535 Bruel and Kjaer, Gentfofte, Denmark; SDD-2000 Multiview Aloka, Tokyo, Japan) with radial endoscopic probe and a 7.5 MHz transducer. The probe was covered with a condom and, after application of a lubricant, introduced into the anal canal with the patient in left lateral or prone position. The probe was slightly withdrawn so all the different levels of the anal sphincter complex could be visualized. The total duration of the anorectal function test combination was between 30 min and 55 min. 96

Test questionnaire The self-administered questionnaire was handed out by a physician before the first test was performed. Patients were requested to take the questionnaire home and to complete the questionnaire after their last test. One researcher (MD) collected all completed questionnaires and contacted patients when no questionnaires were returned. When necessary, extra questionnaires were sent out. The questionnaire consisted of three modules. First, a standard formatted Likert scoring module was used with four items concerning pain, embarrassment, discomfort, and anxiety. The first three items have previously been used in a study of the acceptance of CT colonoscopy by patients [15]. Based on literature data, we added anxiety as the fourth item [13, 16, 17]. Responses were scored on a five-point scale with 1 indicating ‘‘none’’ and 5 indicating ‘‘extreme’’. By adding the item scores, an overall burden score was determined. Second, a comparative assessment module was used, forcing patients to rank the different tests from least to most inconvenient. Finally, a behavioural intent module was used by asking patients whether or not they, if opportune, would recommend each test to friends or relatives. The different modules were collated into one comprehensive questionnaire.

Statistical analyses The general linear model for related samples was used to compare the burden of the different tests. When a statistical difference was found, paired t-tests were used as post hoc tests. Subgroup analyses were performed based on age, duration of faecal incontinence (using a median split) and gender. Unpaired t-tests were used to test for differences between groups of patients with respect to sum burden scores. We also investigated with unpaired t-tests if burden values were different depending on whether patients received their tests in a single centre or in multiple centres. We analysed whether test order and the time lag between date of last test and date of completed questionnaire affected experienced burden, using Pearson correlation coefficients. Additional analyses were performed to study the association between the subjective ranking of a test and the burden variables. For each of the three tests, patients were categorized according to position of that test in their inconvenience ranking. We used analysis of variance to examine differences in the amount of burden between the different patient groups. p-values below 0.05 were considered to represent a statistically significant difference.

Results Patient characteristics From the 270 questionnaires distributed, 240 questionnaires were returned during the study period (response rate: 89%). Data were missing for one or more tests in 29 patients for various reasons: last test occurred after completion of the questionnaire (n515), contraindication (n58), claustrophobia (n53) or unknown (n53). For 211 (23 male; 188 female) patients all test The British Journal of Radiology, February 2006

Patients’ perception of tests Table 1. Order of testing Test order

Frequency

Anorectal function combination, defecography, MRI Anorectal function combination, MRI, defecography Defecography, anorectal function combination, MRI MRI, defecography, anorectal function combination Other

18 (17%) 14 (13%) 32 (30%) 32 (30%)

12 (11%)

data were available and could be analysed. These patients had a mean age of 59.2 (SD¡12.2) years, duration of incontinence 8.5 (SD¡8.4) years and Vaizey incontinence score of 18.0 (SD¡3.1).

Order and timing of the tests Information on the date of testing was available in 157 patients (74%). The mean duration between the first and last test was 62 days (SD¡92). Many tests were performed on the same day. The mean time between last test and completion of the questionnaire was 27 days (SD¡50). As the exact testing times were absent from a number of patient records, the exact test order could be derived for 108 patients (Table 1).

Test burden The reported burden of testing was low for all three tests, with average burden scores in the 1 to 2 range on all four items (Figure 1). Significant between-test differences were noted for embarrassment, pain and discomfort as well as for the total burden sum score. For embarrassment, discomfort and total burden, MRI had the lowest average score (1.56, 1.62 and 6.16, respectively) and defecography the

highest (1.92, 2.00 and 6.85, respectively) (all p,0.001). MRI also scored lowest regarding pain (1.38), whereas the highest pain score was observed for the anorectal function combination (1.73) (all p,0.001). The level of anxiety between tests also reached significance (p50.013) with higher values for MRI (1.6) compared with for defecography (1.4) and anorectal function tests (1.4). Younger patients (below 59 years) had a significantly higher total burden sum score for MRI (6.6 versus 5.7), defecography (7.4 versus 6.3) and anorectal function tests (7.1 versus 6.1) than older patients (Table 2). No differences with respect to the total burden sum scores of the three tests were observed between subgroups characterized by gender or duration of incontinence. Whether patients received their tests in a single centre (63%) or in multiple centres (37%) did not influence experienced burden. Despite these low average scores, a group of patients (24%, n551) reported on at least one test item (anxiety, embarrassment, pain, or discomfort) a high burden score (4 or 5). One or more items of MRI were given a high burden score by 12% (n525) of all patients; this percentage was 16% (n533) for defecography and 12% (n526) for anorectal function test combination. Patients reporting a high burden score for at least one of the items of MRI and anorectal function combination were significantly younger then patients who did not report a high burden score (55 years versus 60 years (p50.045) and 54 years versus 60 years (p50.015), respectively). There were no other significant associations between medical history and the group of patients that gave a high burden score on at least one item of a test. In a subset of 137 we analysed the effect of time lag between last test and completed questionnaire. No relationship was observed between total burden of any of the tests and time-lag (MRI (r520.13; p50.14); anorectal function tests (r520.07; p50.44); defecography (r520.002; p50.98)). Order of testing was present in 108 patients. Analysis showed that test order did not influence the amount of

Figure 1. Burden scores of the three tests in faecal incontinence with respect to pain, embarrassment, discomfort, anxiety and sum burden. *Difference between 3 tests (p,0.05). **Difference between 3 tests (p,0.001). Values indicate mean and 95% confidence interval; n5211. Post hoc tests showed significantly lower burden scores for MRI compared with the combination of anorectal function tests (pain and sum burden) and defecography (embarrassment, discomfort and sum burden). The British Journal of Radiology, February 2006

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M Deutekom, M P Terra, M G W Dijkgraaf et al Table 2. Subgroup analyses on total burden

Gender Age (years) Duration of incontinence (years) Site

Male Female , 59 . 59 ,5 .5 Single Multiple

MRI Total burden

Anorectal function c. Total burden

Defecography Total burden

6.1 6.1 6.6 5.5 6.2 6.0 6.0 6.2

7.0 6.6 7.0 6.3 6.7 6.5 6.9 6.2

6.4 6.9 7.4 6.3 6.9 6.9 6.9 6.9

p50.68 p,0.01 p50.67 p50.71

experienced burden (MRI (p50.36); anorectal function tests (p50.83); defecography (p50.63)). Only a small number of patients would not recommend one of the tests to a friend or relative: 7 for MRI (3.3%), 12 for defecography (5.6 %), and 6 for the anorectal function test combination (2.8%). Reasons for not advising MRI were possibly anxious reactions (n54), fear of loss of stool (n51), headache (n51) and unknown (n51). Defecography was not advised for various reasons: the dislike of the ingestion of the contrast medium before defecography (n53), experienced pain (n55), unclearness aboutuseofresults(n52),anxiety(n51)andthelackofprivacy (n51).Reasons fornotrecommendingtheanorectalfunction testcombinationwerepain(n55)andlongduration(n51). Not all patients responded to the ranking question, therefore analyses were done on the 174 respondents (82%). On the ranking question MRI scored best, with 120 (69%) patients scoring MRI as least inconvenient (Figure 2). Further analysis of all three tests revealed an association between the position in the ranking question (from least to most inconvenient) and the reported burden (Figure 3). Higher rankings (more inconvenient) corresponded with a significant higher burden sum score for that test (MRI: p,0.001, anorectal function combination: p50.03, and defecography: p,0.001).

Discussion This study investigated and compared the burden of endoanal MRI to defecography and the anorectal function test combination. Although endoanal MRI was associated with the highest level of anxiety, MRI was

p50.73 p50.01 p50.57 p50.16

p50.28 p50.02 p50.92 p50.94

found to have the lowest average scores for pain, embarrassment, discomfort and total burden. MRI also did well on the ranking question, with almost 70% of all patients scoring MRI as least inconvenient. Despite the significant differences in burden between tests, we should note that the differences were small and that absolute levels of burden were low for all tests. There existed a group of patients (24%) that reported a high score on at least one aspect of a test. These patients were on average younger, but it appeared to be impossible to identify on this group of patients basis of medical history. Only a small percentage of patients would not recommend one of the tests to a friend or relative. It could be hypothesized that larger differences in perceived burden exist across subgroups. We therefore performed a series of subgroup analyses defined by gender, age, and duration of incontinence. Younger patients (below 59 years) reported a significantly higher total burden sum score for all three tests. This finding could possibly be explained by a diminished pelvic floor sensory enervation in the older patient population or to less anxiety or embarrassment related to the procedure itself. Gender, duration of faecal incontinence complaints, order of tests, single or multiple site testing and location played no significant role in the amount of perceived burden of the tests. A number of potential limitations of this study should be taken into account. The obtained results were derived from data of patients voluntarily seeking help. It is possible that there exists a group of patients who do not request medical care, as they are less willing to undergo diagnostic testing, probably having higher burden scores

Figure 2. Inconvenience ranking of three tests in faecal incontinence. Proportion of patients reporting the test to be most inconvenient (black), least inconvenient (white) or in between (light grey).

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Patients’ perception of tests

Figure 3. Burden sum scores and inconvenience ranking. Values indicate mean and 95% confidence interval.

if ever tested. As patients seeking medical attention may suffer more severely from their complaints they could downgrade the burden of the testing in comparison with the burden of their illness. The observed imbalance between men and women in our study is not due to a form of selection bias but is inherent to the disorder of faecal incontinence [18]. We were forced to combine five tests into an anorectal function combination, as these tests are usually performed in a single testing session. When designing the study we learned that patients found it difficult to differentiate the tests during the testing sequence. We anticipated that in between measurement might interfere with the experience of the testing sequence and decided to rely on a post hoc assessment of the overall burden of the combination of tests. Unfortunately, this prevents us from making separate statements on endosonography. As endosonography and endoanal MRI produce comparable information, it would be interesting to compare the burden of these two imaging modalities. Although some have questioned the role of defecography in the diagnostic work up of faecal incontinence, our research group had decided to include this diagnostic modality within the cohort study. Some authors have underscored the importance of the role of defecography for accurately diagnosing intussusceptions and anterior rectoceles [19] or for determining the aetiology of outlet obstruction symptoms in patients with combined faecal incontinence [20]. In a suggested work-up of faecal incontinent patients by Felt [21], defecography was one of the components of the diagnostic procedures. MRdefecography is primarily employed in patients with prolapse or constipation, while the role of MR-defecography in incontinent patients is unclear. For this reason MR-defecography was not part of this diagnostic cohort study evaluating current practice. Another possible limitation is the non-random test order. The order in which tests were offered to patients varied considerably, but the results of our analysis showed that test order did not significantly affect experienced burden. The British Journal of Radiology, February 2006

It has been shown that past experience with testing can influence the perception of patients of a test [22–24]. It is unusual for patients to undergo repeated testing, so we expected very few of our patients to have undergone one or more of these tests previously. The low prevalence and the lack of data prevent us from exploring explicitly any bias due to prior experience. We believe that in our questionnaire study total, bias is kept to a minimum. Non-response bias is negligible as we achieved a response rate of 89%. The questionnaires were selfadministered so there is no potential for interviewer bias. Response bias was minimized by assuring anonymity of the patient. Furthermore, the questionnaires were handed out by a physician, but patients were requested to complete the questionnaire at home. Finally, due to the subject of the questionnaire we did not expect patients to respond in a sociably desirable manner. We have tried to standardize the information given to patients by handing out a written information sheet prior to testing. However, we cannot claim that all patients received exactly the same oral information by their specialists. To our knowledge this is the first study to investigate the patient burden of endoanal MRI, defecography and anorectal function test combination. The burden of MRI has been studied before, mostly with respect to patient anxiety. In a review by Melendez et al [13] it has been shown that anxiety-related reactions occur in approximately 4–30% of patients undergoing MRI. In this study, three patients did not undergo an MRI because of claustrophobia. None of the other patients became anxious up to a level that the test could not be executed, and none of the examinations had to be discarded because of motion artefacts, also associated with high patient anxiety in the past [17, 25–27]. The percentage of patients reporting high anxiety levels was low in comparison with other studies. One possible reason for this could be that anxiety was measured after the test had been performed. Various studies have reported lower anxiety levels post-MRI 99

M Deutekom, M P Terra, M G W Dijkgraaf et al

compared with pre-MRI [16, 25, 27]. Another possible explanation for the lower number of patients with anxiety reactions could lie in the fact that patients suffered from their incontinence for a long duration and were not afraid that MRI would reveal a certain malignant disease. Studies have shown that test anxiety could result from insecurity about what the test would reveal [16, 28]. Although the amount of anxiety of MRI in this study was low in comparison with earlier studies, the observed values were slightly higher than those for defecography and the anorectal function combination. MRI scored better with respect to other variables than defecography and the anorectal function combination. Small but significantly lower scores for MRI were seen for pain, embarrassment and discomfort. Total burden sum score was also significantly lower for MRI. Because all tests were performed as part of a larger study designed to evaluate the diagnostic performance of these tests, patients received phased-array MRI as well as endoanal MRI in a single session. The present study addressed the burden of that total MRI session. It can be expected that the burden for a diagnostic session with endoanal MRI only would be even somewhat smaller, because of the shorter duration of this single procedure. Overall, MRI was preferred more often than defecography and functional testing, with 120 (69%) patients scoring MRI as least inconvenient. For every test we observed a significant relationship between the given inconvenience rank and the burden sum score. Patients ranking a test as least inconvenient reported significantly less burden than patients who ranked this test as most inconvenient. We feel confident in concluding that the burden sum score, based on a combination of embarrassment, pain, anxiety and discomfort, is a reflection of relative inconvenience. The observed relation between the burden values and the ranking question supports the construct validity of this short and apprehensive questionnaire. In summary, in this study, set up to investigate the burden of diagnostic tests used in the assessment of faecal incontinence, we found significant differences between tests, with MRI scoring significantly better than defecography and the anorectal function combination. As the differences were small and the average burden values were low for all tests, we find it safe to say that the role of burden of testing in the search for an optimal strategy in faecal incontinence will be limited. The preferred diagnostic pathway will most likely be based on maximizing diagnostic accuracy at acceptable costs. Efforts to collect more information on test accuracy and costs are underway.

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5. Mavrantonis C, Wexner SD. A clinical approach to fecal incontinence. J Clin Gastroenterol 1998;27:108–21. 6. Jorge JM, Wexner SD. Etiology and management of fecal incontinence. Dis Colon Rectum 1993;36:77–97. 7. Kamm MA. Obstetric damage and faecal incontinence. Lancet 1994;344:730–3. 8. Toglia MR. Pathophysiology of anorectal dysfunction. Obstet Gynecol Clin North Am 1998;25:771–81, vi. 9. Prather CM. Physiologic variables that predict the outcome of treatment for fecal incontinence. Gastroenterology 2004;126 (Suppl.):s135–s140. 10. Briel JW, Stoker J, Rociu E, Lameris JS, Hop WCJ, Schouten WR. External anal sphincter atrophy on endoanal magnetic resonance imaging adversely affects continence after sphincteroplasty. Br J Surg 1999;86:1322–7. 11. Stoker J, Rociu E. Endoluminal MR imaging of anorectal diseases. J Magn Reson Imaging 1999;9:631–4. 12. Rociu E, Stoker J, Eijkemans MJ, Schouten WR, Lameris JS. Fecal incontinence: endoanal US versus endoanal MR imaging. Radiology 1999;212:453–8. 13. Melendez JC, McCrank E. Anxiety-related reactions associated with magnetic resonance imaging examinations. JAMA 1993;270:745–7. 14. Vaizey CJ, Carapeti E, Cahill JA, Kamm MA. Prospective comparison of faecal incontinence grading systems. Gut 1999;44:77–80. 15. Van Gelder RE, Birnie E, Florie J, Schutter MP, Bartelsman JF, Snel P, et al. A comparison of patient preference of CT colonography and colonoscopy: a five-week follow-up study. Radiology 2004;233:328–37. 16. Katz RC, Wilson L, Frazer N. Anxiety and its determinants in patients undergoing magnetic resonance imaging. J Behav Ther Exp Psych 1994;25:131–4. 17. MacKenzie R, Sims C, Owens RG, Dixon AK. Patients’ perceptions of magnetic resonance imaging. Clin Radiol 1995;50:137–43. 18. Henry MM. Pathogenesis and management of fecal incontinence in the adult. Gastroenterol Clin North Am 1987;16:35–45. 19. Jorge JM, Habr-Gama A, Wexner SD. Clinical applications and techniques of cinedefecography. Am J Surg 2001;182:93–101. 20. Wiersma TG, Mulder CJ, Reeders JW. Dynamic rectal examination: its significant clinical value. Endoscopy 1997;29:462–71. 21. Felt-Bersma RJF, Cuesta MA. Fecal incontinence 1994 which test and which outcome. Neth J Med 1994;44:182–8. 22. Mueller PR, Biswal S, Halpern EF, Kaufman JA, Lee MJ. Interventional radiologic procedures: patient anxiety, perception of pain, understanding of procedure, and satisfaction with medication–a prospective study. Radiology 2000;215:684–8. 23. Murphy KJ, Brunberg JA. Adult claustrophobia, anxiety and sedation in MRI. Magn Reson Imaging 1997;15:51–4. 24. Thorp D, Owens RG, Whitehouse G, Dewey ME. Subjective experiences of magnetic resonance imaging. Clin Radiol 1990;41:276–8. 25. Dantendorfer K, Amering M, Bankier A, Helbich T, Prayer D, Youssefzadeh S, et al. A study of the effects of patient anxiety, perceptions and equipment on motion artifacts in magnetic resonance imaging. Magn Reson Imaging 1997;15:301–6. 26. Sarji SA, Abdullah BJ, Kumar G, Tan AH, Narayanan P. Failed magnetic resonance imaging examinations due to claustrophobia. Australas Radiol 1998;42:293–5. 27. Quirk ME, Letendre AJ, Ciottone RA, Lingley JF. Anxiety in patients undergoing MR imaging. Radiology 1989;170: 463–6. 28. Gebbensleben B, Rohde H. Anxiety before gastrointestinal endoscopy–a significant problem? Dtsch Med Wochenschr 1990;115:1539–44.



Enhanced visualization and quantification of magnetic resonance diffusion tensor imaging using the p:q tensor decomposition 1,2,3

A PENÃ, PhD, 1,3H A L GREEN, MBChB, 3T A CARPENTER, 2,3 MChir, FRCS, FMedSci and J H GILLARD, BSc, MD, FRCR

PhD,

1,2,3

S J PRICE,

FRCS,

1,3

J D PICKARD,

Departments of 1Neurosurgery, 2Radiology and the 3Wolfson Brain Imaging Centre, Addenbrooke’s Hospital and the University of Cambridge, Cambridge CB2 2QQ, UK ABSTRACT. Many scalar measures have been proposed to quantify magnetic resonance diffusion tensor imaging (MR DTI) data in the brain. However, only two parameters are commonly used in the literature: mean diffusion (D) and fractional anisotropy (FA). We introduce a visualization technique which permits the simultaneous analysis of an additional five scalar measures. This enhanced diversity is important, as it is not known a priori which of these measures best describes pathological changes for brain tissue. The proposed technique is based on a tensor transformation, which decomposes the diffusion tensor into its isotropic (p) and anisotropic (q) components. To illustrate the use of this technique, diffusion tensor imaging was performed on a healthy volunteer, a sequential study in a patient with recent stroke, a patient with hydrocephalus and a patient with an intracranial tumour. Our results demonstrate a clear distinction between different anatomical regions in the normal volunteer and the evolution of the pathology in the patients. In the normal volunteer, the brain parenchyma values for p and q fell into a narrow band with 0.976,p,1.063 6 1023 mm2 s21 and 0.15,q,1.08 6 1023 mm2 s21. The noise appeared as a compact cluster with (p,q) components (0.011, 0.141) 6 1023 mm2 s21, while the cerebrospinal fluid was (3.320, 0.330) 6 1023 mm2 s21. In the stroke patient, the ischaemic area demonstrated a trajectory composed of acute, subacute and chronic phases. The components of the lesion were (0.824, 0.420), (0.884, 0.254), (2.624, 0.325) at 37 h, 1 week and 1 month, respectively. The internal capsule of the hydrocephalus patient demonstrated a larger dispersion in the p:q plane suggesting disruption. Finally, there was clear white matter tissue destruction in the tumour patient. In summary, the p:q decomposition enhances the visualization and quantification of MR DTI data in both normal and pathological conditions. Magnetic resonance (MR) diffusion tensor imaging (DTI) is a technique which allows the in vivo measurement of water diffusion in biological tissues from which tissue microstructure can be inferred [1–5]. It has been used successfully to investigate a number of neurological disorders that involve the disruption of white matter fibres including schizophrenia [6], head injury [7], multiple sclerosis [8] and stroke [9, 10]. In addition, DTI data can be used with a set of computational techniques called ‘‘tractography’’ [11] to reconstruct in vivo white matter tracts in the human brain, which is a very promising field, for example, to investigate their disruption due to an expanding tumour [12]. Diffusion is properly described by a high-dimensional mathematical quantity called a tensor. A tensor represents the generalization of scalars and vectors and as such, it contains more information than these. In three dimensions a scalar has one element, a vector three elements and a tensor nine elements. In order to quantify pathological changes in the diffusion tensor, a transformation is required which reduces the dimensionality of the tensor and to this end a number of tensor scalar measures have Address correspondence to: Dr Jonathan H Gillard, University Department of Radiology, Addenbrooke’s Hospital, Cambridge CB2 2QQ, UK.


Received 22 April 2003 Revised 24 May 2005 Accepted 1 June 2005 DOI: 10.1259/bjr/24908512 ’ 2006 The British Institute of Radiology

been proposed [1, 13]. From a theoretical point of view, tensor calculus establishes that many such measures exist. These include the lattice index (LI), relative anisotropy (RA), fractional anisotropy (FA), the volume ratio (VR) and ratios of the various eigenvalues (li), the mean diffusivity (D), the Euclidean length of the tensor (L), its anisotropy angle (w) and any algebraic combination of the first, second and third invariants of the tensor [14]. From a practical point of view, however, only a limited number of these measures are actually used in clinical studies. In the MR DTI literature, the most common of these measures are FA and D. Out of 30 recent studies on clinical applications of DTI, encompassing diseases such as schizophrenia, Alzheimer’s disease, stroke, multiple sclerosis and head injury, 26 reported their results using both FA and D [7, 9, 12, 15–36], while only four reported D alone [37–40]. The caveat with exclusively using D and FA to characterize pathology in clinical applications is that it is not known a priori which tensor scalar measure is the most appropriate to quantify pathological changes in brain tissue. It is conceivable, for example, that a study might fail to show significant changes when the diffusion tensor is measured using FA but it may show differences when using RA or L or some other measure. We have previously 101

A Penã, H A L Green, T A Carpenter et al

shown this to be the case in acute stroke [34]. The identification of which is the ‘‘best’’ measure of the diffusion tensor is an empirical process, which will only be resolved after a large number of experiments are conducted and corroborated with external empirical information, such as histology. In these circumstances it seems reasonable to analyse as many scalar measures as possible, and not rely solely on D and FA. This article explores the novel application of a mathematical technique to enhance the visualization and quantification of brain tissue in MR DTI, which we will term ‘‘p:q decomposition’’. The technique is based on a tensor transformation, which decomposes the diffusion tensor into its isotropic (p) and anisotropic (q) components. In contrast to the standard practice in the literature where only D and FA are analysed, this technique permits the visualization simultaneously of seven scalar measures. These are D, p, q, RA, FA, w, and L. This technique is based on a classical decomposition used in tensor calculus, already observed by the major contributions of Basser et al [13] and Pierpaoli et al [5], but which has not been applied yet to visualize and quantify the diffusion tensor in MR DTI. In the following sections we will describe the technique and apply it to data from a control volunteer and three clinical examples.

Materials and methods Theory Diffusion in tissue can be mathematically represented as a second-order Cartesian tensor, which in matrix form is: 2 3 Dxx Dxy Dxz 6 7 ð1Þ Dij ~4 Dyx Dyy Dyz 5 Dzx Dzy Dzz Given that the tensor is symmetric along its principal diagonal, i.e. Dyx5Dxy, it has only six independent components. From the tensor, the eigenvalues li are calculated using a standard methodology, such as singular value decomposition [41], as l1, l2 and l3. According to tensor calculus, based on the eigenvalues many possible scalar measures of the diffusion tensor can be devised [42, 43]. The standard methodology in the DTI literature, however, is to calculate only two scalar measures of Dij. These are the mean diffusion (D) defined as: 1 l1 zl2 zl3 ð2Þ D~ tr(Dij )~ 3 3 where tr represents the trace of the tensor, and the fractional anisotropy (FA) or the relative anisotropy (RA) defined as: qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi rffiffiffi 2 2 2 3 (l1 {D) z(l2 {D) z(l3 {D) qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi FA~ 2 l21 zl22 zl23 ð3Þ qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi rffiffiffi 2 2 2 1 (l1 {D) z(l2 {D) z(l3 {D) RA~ D 3 102

The proposed technique is based on a classical tensor decomposition, already observed by Basser et al [13] and Pierpaoli et al [5]. Our contribution is to construct a graphical representation of the diffusion tensor based on this decomposition. This transformation has its conceptual roots in the mathematical theory of continuum mechanics [44, 45]. We will term the technique p:q decomposition. Using this methodology, the first step is to decompose the diffusion tensor from Equation (1) according to the next equation: Dij ~DIij z½Dij DIij

ð4Þ

into two tensors P and Q, i.e. Dij5Pij+Qij. Here Iij is the identity tensor Iij5diag(1,1,1). The first term on the right hand side of Equation (4) is the isotropic tensor, while the second term (in brackets) represents the deviatoric tensor. The magnitude of these tensors can be denoted by its isotropic (p) and anisotropic (q) components. The values of p and q can be computed as: pffiffiffi p~ 3D ð5Þ and

qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi q~ (l1 {D)2 z(l2 {D)2 z(l3 {D)2

ð6Þ

According to these definitions p is therefore a scaled measure of the mean diffusion in the tensor, while q is a measure of the variance or deviation of the eigenvalues with respect to the mean diffusion of the tensor. The second step is to plot each tensor as a point in a Cartesian plane with p taken as the x-axis and q as the yaxis, as in Figure 1a. This plane will be denoted as the p:q plane. The effect of this transformation is to reduce the dimensionality of the tensor from six dimensions to two. The third step is to use the p:q plane to deduce the five additional tensor scalar measures: D, RA, FA, L and w. Four of these seven tensor measures (q, RA, FA, w) are anisotropy measures, while D and p are measures of the magnitude of diffusion and L is a measure of the total diffusion of the tensor. D, p, q and L have units of 1023 mm2 s21, RA and FA are dimensionless, and w has units of degrees. These scalar measures can be deduced either analytically or graphically. The analytical method is to directly compute the measures based on the p,q components using the formulae: rffiffiffi pffiffiffiffiffiffiffiffiffiffiffiffiffiffi q 3q {1 q 2 2 L~ p zq , RA~ , FA~ , w~ tan ð7Þ p 2L p However, the real advantage of using the p:q plane is that we can obtain these values directly from the graph as follows. Consider a tensor A from which we can compute its location in the p:q plane using Equations (5) and (6) as po and qo. Therefore, Lo is the distance between the origin of coordinates and the point (po,qo); wo is the angle subtended between the p axis and a line originating in the centre of coordinates and passing through point (po,qo), i.e. the segment Lo; RAo is the ratio between qo and FAo is the ratio between qo and Lowith scale factor po;ffiffiffiffiffiffi q 3= &1:22; D is p1ffiffi &0:57 of the value in the p axis. 2 3 The British Journal of Radiology, February 2006

Enhanced visualization and quantification in MR DTI Figure 1. (a) A point representing a sample of tissue in the p:q plane. The x axis corresponds to the isotropic component of diffusion (p) and the y axis the anisotropic component of diffusion (q). Any tensor can be decomposed into its p and q components po and qo, which will correspond to a point in the p:q plane. (b) Starting from a point in the p:q plane, we can deduce the standard anisotropy measures RA and FA using simple geometry. Both of these measures will be proportional to the angle w, in fact RA is proportional to the tangent and FA proportional to the sine. (c) Two tissues will, in general, have different p and q components. Thus a tissue A with components pA and qA, will have a different location from a tissue B with components pB and qB. (d) A tissue A in general will have different p and q components at different times. By plotting these different components in the p:q plane we can obtain a trajectory that illustrates the evolution of tissue in time. In this example we see a trajectory demonstrating three time points for tissue A.

Figure 1b illustrates the geometrical relationship between these various quantities.

Data acquisition In order to illustrate the use of the p:q decomposition with clinical data, four representative cases were selected. The first comprised five regions of interest (ROIs) in a healthy 27-year-old volunteer (to illustrate spatial variation in the tensor field, as in Figure 1c). The second reports the findings in a 76-year-old hypertensive woman who presented with a sudden onset of expressive dysphasia and right-sided hemiparesis. Imaging of her left middle cerebral artery territory stroke was undertaken at 37 h, 1 week and 3 months from stroke onset to illustrate the temporal variation in the tensor field (Figure 1d). The third case is a 85-year-old female hydrocephalus patient with a history of gait ataxia, falls and memory problems. And finally, the fourth case is a 46-year-old male patient with a WHO Grade II oligodendroglioma. The Local Research Ethics Committee approved the study and informed consent was obtained. The diffusion tensor data sets were acquired using a 5 mm slice thickness. Imaging was performed on a 3 Tesla magnetic resonance machine (Bruker Medspec S300; Bruker Medical, Ettlingen, Germany). A single shot spin echo, echo planar imaging technique, with Stejskal-Tanner diffusion sensitizing pulses [46] was used. Imaging parameters were: repetition time (TR)55070 ms, echo time (TE)5107 ms, a590˚, d521 ms and D566 ms. Eight interleaved supratentorial slices were acquired with a phase template in a near axial plane, using a 128 6 128 matrix, field of view of 25 cm 6 25 cm. For each slice, images were collected from 12 non-collinear gradient The British Journal of Radiology, February 2006

directions [47]. For each gradient direction an unweighted bo image and five diffusion weighted images were collected at equally spaced b-values in the range bmin5318 s mm22 to bmax51541 s mm22. Using a specially-written program in MATLAB (The MathWorks Inc., Natick, MA) the diffusion tensor was computed on a voxel by voxel basis, using a singular value decomposition algorithm to fit the signal intensities to the Stejskal-Tanner equation, following the method proposed by Basser et al [1, 2]. From the tensor, the p and q components were calculated based on Equations (5) and (6), and D, RA, FA, w, and L using Equations (2) and (7).

Results Normal volunteer D and FA maps of the volunteer were used to select square anatomical ROIs of 5 6 5 voxels, which were subsequently averaged to obtain a mean value for the ROI. These were placed in the corpus callosum (CC), occipital cortex (Cx), cerebrospinal fluid (CSF), internal capsule (IC) and noise regions, as illustrated in Figure 2a,b. When these ROIs were plotted in the p:q plane, they formed clearly segregated clusters (Figure 2c). The spherical diffusion and deviatoric diffusion fell within a narrow band with 0.976,p,1.063 6 1023 mm2 s21 for the three structures in brain parenchyma, i.e. CC, IC, Cx, D50.607, 0.563, 0.613 6 1023 mm2 s21 and p51.052, 0.976, 1.063 6 1023 mm2 s21, respectively. In contrast, all the other measures varied substantially, with a range bounded by maximum values for the CC (q51.086, L51.032) 6 1023 mm2 s21 and RA50.879, 45.9 ˚, FA51.512, and 103


(a)

(b)

(c) Figure 2. (a) Map of the mean diffusion (D) for a horizontal slice of the normal volunteer investigated, demonstrating the regions of interest used in the study, which are (from top to bottom) noise (N), cerebrospinal fluid (CSF), internal capsule (IC), splenium corpus callosum (CC) and occipital cortex (Cx). The scale on the right indicates the magnitude of D. (b) Map of the fractional anisotropy (FA) for the same horizontal slice of the normal volunteer investigated, demonstrating the location of the same regions of interest. The scale on the right indicates the dimensionless magnitude of FA. (c) p:q plane illustrating the defined regions of interest (ROIs) in the normal volunteer. Three clusters are observed for the noise (N), with small components for both p and q. The three parenchyma ROIs (CC, IC, Cx) are located along a line with approximately the same value of p, but significantly different values of q. The CSF has a much larger dispersion and a larger value of mean diffusion.

minimum values for the Cx (q50.150, L51.073) 6 1023 mm2 s21 and RA50.141, 8.06 ˚, FA50.171. Noise appeared as a cluster close to the origin of the coordinates (D50.007, p50.011, q50.141) 6 1023 mm2 s21, but whose additional scalar measures were amongst the highest (L51.073 6 1023 mm2 s21, RA512.145, 85.29˚, FA51.220. CSF presented the opposite characteristics, being the most distant to the origin of coordinates (D51.917, p53.320, q50.330) 6 1023 mm2 s21, and having small additional scalar measures (L51.073 6 1023 mm2 s21, RA50.099, 5.68 ˚, FA50.121). Statistically significant differences between the various ROIs were investigated using unpaired Student’s t-tests. The isotropic diffusion (p) between the noise and the parenchyma (CC, IC, Cx) was significantly different (pvalue,0.01), and between the CSF and the parenchyma 104

(p-value,0.01). It was significantly different between the CC and the IC (p-value,0.01), between the IC and the Cx (p-value,0.05), but not between the CC and the Cx (pvalue50.5473). The deviatoric diffusion (q) was different for the three parenchyma ROIs (CC, IC, Cx). It was significantly different between the CC and IC (pvalue,0.01), between the CC and Cx (p-value,0.01), and between the IC and Cx (p-value,0.01). Results for the seven scalar measures are presented in Table 1.

Stroke patient Lesion and contralateral control square anatomical ROIs of 5 6 5 voxels, were selected in the stroke patient using the FA and D maps at 37 h, 1 week and 1 month The British Journal of Radiology, February 2006

Enhanced visualization and quantification in MR DTI Table 1. Data from the MR diffusion tensor imaging (DTI) acquisitions for the normal volunteer (upper set) and the stroke patient (lower set) Normal volunteer

Corpus callosum Internal capsule Cortex Noise Cerebrospinal fluid

D

p

q

RA

FA

w

L

0.607 0.563 0.613 0.007 1.917

1.052 0.976 1.063 0.011 3.320

1.086 0.684 0.150 0.141 0.330

1.032 0.701 0.141 12.145 0.099

0.879 0.703 0.171 1.220 0.121

45.92 35.05 8.06 85.29 5.68

1.512 1.192 1.073 0.142 3.336

D

p

q

RA

FA

w

L

0.476 0.666 0.511 0.709 1.515 0.743

0.824 1.154 0.884 1.229 2.624 1.288

0.420 0.475 0.254 0.777 0.325 0.586

0.556 0.466 0.338 0.654 0.150 0.507

27.03 22.39 16.02 32.29 7.07 24.48

0.925 1.248 0.920 1.454 2.644 1.415

Stroke patient

37 h (lesion) 37 h (control) 1 week (lesion) 1 week (control) 1 month (lesion) 1 month (control)

0.510 0.412 0.287 0.632 0.124 0.455

Mean diffusion (D), isotropic component of diffusion (p), anisotropic component of diffusion (q), anisotropy angle (w), fractional anisotropy (FA), relative anisotropy (RA), Euclidean length (L).

(Figure 3a). To clarify presentation and due to the multiple number of ROIs used, all variables analysed were averaged within each ROI to obtain a mean value and standard deviation. For each ROI the mean is in the crossing of the bars which represent the magnitude of standard deviation. The lesion (ischaemic region) in the stroke patient described a trajectory in the p:q plane composed of three phases, which occupied the regions: 0.824,p,2.624 6 1023 mm2 s21 and 0.25,q,0.42 6 1023 mm2 s21. The (p,q) components of the lesion ROI were (0.824, 0.420), (0.884, 0.254), (2.624, 0.325) at 37 h, 1 week and 1 month, respectively. All these results are shown in Figure 3b. The corresponding contralateral control ROIs (Figure 3a, shown in blue), in contrast, demonstrated only a small degree of change, remaining in the region 1.154,p,1.288 6 1023 mm2 s21 and 0.47,q,0.77 6 1023 mm2 s21. The (p,q) components of the control ROI were (1.154, 0.475), (1.229, 0.777), (1.288, 0.586) at 37 h, 1 week and 1 month, respectively. Results for the seven scalar measures are presented in Table 1.

Hydrocephalus patient We have investigated microstructural changes in the internal capsule associated with the ventricular dilatation in this patient. For this analysis, four ROIs have been selected (two in the patient bilaterally and two in the control volunteer bilaterally) in one axial slice corresponding to the posterior limb of the internal capsule (IC) at the level of the foramen of Monro. Each ROI was composed of nine voxels. As a comparison, the same regions were selected in a control volunteer in the same manner. The results were: for the patient p51.29¡0.466, q50.85¡0.054 for the left IC and p50.96¡0.195, q50.86¡0.147 for the right IC; and for the control, p51.01¡0.023, q50.80¡0.084 for the left IC, and p50.96¡0.034, q50.84¡0.159 for the right IC. All the p, q units are in 1023 mm2 and are illustrated in Figure 4. These results demonstrate that while the mean values of the four ROIs are roughly similar, there is a marked increase in the dispersion of the voxels in the IC of the patient. The British Journal of Radiology, February 2006

Tumour patient In this patient we investigated the hypothesis that there are significant differences between the microstructure in the tumour region and an equivalently-located contralateral control. For this analysis, two ROIs have been selected in the patient in one axial slice: one in the tumour region (the posterior pericallosal region) and another in the equivalent normal-appearing contralateral control region. For the tumour: p52.26¡0.210, q5 0.37¡0.049, while for the contralateral control region p 51.331¡0.128, q50.71¡0.134. All the p, q units are in 1023 mm2 and are illustrated in Figure 5. These results demonstrate both an increase in the isotropic component of diffusion and a decrease in the deviatoric component. We can interpret these changes as suggesting a loss in the microstructure of tissue, as a decrease in the density of fibres (more intercellular space) and a loss in fibre coherence. These changes are consistent with previous reports [48].

Discussion We have presented a technique that permits the simultaneous visualization of multiple tensor scalar measures from MR DTI data. In particular, we have shown how, from a single graph (the p:q plane), it is possible to deduce seven scalar measures of the diffusion tensor, including D, p, q, RA, FA, w, and L. This represents an improvement on the standard methodology in the MR DTI literature in which only two scalar measures (typically FA and D) are displayed. There have already been a number of studies in the literature that have considered plotting simultaneously two tensor scalar measures, particularly FA and D. These include Pierpaoli et al [5] who distinguished various brain regions based on decomposing the tensor in terms of D and the volume ratio (VR). Werring et al [8] in an investigation of normal-appearing white matter lesions in multiple sclerosis, and Wieshmann et al [49] and Jones et al [50] have also demonstrated the potential of plotting FA and D simultaneously. Plotting D vs FA, however, 105


(a)

(b)

Figure 3. (a) Mean diffusion (D) and fractional anisotropy maps (FA) for a stroke patient at three time points: 37 h (left column), 1 week (central column) and 3 months (right column). FA is lower row and D is upper row. These maps demonstrate the regions of interest (ROIs) used in this study. The lesion ROIs are presented in orange and the control ROIs in green. Each of these ROIs consisted of 5 6 5 6 1 voxels. (b) This figure illustrates the p:q plane for the stroke patient, with lesion and control ROIs at (a) 37 h, (b) at 1 week and (c) at 3 months. The control ROIs are denoted in blue and the lesion ROIs in red. The arrows demonstrate the trajectory followed by the lesion in this patient and show schematically how, while the control ROIs remain in roughly the same region in the p:q plane, the ischaemic lesion demonstrates a trajectory composed of acute (reduction in p, reduction in q), subacute (normalization of p while q remains low) and chronic (increased p while q remains low) phases. The inset shows schematically the location of the lesion ROI with respect to the control ROI and a line of constant fractional anisotropy. As FA is function of the angle w, the figure indicates that at 37 h the lesion has a higher FA than the control, while at 1 week it has a lower FA than the control.

does not allow visualization or quantitative analysis of the other scalar measures (e.g. p, q, RA, w, L) from a single graph, while our study does. To illustrate the use of our method for clinical data, we have applied it to a healthy volunteer, a sequential study in a patient with recent stroke, a patient with hydrocephalus and a patient with an intracranial tumour. In all cases the p:q plane offers the analyst a concise and easy-to-use representation of the diffusion tensor. The first case illustrates the spatial variation in the tensor

field and statistically significant differences between different tissue types (e.g. grey matter, white matter), while the second case illustrates the temporal variation in the tensor field and thus the evolution of the lesion (e.g. lesion, contralateral control). We propose that the p:q decomposition is a powerful aid not only in the visualization of the data, but also in its analysis, by offering a unique opportunity to assess the additional non-standard tensor scalar measures (p, q, w, L) and their relationship with the standard measures (D,

Figure 4. This is the p:q diagram for the internal capsule (IC) of a hydrocephalus patient. Regions of interest (ROIs) have been selected on the IC bilaterally at the level of the foramen of Monro. The same ROIs have been selected in a control subject. ROI location is shown in the insets (patient, above; control, below). The p:q diagram demonstrates an increased dispersion (disorganization) of the white matter tracts of the IC in the hydrocephalus patient as compared with the control. 106


Enhanced visualization and quantification in MR DTI

Figure 5. This is the p:q diagram of a patient with a Grade II oligodendroglioma in the posterior pericallosal white matter. The location of the regions of interest (ROIs) is indicated in the inset (above), and the patient’s MR fluid attenuation inversion recovery (FLAIR) image (below). Compared with the control region, the tumour region demonstrates both an increase in isotropic diffusion (p) and a decrease in deviatoric diffusion (q). This tissue signature is consistent with the destruction of white matter tracts in the tumour region.

RA, FA). In this context, several interesting observations from both the normal volunteer and the stroke patient have been possible by using the p:q technique.

Normal volunteer The first observation is that the p:q plane provides a graphical means to understand the complex equations that describe RA and FA. From Figure 1b and Equation (7), one can easily observe that both RA and FA are composite measures of other more basic tensor quantities. In particular, RA is simply the ratio between q and ffiffiffiffiffiffi FA the ratio between q and L scaled by a factor p, q and of 3=2&1:22. The second observation is that the p:q plane explains some anomalies when using FA and RA as measures of anisotropy. For example, if we take the value of FA for noise from Table 1, we obtain the theoretical maximum FA value of 1.22. This is confusing, as one would not expect empty space to have a large degree of organization (anisotropy). This paradoxical result is in fact a methodological artefact in using FA as an anisotropy measure, and is due to the presence of L in the denominator of FA. Given that the diffusion of empty space should be zero (or close to zero due to experimental error), a very small L will imply a very large FA. Thus FA is a measure of tissue anisotropy, but weighted by its total diffusion. The same argument applied to the value of RA, which gives the enormous value of 12 (while the corpus callosum, for instance, is 1.032). The third observation is the insight that q might offer as a measure for the background noise in the data. Equation (6) can be interpreted in statistical terms such that q is a measure of the variance of the eigenvalues of the tensor with respect to the mean diffusion D. Therefore, isotropic elements in the data (such as the empty space, Cx and the CSF) should theoretically have all the eigenvalues equal and thus a q equal to zero. However, due to experimental error, background noise and other MR acquisition influences, there is a small The British Journal of Radiology, February 2006

discrepancy and the eigenvalues are not exactly the same. Our results show that Cx and the noise have approximately the same q values (0.141 and 0.15), while the CSF presented a larger dispersion (q50.330), which might be attributed to the contribution of diffusion and bulk flow during the acquisition time.

Stroke patient The first observation is the ability of the p:q planes to visually convey simultaneous changes in the isotropic and anisotropic components of the diffusion tensor as they change in time, in other words the ‘‘trajectory’’ of the tensor. In addition to the qualitative nature of the trajectory, the magnitude of tensor changes can be read directly from the p and q axes of the plots. In our example, the trajectory describing the lesion evolution is composed of three segments or phases (Figure 3b), which can be interpreted in biological terms as the acute (reduction in p, reduction in q), sub-acute (pseudonormalization of p, while q remains reduced) and chronic (increase in p, while q remains reduced) phases that have been well-documented in association with stroke [9]. A second observation demonstrates another methodological artefact or anomaly of FA and RA. Close inspection of Figure 3b demonstrates that the lesion clusters (shown in red) with respect to the control clusters (shown in blue) are displaced first above a line of constant FA in the acute phase (37 h) and subsequently below this line in the sub-acute phase (1 week), as shown in the inset. As both RA and FA are functions of the angle w, this would imply that they are increased in the lesion as compared with the control, which is absurd. This paradox of increased tissue anisotropy (as measured in terms of RA or FA) was reported by Nusbaum and colleagues [51] in normal ageing. As we have explored in more detail in the case of acute stroke [34], the p:q technique provides a graphical explanation of why this can be the case, and that this apparent increase in anisotropy (as measured in terms of RA or FA) can be 107


purely a graphical consequence of the manner in which FA and RA are calculated and thus a methodological artefact. Quantitatively, the anisotropy measured with FA and RA of the lesion’s acute phase (37 h): RA increased from 0.412 to 0.510 (or +24%) and FA increased from 0.466 to 0.556 (or +19%). In contrast, q decreased from 0.475 to 0.420 (or 212%). This behaviour suggests, albeit tentatively, that theoretically q may detect early changes in tissue anisotropy that are misrepresented by RA and FA. A third observation is that during the sub-acute phase (1 week after the stroke), the best sensitivity to the pathology is offered by q rather than by FA or RA. q was reduced from 0.777 to 0.254 (267%), In contrast, FA and RA decreased by the smaller amounts of 0.632 to 0.287 (255%), and 0.654 to 0.338 (248%), respectively. Based on the previous observations, we can conclude that, at least in some circumstances, some non-standard anisotropy measures (such as q) can provide a higher sensitivity to detect pathological conditions than standard measures such as RA and FA. We have also shown that RA and FA have the potential to give ‘‘paradoxical’’ results and thus must be used with caution. However, this analysis does not resolve what is perhaps the most important question in MR DTI: from all the various tensor measures, which one is the best one to characterize damage to brain tissue? As has been recently noted by Pierpaoli et al [23], the fact remains that we do not know a priori which is the best measure because this is not a theoretical question but an empirical one. It is equivalent to asking which statistical measure, e.g. the mean or the variance for example, will better describe a population. They describe different aspects of a population and therefore will be useful in answering different questions. Tensor calculus can only help by defining which measures can be used in our analysis. Which one best describes some aspect of the brain (be it a tissue type or a pathological condition, such as oedema or necrosis) can only be answered empirically, by relating the observed tensor measures with independent biological data such as histology, other imaging modalities and/or cognitive tests.

Hydrocephalus patient The disruption observed in IC using the p:q diagram from the hydrocephalus patient is encouraging. In patients with hydrocephalus it is common to observe clinical symptoms that are thought to be associated with the disruption of deep white matter tracts. Similar findings were observed in this patient at the level of the internal capsule (IC). It has been suggested that during ventricular dilatation, these tracts are being stretched and thus become mechanically compromised. Our results support this notion by demonstrating that the MR DTI diffusion signature of the IC is altered. In particular, this disruption is not due to changes in the anisotropy of tissue but to changes in its mean diffusion. This suggests that the white matter tracts have been disrupted.

Tumour patient In the case of the tumour patient, the p:q decomposition was useful to illustrate simultaneously changes in 108

both the isotropic and the anisotropic components of the diffusion tensor. There was a decrease in anisotropy (q) and an increase in mean diffusion (p). These changes are thought to be associated with white matter destruction. There are a number of limitations in using the p:q decomposition. Just like D, RA and FA, the location of tissue in the p:q plane does not give information about the directionality of diffusion. Also, there are other important tensor scalar measures that are not directly conveyed by the p:q plane, such as the eigenvalues and the tensor invariants. Finally, from a practical point of view, the p:q decomposition must be applied to other brain pathologies in order to establish how beneficial it might be in those situations.

Conclusion The p:q tensor decomposition enhances the visualization and quantification of MR DTI data in both normal and pathological conditions. In particular it is an aid to visualize simultaneously seven scalar tensor measures. We have also shown the pitfalls of using FA and RA exclusively, and the potential of using other tensor measures, particularly q. However, it is important to note that, despite the enhanced visualization and quantification provided by our technique, the choice of which tensor scalar measure best describes brain tissue and its changes remains an empirical matter. We hope that the enhanced repertoire of analysis tools that we propose might enable improved categorization of tensor abnormalities in pathology.

Acknowledgments AP is in receipt of a Wellcome Trust Fellowship in Mathematical Biology. The Cambridge Commonwealth Trust supports HALG. The Medical Research Council Technology Foresight grant and the Wolfson Foundation support the Wolfson Brain Imaging Centre. We acknowledge the help of radiographers Tim Donovan, Victoria Lupson and Ruth Bisbrown-Chippendale, the many useful discussions with Dr Neil G Harris, Dr Brian K Owler, Dr Luzius A Steiner and Dr Shahan Momjian, as well as the excellent computing support of Mr Julian Evans.

References 1. Basser PJ. Inferring microstructural features and the physiological state of tissues from diffusion-weighted images. NMR Biomed 1995;8:333–44. 2. Basser PJ, Mattiello J, LeBihan D. Estimation of the effective self-diffusion tensor from the NMR spin echo. J Magn Reson B 1994;103:247–54. 3. Basser PJ, Mattiello J, LeBihan D. MR diffusion tensor spectroscopy and imaging. Biophys J 1994;66:259–67. 4. Pierpaoli C, Basser PJ. Toward a quantitative assessment of diffusion anisotropy. Magn Reson Med 1996;36:893–906. 5. Pierpaoli C, et al. Diffusion tensor MR imaging of the human brain. Radiology 1996;201:637–48. 6. Foong, J, et al. Neuropathological abnormalities of the corpus callosum in schizophrenia: a diffusion tensor imaging study. J Neurol Neurosurg Psychiatry 2000;68:242–4. 7. Arfanakis K, et al. Diffusion tensor MR imaging in diffuse axonal injury. AJNR Am J Neuroradiol 2002;23:794–802.


Enhanced visualization and quantification in MR DTI 8. Werring DJ, et al. Diffusion tensor imaging of lesions and normal-appearing white matter in multiple sclerosis. Neurology 1999;52:1626–32. 9. Sorensen AG, et al. Human acute cerebral ischemia: detection of changes in water diffusion anisotropy by using MR imaging. Radiology 1999;212:785–92. 10. Gillard JH, et al. MR diffusion tensor imaging of white matter tract disruption in stroke at 3 T. Br J Radiol 2001;74:642–7. 11. Mori S, et al. Imaging cortical association tracts in the human brain using diffusion-tensor-based axonal tracking. Magn Reson Med 2002;47:215–23. 12. Wieshmann UC, et al. Diffusion tensor imaging demonstrates deviation of fibres in normal appearing white matter adjacent to a brain tumour. J Neurol Neurosurg Psychiatry 2000;68:501–3. 13. Basser PJ, Pierpaoli C. Microstructural and physiological features of tissues elucidated by quantitative-diffusiontensor MRI. J Magn Reson B 1996;111:209–19. 14. Skare S, et al. Noise considerations in the determination of diffusion tensor anisotropy. Magn Reson Imaging 2000;18: 659–69. 15. Foong J, et al. Neuropathological abnormalities of the corpus callosum in schizophrenia: a diffusion tensor imaging study. J Neurol Neurosurg Psychiatry 2000;68: 242–4. 16. Werring DJ, et al. Diffusion tensor imaging of lesions and normal-appearing white matter in multiple sclerosis. Neurology 1999;52:1626–32. 17. Wieshmann UC, et al. Reduced anisotropy of water diffusion in structural cerebral abnormalities demonstrated with diffusion tensor imaging. Magn Reson Imaging 1999;17:1269–74. 18. Werring DJ, et al. Diffusion tensor imaging can detect and quantify corticospinal tract degeneration after stroke. J Neurol Neurosurg Psychiatry 2000;69:269–72. 19. Rugg-Gunn FJ, et al. Diffusion imaging shows abnormalities after blunt head trauma when conventional magnetic resonance imaging is normal. J Neurol Neurosurg Psychiatry 2001;70:530–3. 20. Tievsky AL, Ptak T, Farkas J. Investigation of apparent diffusion coefficient and diffusion tensor anisotrophy in acute and chronic multiple sclerosis lesions. AJNR Am J Neuroradiol 1999;20:1491–9. 21. Jones DK, et al. Characterization of white matter damage in ischemic leukoaraiosis with diffusion tensor MRI. Stroke 1999;30:393–7. 22. Rugg-Gunn FJ, et al. Diffusion tensor imaging of cryptogenic and acquired partial epilepsies. Brain 2001;124:627–36. 23. Pierpaoli C, et al. Water diffusion changes in Wallerian degeneration and their dependence on white matter architecture. Neuroimage 2001;13:1174–85. 24. Rovaris M, et al. Cognitive dysfunction in patients with mildly disabling relapsing-remitting multiple sclerosis: an exploratory study with diffusion tensor MR imaging. J Neurol Sci 2002;195:103–9. 25. Bozzali M, et al. Quantification of tissue damage in AD using diffusion tensor and magnetization transfer MRI. Neurology 2001;57:1135–7. 26. Sinha S, et al. Diffusion tensor MR imaging of high-grade cerebral gliomas. AJNR Am J Neuroradiol 2002;23:520–7. 27. Guo AC, Jewells VL, Provenzale JM. Analysis of normalappearing white matter in multiple sclerosis: comparison of diffusion tensor MR imaging and magnetization transfer imaging. AJNR Am J Neuroradiol 2001;22:1893–900. 28. Cercignani M, et al. Mean diffusivity and fractional anisotropy histograms of patients with multiple sclerosis. AJNR Am J Neuroradiol 2001;22:952–8.


29. O’Sullivan, M, et al. Evidence for cortical ‘‘disconnection’’ as a mechanism of age-related cognitive decline. Neurology 2001;57:632–8. 30. Ciccarelli O, et al. Investigation of MS normal-appearing brain using diffusion tensor MRI with clinical correlations. Neurology 2001;56:926–33. 31. Guo AC, MacFall JR, Provenzale JM. Multiple sclerosis: diffusion tensor MR imaging for evaluation of normalappearing white matter. Radiology 2002;222:729–36. 32. Mukherjee P, McKinstry RC. Reversible posterior leukoencephalopathy syndrome: evaluation with diffusion-tensor MR imaging. Radiology 2001;219:756–65. 33. Lim KO, et al. Reduced frontal white matter integrity in cocaine dependence: a controlled diffusion tensor imaging study. Biol Psychiatry 2002;51:890–5. 34. Green HA, et al. Increased anisotropy in acute stroke: a possible explanation. Stroke 2002;33:1517–21. 35. Rovaris M, et al. Assessment of normal-appearing white and gray matter in patients with primary progressive multiple sclerosis: a diffusion-tensor magnetic resonance imaging study. Arch Neurol 2002;59:1406–12. 36. Rocca, MA, et al. A diffusion tensor magnetic resonance imaging study of brain tissue from patients with migraine. J Neurol Neurosurg Psychiatry 2003;74:501–3. 37. Molko N, et al. Monitoring disease progression in CADASIL with diffusion magnetic resonance imaging: a study with whole brain histogram analysis. Stroke 2002;33:2902–8. 38. Inglese M, et al. Magnetization transfer and diffusion tensor MR imaging of acute disseminated encephalomyelitis. AJNR Am J Neuroradiol 2002;23:267–72. 39. Bozzali M, et al. Quantification of brain gray matter damage in different MS phenotypes by use of diffusion tensor MR imaging. AJNR Am J Neuroradiol 2002;23:985–8. 40. Codella M, et al. A preliminary study of magnetization transfer and diffusion tensor MRI of multiple sclerosis patients with fatigue. J Neurol 2002;249:535–7. 41. Golub GH, Van Loan CF. Matrix computations 3rd edn. The Johns Hopkins University Press, Baltimore, MD, 1999. 42. Alexander AL, et al. A geometric analysis of diffusion tensor measurements of the human brain. Magn Reson Med 2000;44:283–91. 43. Bahn MM. Invariant and orthonormal scalar measures derived from magnetic resonance diffusion tensor imaging. J Magn Reson 1999;141:68–77. 44. Mase GE. Continuum Mechanics. Schaum Outline Series. McGraw-Hill Publishing Company, 1970. 45. Fung YC. A first course in continuum mechanics. Englewood Cliffs, NJ: Prentice Hall, 1994. 46. Stejskal EO, Tanner JE. Spin diffusion measurements: spin echoes in the presence of a time dependent field gradient. J Chem Phys 1965;42:288–92. 47. Papadakis NG, et al. A comparative study of acquisition schemes for diffusion tensor imaging using MRI. J Magn Reson 1999;137:67–82. 48. Price SJ, et al. Diffusion tensor imaging of brain tumours at 3T: a potential tool for assessing white matter tract invasion? Clin Radiol 2003;58:455–62. 49. Wieshmann UC, et al. Reduced anisotropy of water diffusion in structural cerebral abnormalities demonstrated with diffusion tensor imaging. Magn Reson Imaging 1999;17:1269–74. 50. Jones DK, et al. Characterization of white matter damage in ischemic leukoaraiosis with diffusion tensor MRI. Stroke 1999;30:393–7. 51. Nusbaum AO, et al. Regional and global changes in cerebral diffusion with normal aging. AJNR Am J Neuroradiol 2001;22:136–42.

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PET-CT findings in surgically transposed ovaries 1,2,3 1,3

R ZISSIN, MD, 1U METSER, E EVEN-SAPIR, MD, PhD

MD,

1

H LERMAN,

MD,

1

G LIEVSHITZ,

MD,

4

T SAFRA,

MD

and

Department of 1Nuclear Medicine and 4Oncologic Surgery Unit, Tel-Aviv Sourasky Medical Center and the Department of 2Diagnostic Imaging, Sapir Medical Center, Kfar Saba, affiliated to the 3 Sackler Faculty of Medicine, Tel-Aviv, Israel ABSTRACT. The aim of this study is to present the PET/CT findings of surgically transposed ovaries. PET/CT studies and associated abdominal imaging studies of seven women, aged 28–43 years, with 11 transposed ovaries were retrospectively reviewed. Attention was directed to the location and the 18F-Fluorodeoxyglucose (FDG) avidity of the transposed ovaries. On the CT part of the PET/CT, location of the transposed ovaries was in the ipsilateral iliac fossa or paracolic gutter abutting the anterior aspect of the ipsilateral colon (n56), posterolateral to the cecum (n54) and in the anterior abdominal cavity (n51). Ovaries were of soft-tissue density (n510 with a hypodense region in two) and one was cystic. In three patients, the transposed ovary was associated with increased FDG uptake with standard uptake values ranging from 2.4 to 4.8. Two of the latter patients had more than one PET/CT study. FDG uptake altered between studies, probably related to the performance of the study on different phases of the cycle. Menstrual history in one of the patients confirmed that the study was performed at the ovulatory-phase of the cycle. To conclude, a transposed ovary may appear on a PET-CT study as a mass with occasionally increased FDG uptake that may be related to its preserved functionality. Physicians interpreting PET/CT should be aware of surgically transposed ovaries in young female patients to avoid misdiagnosing it as tumour. Pelvic radiation therapy for cervical, vaginal or colorectal cancer often leads to ovarian failure. Ovarian transposition outside the radiation field, to the paracolic gutter or iliac fossa, is a surgical procedure performed to preserve ovarian function mainly in young females with early stages of cervical carcinoma [1]. On imaging, the transposed ovary may appear as a small soft-tissue mass, often with one or more tiny cysts, or alternatively as a larger intraperitoneal cystic mass which may show functional, periodic changes on follow-up studies, according to the expected changes in the ovary during the different phases of the menstruation cycle. Surgical clips are usually placed to permit identification of the transposed ovary [2–5]. In oncologic patients, the recognition of the position and the appearance of the transposed ovary are crucial to avoid misinterpreting it as a tumour. We have encountered 10 18F-Fluorodeoxyglucose (FDG) PET/CT studies in 7 females with 11 surgically transposed ovaries and we present their imaging findings on PET/CT and conventional abdominal imaging.

Materials and methods We reviewed the clinical data and imaging studies of seven female oncologic patients (aged 28–43 years) after ovarian transpositions who were referred for PET/CT studies. Ovarian transposition was bilateral in four Address correspondence to: Einat Even-Sapir, Department of Nuclear Medicine, Tel-Aviv Sourasky Medical Center, 6 Weizman Street, Tel-Aviv, 64239 Israel.

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Received 17 February 2005 Revised 21 May 2005 Accepted 15 June 2005 DOI: 10.1259/bjr/33143536 ’ 2006 The British Institute of Radiology

patients and unilateral in the other three. Five women had carcinoma of the cervix, one had a rectovaginal cleft mucinous adenocarcinoma and one had uterine nonHodgkin’s lymphoma. Six patients reported amenorrhoea after hysterectomy and one was menstruating. Two of the study patients had more than one PET/CT study, at different time points in the menstruation cycle, available for assessment. One patient had two studies and the other had three. Five PET/CT studies were performed for findings suggestive of recurrence that were detected on physical examination and/or seen on MRI or diagnostic CT performed for follow-up. In two patients, five follow up PET-CT studies were performed, for re-staging and for monitoring response to treatment. PET-CT scan was performed following the administration of iodinated oral contrast material and after intravenous injection of 370–666 MBq (10–18 mCi) of 18 FDG. Low-dose CT scanning was performed (140 kV, 80 mA, 0.8 s per CT rotation, pitch of 6, and table speed of 22.5 mm s21) during normal respiration. PET scanning was performed immediately following the CT without changing the patient position. Images were interpreted at a work-station (Xeleris Elgems, Haifa, Israel) equipped with fusion software that enables the display of PET, CT and fused PET/CT images.

Results The clinical and imaging findings of the patients are summarized in Table 1. All 11 transposed ovaries were The British Journal of Radiology, February 2006

PET-CT findings in surgically transposed ovaries Table 1. The clinical data and imaging findings of 7 patients with transposed ovaries Patient no., age (years), primary tumour

Medical history

Indication for PET/CT

PET/CT findings

1. 35, carcinoma of cervix

6 months after Lt. SO, Rt. OT and 3 months after combined chemo-radiotherapy

A Rt. gutter ST mass on CT – suspicion of recurrence


18 months after radical hysterectomy, Rt. SO and Lt. OT

An intra-abdominal ST mass on CT – suspicion of recurrence


10 months after radical hysterectomy, pelvis lymphadenectomy and bilateral OT

Suspected mesenteric lymphadenopathy on CT


10 years after radical hysterectomy, pelvis lymphadenectomy and bilateral OT -S/P fluid aspiration from a Lt. ovarian cyst, 5 years earlier

A 5 cm cystic (necrotic?) RLQ mass on MRI – suspicion of recurrence


6 years after radical hysterectomy, pelvis lymphadenectomy and bilateral OT

Clinical suspicion of recurrence

6. 39, uterine non-Hodgkin’s lymphoma

5 years after radical hysterectomy, pelvis lymphadenectomy and bilateral OT

1st: Clinical suspicion of recurrence

A 2.3 cm63.2 cm ST mass, with central hypodensity, near surgical clips, in Rt. gutter, posterolateral to AC, cranially to a normal appendix. No FDG uptake A 1.7 cm63.7 cm ST mass, near surgical clips, in the anterior mid-abdomen, between bowel loops and Lt. rectus abdomini muscle. No FDG uptake Rt. A 2.1 cm63.3 cm ST mass, near surgical clips, in the Rt. iliac fossa, posterolateral to the cecum. No FDG uptake Lt. A 2.4 cm62 cm ST mass, near surgical clips, in the Lt. iliac fossa, anterior to DC. No FDG uptake Rt. A 2.7 cm61.5 cm ST mass, near surgical clips, in the Rt. iliac fossa, anterior to the cecum. No FDG uptake Lt. A 3.8 cm63 cm hypodense mass, near surgical clips, in the Lt. iliac fossa, anterior to DC. Mild FDG uptake (SUV-2.4) Rt. A 1.9 cm61 cm ST mass, near surgical clips, in the Rt. gutter posterolateral to AC. No FDG uptake Lt. A 2.6 cm60.6 cm ST mass, near surgical clips, in the Lt. gutter posterolateral to DC. No FDG uptake 1st: Rt. A 2.9 cm61.7 cm ST mass near surgical clips, in the Rt. gutter, posterolateral to the cecum with FDG uptake (SUV-4.8) Lt. A 2.1 cm62.4 cm ST mass with hypodense centre, near surgical clips, in the Lt. gutter, lateral to DC. No FDG uptake 2nd. Rt. A 2.9 cm62.7 cm ST mass. No FDG uptake Lt. Same as in the previous study 3rd: no change from previous study 1st: A 2.3 cm61.8 cm hypodense mass, near surgical clips, in the Rt. gutter anterior to AC. No FDG uptake – Pararectal local recurrence 2nd: A 3 cm61.9 cm ST mass with FDG uptake (SUV-3.6)– Progression of local pelvic disease

2nd: follow-up 6 months later

7. 28, rectovaginal cleft mucinous adenocarcinoma, S/P breast cancer

4 months after limited surgical excision of the tumour, Rt OT + chemo-radiotherapy

3rd: follow-up 6 months later 1st: Suspected local recurrence in the Rt. pararectal space on MRI

2nd: 3 months later (on mid-cycle) – to monitor response to therapy

SO, salpingo-oophorectomy; Lt., left; Rt., right; LLQ, left lower quadrant; RLQ, right lower quadrant; ST, soft tissue; AC, ascending colon; DC, descending colon; OT, ovarian transposition.


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recognized on the CT part of the PET/CT study, adjacent to surgical clips. Their location was in the ipsilateral iliac fossa or paracolic gutter (n510), either abutting the anterior or lateral aspect of the ipsilateral colon (n56) (Figure 1) or posterolateral to the cecum (n54) (Figure 2), and in the anterior abdominal cavity between small bowel loops and the left rectus abdomini muscle at the level of L3 vertebra (n51). Ten ovaries were of softtissue density, with a hypodense region in two of them, while the remaining one showed periodic CT changes, related to the menstruation cycle, which varied from a ‘‘cystic’’ to a soft-tissue attenuating mass. In three patients, the transposed ovary was associated with increased FDG uptake. One patient, with bilateral ovarian transposition, was referred for PET/CT for the assessment of a ‘‘necrotic’’ mass demonstrated on MRI (Figure 1a). On PET/CT, performed 1 month later, the lesion showed significant diminution in size without FDG uptake, while minimal uptake (standard uptake value of 2.4) was seen in the contralateral transposed

ovary (Figure 1b,c). As the patient was amenorrhoeic following hysterectomy, we could only assume that the MRI and PET/CT findings represented periodic changes in bilaterally transposed ovaries. The second patient, with bilateral ovarian transposition after hysterectomy, had three PET-CT studies. On the first study, the right transposed ovary presented as a soft-tissue mass with increased FDG uptake (standard uptake value of 4.8) (Figure 2a,b). On the second study, 6 months later, without any treatment in the interim, the same ovary presented as a soft-tissue mass with no uptake (Figure 2c). These findings remained unchanged on a third follow-up study. In the third patient, who was still menstruating as she had an intact uterus, a rim of FDG uptake (standard uptake value of 3.6) was detected in the transposed ovary on a study performed 14 days after menstruation. That ovary was demonstrated on the CT part of the study as a soft-tissue mass. Based on the menstrual history of the patient, it appeared that the patient was in the ovulatory-phase of the cycle.

(a)

(b) Figure 1. A 30-year-old woman, 10 years after radical hysterectomy and bilateral ovarian transposition for carcinoma of the cervix, referred for PET/CT for suspected recurrence on MRI (patient no. 4). (a) An axial T2 weighted MR image at the pelvic inlet shows the transposed right ovary (RO) anteriorly to the ascending colon (AC) as a 5 cm hyperintense mass with a thin hypointense rim, suspected to be a necrotic tumour recurrence. Note also the left transposed ovary (LO), abutting the anterior aspect of the descending colon (DC), as a hypointense lesion. That ovary was not reported on the MRI. (b) Axial PET/CT images (from left to right: CT, PET and fused PET/CT images). On the CT, the bilateral transposed ovaries are seen (thin white arrows). Note the diminution in size of the right ovary in comparison with the previous MRI performed 1 month earlier, most likely related to its periodic functional changes. The left transposed ovary shows increased FDG uptake on the PET and on fused images (thin arrows). Additional physiological sites of FDG uptake are seen, including bowel (arrowhead), bone marrow (medium-size arrow) and iliac blood vessels (large arrow). (Continued)

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PET-CT findings in surgically transposed ovaries

(c) Figure 1. (Cont.) (c) Coronal PET-CT images (from left to right: CT, PET and fused PET/CT images) show mild increased FDG uptake in the left transposed ovary (arrows). Physiological FDG uptake is seen in the brain, myocardium, bowel and liver.

This increased ovarian uptake was not detected on a previous PET/CT study, performed not at the ovulatory phase, associated with an altered appearance of the ovary, seen on the CT part of that study, as a hypodense mass.

Discussion Ovarian transposition was described by McCall et al in 1958 for young (,40 years old) females with an early-stage cervical carcinoma planned for pelvic radiosurgical treatment, to maintain ovarian function [6]. The procedure may be unilateral or bilateral, performed at the time of the radical hysterectomy or staging lymphadenectomy [1]. The repositioning of the ovary outside the radiation field may be above the iliac crest, into the ipsilateral paracolic gutter or lower down, below the iliac crest lateral to the iliopsoas muscle [2]. The normal transposed ovary may appear on abdominal CT as a soft-tissue mass, sometimes with small cysts or as a predominant cystic lesion, mimicking a peritoneal or retroperitoneal tumour implants. The location of the transposed ovary on CT is generally either adjacent to The British Journal of Radiology, February 2006

the ascending or descending colon, or in the upper pelvis lateral to or anterolateral to the psoas muscle [2–5]. However, in one of our patients, the transposed ovary was in an atypical location, i.e. in the anterior abdominal cavity between the abdominal wall musculature and the small bowel loops, mimicking a peritoneal implant. Adjacent surgical clips assisted in identifying it as a transposed ovary. Lack of familiarity with the procedure as well as with the CT features of a transposed ovary may lead to a diagnostic error in the interpretation of abdominal CT or MR imaging, misdiagnosing the transposed ovary as a metastatic deposit. It was the case in five of our patients, who were referred for a PET/CT study for a ‘‘suspected’’ tumoural recurrence on either CT or MRI. A right-sided transposed ovary should also be differentiated from a mucocele of the appendix, although an appendectomy is usually performed at the time of the surgical procedure [5]. In one of our patients, the appendix was not removed and was identified separately from the ovary on the CT part of the study, obviating such an interpretation mistake. Recently, hybrid systems composed of PET and CT have been introduced in addition to conventional 113

R Zissin, U Metser, H Lerman et al

(a)

(b)

(c)

114


PET-CT findings in surgically transposed ovaries Figure 2. A 39-year-old patient with a previous hysterectomy for uterine non-Hodgkin’s lymphoma, referred for suspected central nervous system recurrence (patient no. 6). (a) On PET data (from left to right: coronal, sagittal and transaxial images) increased FDG uptake was detected in the right abdomen (arrows). (b) On PET-CT (from left to right: CT, PET and fused PET/CT images) the increased uptake corresponded in location to a soft-tissue mass posterolateral to the cecum, adjacent to surgical clips, identified as a transposed ovary (arrows). (c) PET/CT study performed 6 months later, without treatment in the interim. The transposed ovary shows no increased uptake, confirming the functional aetiology of FDG uptake on the first study (arrow).

cross-sectional imaging methods in the routine practice of oncologic patients for staging, monitoring response to treatment and assessment of recurrence. PET and CT data, acquired at the same clinical setting, with generation of fused PET/CT images, provide both functional and anatomical information [7]. A transposed ovary may show increased FDG uptake on the PET part of the study due to functional changes, as was seen in three of our patients. FDG uptake in normal ovaries was reported in pre-menopausal patients without a known ovarian malignancy at mid-menstrual cycle. In oligomenorrhoeic patients too, FDG uptake may be high and resemble the uptake values found at mid-cycle [8]. In menstruating patients, the physiological cause of uptake may be sorted out by discussing the menstruation history with the patient. However, as ovarian transposition is carried out primarily in patients with gynaecological malignancies that are often post-hysterectomy, their ovulatory-phase cannot be determined by history alone. Therefore, when detecting a focal increased abdominal uptake on PET in a young female patient, the possibility of a functional uptake in a transposed ovary should be born in mind and adjacent surgical clips should be looked for on the CT part of the study. Directly interviewing the patient may also assist, as unfortunately the information of ovarian transposition is often omitted from the referral sheath for a PET/CT study. It was, indeed, not provided in any of our patients. In conclusion, the physician interpreting a PET/CT study should be familiar both with the clinical history and the imaging findings of ovarian transposition.


Increased FDG uptake in a transposed ovary may be related to its preserved functionality.

References 1. Morice P, Juncker L, Rey A, El-Hassan J, Haie-Meder C, Castaigne D. Ovarian transposition for patients with cervical carcinoma treated by radiosurgical combination. Fertil Steril 2000;74:743–8. 2. Newbold R, Safrit H, Cooper C. Surgical lateral ovarian transposition: CT appearance. AJR Am J Roentgenol 1990;154:119–20. 3. Goldberg RE, Sturgeon JF. Surgically transposed ovary presenting as an intraperitoneal mass on computed tomography. Can Assoc Radiol J 1995;46:229–30. 4. Kier R, Chambers SK, Kier R, Chambers SK. Surgical transposition of the ovaries: imaging findings in 14 patients. AJR Am J Roentgenol 1989;153:1003–6. 5. Bashist B, Friedman WN, Killackey MA. Surgical transposition of the ovary: radiologic appearance. Radiology 1989;173:857–60. 6. McCall ML, Keaty EC, Thompson JD. Conservation of ovarian tissue in the treatment of the carcinoma of the cervix with radical surgery. Am J Obstet Gynecol 1958;75:590–600. 7. Rohren EM, Turkington TG, Coleman RE. Clinical applications of PET in oncology. Radiology 2004;231: 305–32. 8. Lerman H, Metser U, Grisaru D, Fishman A, Lievshitz G, Even-Sapir E. Normal and abnormal 18F-FDG endometrial and ovarian uptake in pre- and postmenopausal patients: assessment by PET/CT. J Nucl Med 2004;45:266–71.

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An audit of imaging test utilization for the management of lymphoma in an oncology hospital: implications for resource planning? A SCHWARTZ, BSc, M K GOSPODAROWICZ, A KELLER, MD and R W TSANG, MD

MD,

K KHALILI,

MD,

M PINTILIE,

MSc,

S GODDARD,

BSc,

University of Toronto, Princess Margaret Hospital, University Health Network, 610 University Avenue, Toronto, Ontario, M5G 2M9 Canada ABSTRACT. The purpose of this study was to assist with resource planning by examining the pattern of physician utilization of imaging procedures for lymphoma patients in a dedicated oncology hospital. The proportion of imaging tests ordered for routine follow up with no specific clinical indication was quantified, with specific attention to CT scans. A 3-month audit was performed. The reasons for ordering all imaging procedures (X-rays, CT scans, ultrasound, nuclear scan and MRI) were determined through a retrospective chart review. 411 lymphoma patients had 686 assessments (sets of imaging tests) and 981 procedures (individual imaging tests). Most procedures were CT scans (52%) and chest radiographs (30%). The most common reasons for ordering imaging were assessing response (23%), and investigating new symptoms (19%). Routine follow up constituted 21% of the assessments (142/686), and of these, 82% were chest radiographs (116/142), while 24% (34/142) were CT scans. With analysis restricted to CT scans (296 assessments in 248 patients), the most common reason for ordering CT scans were response evaluation (40%), and suspicion of recurrence and/or new symptom (23%). Follow-up CT scans done with no clinical indication comprised 8% (25/296) of all CT assessments. Staging CT scans were under-represented at 6% of all assessments. Imaging with CT scans for follow up of asymptomatic patients is infrequent. However, scans done for staging new lymphoma patients were unexpectedly low in frequency, due to scans done elsewhere prior to referral. This analysis uncovered utilization patterns, helped resource planning and provided data to reduce unnecessary imaging procedures.

Lymphoma clinicians rely heavily on imaging techniques to determine the stage of disease at initial presentation, to assess the response to treatment and to follow the disease over time [1]. CT scans remain the standard for evaluation of nodal disease [2], while MRI gives additional information for some extranodal sites. Gallium scans and/or 18FDG-PET scans are also useful tools in the staging and follow-up where they help to distinguish residual fibrotic mass from viable lymphoma [3, 4]. After treatment has been completed and providing a complete remission has been achieved, the goal of follow-up investigations is to identify recurrent disease before symptoms develop [5]. However, routine CT imaging for follow up has not been shown to be costeffective, as investigation of symptoms is the most cited reason for finding recurrent disease [6–13]. Several studies documented that only 5–9% of relapses were imaging-detected before the development of symptoms [5, 11, 13]. At a dedicated oncology hospital, the policy for lymphoma patients has not been to perform routine CT imaging for follow up of asymptomatic patients beyond

Address correspondence to: Dr Richard W Tsang,

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Received 26 April 2005 Revised 1 June 2005 Accepted 22 June 2005 DOI: 10.1259/bjr/27372198 ’ 2006 The British Institute of Radiology

the attainment of complete remission of disease. Taking a chest radiograph has been left to the discretion of the attending physician. This is in agreement with published studies that suggested CT scan should be performed according to clinical indications, not strictly as routine actions [8, 10]. Edelman et al also proposed that ‘‘eliminating unnecessary testing would decrease the risk of further physical and psychological harm from the inevitable occurrence of false-positive tests’’ [9]. The subject of whether clinicians are optimally utilizing the available imaging modalities has been seldom studied [10]. This is particularly important in an environment of limited resources, as there is usually a waiting list to access certain imaging procedures such as CT and MRI scans. Concerns were expressed by the imaging department that the lymphoma group may be ordering an excessive number of scans unnecessarily for routine follow ups, hence making the resource less available for staging or other urgent reasons in a timely fashion. Therefore, in this study, the pattern of physician utilization of imaging investigations in the management of lymphoma was examined. The goal was to ascertain the indications for each imaging examination, for example: staging, evaluating response to treatment, suspected or confirmed recurrence, and routine The British Journal of Radiology, February 2006

Imaging utilization in lymphoma patients

follow-up monitoring. The aim was to determine the relative frequencies of the utilization of various imaging modalities, for follow-up monitoring versus for staging and response assessment. There was specific interest in determining if CT scans were often requested for routine follow-up in asymptomatic patients, to understand if this resource was overutilized, possibly at the expense of patients who may require the scans more urgently for assessment of disease.

Methods A 3-month audit of imaging procedures performed on lymphoma patients from January 1st to March 31st 2003 at a dedicated oncology hospital was performed. The Research and Ethics Board of the hospital approved the study. Patients were identified from the Imaging Department database and all were listed with a diagnosis of lymphoma. For this study, patients with a diagnosis of myeloma, leukaemia (acute and chronic) and benign haematological conditions were excluded. A record of all plain radiographs, CT scans, MRI, gallium, mammograms, bone scans, and ultrasound examinations were kept for the 3-month period. Patient demographics, disease extent, and treatment information were collected on each patient through a chart review. Details of histology, Ann Arbor stage, treatment, and response were abstracted. The oncologist responsible for each patient was recorded. For the purpose of this study, a ‘‘procedure’’ was considered a single imaging examination. For example, CT thorax, CT abdomen/pelvis, gallium scan and ultrasound were each counted as separate procedures (total: four procedures). An ‘‘assessment’’ was defined as a set of imaging examinations all carried out over a 2 weeks period and requested for the same purpose. By definition, CT thorax, CT abdomen/pelvis, gallium scan and ultrasound, if all done for staging, were counted as one assessment. A ‘‘new’’ patient in this study was a patient referred with a new diagnosis of lymphoma. Imaging tests performed up to 6 months after completion of definitive therapy were counted as performed for a ‘‘new patient’’. An ‘‘old’’ patient was defined as one who had had imaging performed more than 6 months after completing initial therapy. A patient with a previous diagnosis of lymphoma, but referred for

management of relapse beyond 6 months of completing initial therapy was considered an ‘‘old’’ patient. If a patient was under observation, for example in asymptomatic advanced stage follicular lymphoma, and the observation period lasted for more than 6 months from the time of referral, they would then be considered an ‘‘old’’ patient as well. The purpose of this distinction was to separate imaging utilization between patients referred with a new diagnosis of lymphoma for management (new), and those beyond the stage of initial treatment and attainment of complete remission (old). The indication for ordering each assessment was determined based on the physician’s clinical notes in the medical record. Reasons for ordering imaging assessments were categorized into: staging, response assessment, evaluation of residual disease, investigation of new symptoms, suspicion of recurrence, routine follow-up with no specific clinical indication, procedure related assessments such as biopsies, assessments mandated by study protocol, assessments performed for unrelated medical problems, assessments recommended by radiologists, surveillance for a secondary malignancy and assessments done for treatment complications. Questionable cases were reviewed by additional clinicians and a reason assigned by consensus.

Results 411 patients were included in the study. Patient characteristics are shown in Table 1. The most common non-Hodgkin’s lymphoma histologies were diffuse large B-cell lymphoma (22%), follicular lymphoma (20%), and others (23%). The initial diagnosis date is shown in Figure 1. 50% of the patients were diagnosed before Table 1. Patient characteristics at initial diagnosis (n5411) Characteristic

Number (%)

Age Gender

Median 47 years (range 8–97 years) Male 222 (54.0%) Female 189 (46.0%) Hodgkin’s disease 140 (34.1%) NHL 271 (65.9%) I–II 237 (57.7%) III–IV 174 (42.3%)

Diagnosis Ann Arbor stage

NHL, non-Hodgkin’s lymphoma.

Figure 1. Patient’s initial diagnosis date.


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A Schwartz, M K Gospodarowicz, K Khalili et al

Figure 2. Number of assessments per patient within the 3 month period.

2000, and 24% in 2002, which had the greatest proportion of patients diagnosed in a single year. The group of 411 patients generated a total of 686 imaging assessments and 981 procedures within the 3month period. Of these assessments, 25% (171/686) were for ‘‘new’’ patients and 75% (515/686) were for ‘‘old’’ patients. Most patients had one assessment (72%, Figure 2), while the majority of assessments (70.1%) consisted of one procedure (Figure 3). The total number of procedures performed per patient within the 3-month period is shown in Figure 4. CT scans constituted 52% of imaging procedures performed on lymphoma patients, followed by chest radiographs (30%), while others account for ,10% each (Figure 5). The most common indications for assessments were response assessment (23%), investigation of new symptoms (19%), and routine follow-up (21%) (Figure 6). Staging constituted only 4% of assessments (Figure 6). Other indications for scans accounted for 18% of

assessments, but could be broken down into procedure related (5%), study protocol (4%), unrelated medical problem (4%), recommended by radiologist (3%), surveillance for a secondary malignancy (1%) and investigation of treatment complications (1%). A comparison of the indications for assessments between ‘‘new’’ and ‘‘old’’ patients is shown in Figure 7. The largest differences between the new and old patients are in staging with a 14.2% difference, response assessment with a 25.6% difference and routine follow-up with a 26% difference.

Imaging requested for routine follow-up Within the 3-month period, 140 patients had routine follow-up imaging with no specific clinical indication. These patients received 142 assessments and 152 procedures, which comprised 16% of all procedures.

Figure 3. Number of imaging procedures per assessment.

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Figure 4. Total number of procedures per patient within the 3 month period.

The types of imaging procedures are shown in Table 2. Chest radiographs accounted for the majority (82%) The follow-up chest radiographs constituted 116 assessments of 295 performed over the 3-month period (39% of chest radiograph assessments).

Utilization of CT scans CT scans were performed on 248 patients. These patients received 296 assessments and 513 procedures (Table 3). 30% of these patients were ‘‘new’’ and 70% were ‘‘old’’. Over the 3-month period, the majority of patients (85%) received one CT assessment (Table 3), but

each assessment may involve 1–3 CT procedures (Table 4). The indications for CT scans included 40% for response assessment, 13% for suspicion of recurrence, 11% for residual disease, 10% for investigation of new symptoms, 8% for routine follow up, 6% for staging and 12% for other reasons (Figure 8). When comparing the indications for CT between ‘‘new’’ and ‘‘old’’ patients, the largest differences were 38% for response assessment, and 17% for staging, and a difference of 17% for suspicion of recurrence and 12% for routine follow up (Figure 9). All the routine follow up CT assessments (n525) were done for ‘‘old’’ patients, and accounted for 12.3% of the CT scan assessments done for ‘‘old’’ patients. Medical oncologists requested 92% (23/25) and radiation oncologists requested 8% (2/25) of the routine follow up CT scans. Of the follow-up CT scans, 72% (18/25) were in patients diagnosed in 2000–2003.

Discussion

Figure 5. Imaging procedures performed during the 3-month study on lymphoma patients. This graph shows the percentage of procedures that each imaging examination comprises.


At the time when this study was initiated, there were two main concerns at the hospital regarding the utilization of imaging resources by the lymphoma group. A first concern for clinicians was that scans ordered as staging investigations might overwhelm the imaging resource, as it is known that all new patients must be staged with imaging examinations [14], specifically CT scans of head and neck, thorax, abdomen and pelvis [1]. The results showed that staging accounts for only 4% of all the imaging assessments, and for an analysis restricted to CT scans it was 6% of assessments. The differences between the ‘‘new’’ patients and ‘‘old’’ patients showed an expected trend of ‘‘new’’ patients receiving more assessments for staging and evaluation of response. However, even for ‘‘new’’ patients the utilization of imaging for staging is low, and since all patients are staged with imaging, this implied that the majority had initial imaging performed prior to referral. The audit was conducted at a tertiary oncology hospital, with 119


Figure 6. Reason for ordering assessment for cohort of patients during 3-month period.

Figure 7. Comparing reasons for assessments between ‘‘new’’ and ‘‘old’’ patients.

Table 2. Follow up imaging assessments with no specific clinical indication (140 patients with 142 assessments) Procedure

No. of assessments

Percentage

Chest radiograph CT scan Ultrasound Total

116 25a 1 142

81.7% 17.6% 0.7% 100%

Table 3. Utilization of CT scans (248 patients with 296 imaging assessments and 513 individual procedures) Number of assessments performed per patient

Number of patients (%)

1 2 .2

211 (85.1%) 30 (12.1%) 7 (2.8%)

Number of CT scan procedures per patient

Number of patients (%)

1 2 3 .3

83 95 54 16

(33.5%) (38.3%) (21.7%) (6.5%)

many patients seen by external specialists and hence were fully assessed with imaging procedures prior to their referral. This is especially true for patients with stage I–II disease referred for radiation therapy. The Radiation Oncology Department saw 88 new patients in the same 3-month period of this study. Of these patients, 120

Table 4. CT scan procedures (n5513) per CT assessment Number of CT procedures per assessment

Number of assessments (%)

1 2 3

130 (43.9%) 115 (38.9%) 51 (17.2%)

40% (35/88) received radiation therapy. Only 26% (9/35) of those who received radiation therapy were staged with imaging procedures at the study hospital and included in the audit. This infers that 74% of new patients who received radiation therapy had staging scans performed elsewhere prior to their referral and were not even included in this study. ‘‘Old’’ patients had proportionately more assessments for residual disease, suspicion of recurrence and routine follow up. These trends are easily understood by the definition used for ‘‘new’’ patients, as those actively receiving their primary treatment, or those within 6 months of treatment completion when scans are performed to document response. A second concern stems from the waiting time for accessing CT scans, which was up to 2–3 weeks from the time of the request at the time this study was conducted. It was important to determine if there was a disproportionately large number of patients being scanned for routine follow-up with no specific clinical indication, thereby making the resource less available to requests for more urgent reasons. It was anticipated that a reduction in routine follow-up scans would free up resources and therefore reduce the waiting time for scans. In this study, imaging assessments performed as part of routine The British Journal of Radiology, February 2006


Figure 8. Analysis of CT scans, reasons for ordering assessments.

Figure 9. Comparing reasons for CT scans between ‘‘new’’ and ‘‘old’’ patients.

follow-up with no discernable clinical indication accounted for 21% of the total assessments and 16% of the total procedures. Follow-up represents a large proportion of assessments when all imaging procedures are grouped together, but chest radiographs account for the majority of follow-up procedures (82%), and only 24% consisted of CT scan procedures. A chest radiograph is less costly and more widely available compared with a CT scan [8, 9, 11, 15, 16]. Indeed, for routine follow-up, chest radiographs were more frequently ordered by clinicians compared with CT scans in this study, but it is less sensitive compared with a thorax CT scan. Studies have found that CT scans are minimally effective for follow-up in identifying relapses [1, 10, 17] as relapses are most often detected by patients developing diseaserelated symptoms [6–9, 11–13, 18]. Therefore, this implies that the practice of ordering routine chest radiographs has questionable clinical benefit, although one study did suggest a role in following Hodgkin’s disease in the first 3 years after treatment [12]. Perhaps the common practice of using chest radiographs as follow up is more due to its wide availability, low cost, and minimal X-ray exposure. The majority of CT scans were performed for ‘‘old’’ patients. Most of these patients received one assessment (85%), but for two-thirds of patients this assessment constituted more then 1 CT procedure. The reason for ordering CT scans (Figure 8) differed slightly from the reason for ordering all of the imaging procedures when considered together (Figure 6). CT scans were used more to evaluate the response to treatment, rather than to evaluate new problems or for routine follow up. Perhaps this reflected a high degree of success in the initial primary management of lymphomas. The data showed The British Journal of Radiology, February 2006

that although follow-up accounted for 21% of all imaging assessments, only 25 CT assessments were done for routine follow-up within the 3-month period (i.e. 8% of all CT assessments). Of these, 92% were ordered by medical oncologists and 8% by radiation oncologists. The reason for this discrepancy could be due to differences in the patient factors seen by the two specialties (a higher patient volume, with a higher risk of relapse in more advanced stage patients seen by medical oncologists) or, alternatively, patients seen by radiation oncology may be more likely to have follow up imaging performed elsewhere. Many studies have suggested surveillance follow-up routines based on effectiveness, both in terms of the ability to detect relapse and cost [2, 5, 7–9, 11–13]. Not all of the recommended standards are in agreement, ranging from basic history, physical examination and serum lactic dehydrogenase (LDH) tests for follow up [8], which is similar to the practice at our hospital, to a combination of physical examination, blood work, chest radiograph and additional imaging tests such as CT and gallium scans left to the discretion of the investigator [11]. The United States National Comprehensive Cancer Network 2004 practice guidelines for Hodgkin’s disease, which are based on consensus rather than published data, outline an even more intensive follow up routine that entails chest imaging (CT scan or radiograph), to be performed every 3–6 months during the first 3 years post-treatment and annually from the 4th year after treatment [19]. Abdominopelvic imaging was recommended every 6–12 months in the first 3 years post treatment, and annually in years 4 and 5. For patients treated with radiation therapy to the chest, mammographic screening was suggested 8–10 years 121


post-therapy. The follow up guidelines for NHL were not detailed, except for follicular lymphoma where follow up imaging was regarded as necessary but ordered as clinically indicated, about every 6 months [20]. An international workshop that established standardized response criteria in NHL stated that ‘‘imaging studies may be added for relevant clinical indications, but specific tests cannot be currently recommended’’, yet acknowledges that the issue is still controversial and that good clinical judgement is the most important component of patient follow up [2]. In one study of Hodgkin’s disease, among patients with recurrence of disease, imaging-detected cases did not have a better overall clinical outcome with salvage therapy compared with patients whose recurrence were detected by symptoms [12]. A limitation of this study is the retrospective nature of the review, as it is possible that a clinician had ordered an imaging procedure with a legitimate indication, but did not document this in the medical record either before or after the imaging procedure was performed. Such a situation will be misclassified under the category of ‘‘routine follow up with no specific clinical indication’’. Given this limitation, the 8% rate of utilization of CT scans for this reason is probably acceptable within a practice environment where there had been significant variation in the follow up recommendations as cited above. In addition, other institutions with different referral pattern and case-mix will invariably find a very different spectrum for the reasons behind imaging utilization. However, it is the feasibility of the methodology and the potential usefulness of the auditing procedure in assessing and assigning resource utilization in this study that should be emphasised. Increasingly, stakeholders of the healthcare delivery system such as government, health authorities and hospital boards demand accurate utilization data to assign resources, and the type of information requested is often along the same vein as that provided in this study. In conclusion, in the lymphoma patients seen at a tertiary oncology hospital, imaging assessments requested for staging are under-represented. A substantial proportion of patients having had imaging tests completed elsewhere prior to referral explained this. Imaging requested for routine follow up of asymptomatic lymphoma patients is infrequent, apart from chest radiographs. This study reflects the utilization patterns of imaging within a disease group and would assist in planning the assignment of imaging resources based on case-mix. It also reassured the physicians and the institution that the majority of CT scans were ordered for valid indications. It is hoped that the study also raised the awareness of clinicians in the importance of continually adhering to proper indications for ordering imaging tests.

References 1. Schiepers C, Filmont JE, Czernin J. PET for staging of Hodgkin’s disease and non-Hodgkin’s lymphoma. Eur J Nucl Med Mol Imaging 2003;30 Suppl. 1:S82–8. 2. Cheson BD, Horning SJ, Coiffier B, Shipp MA, Fisher RI, Connors JM, et al. Report of an international workshop to standardize response criteria for non-Hodgkin’s

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lymphomas. NCI Sponsored International Working Group. J Clin Oncol 1999;17:1244–53. Naumann R, Vaic A, Beuthien-Baumann B, Bredow J, Kropp J, Kittner T, et al. Prognostic value of positron emission tomography in the evaluation of post-treatment residual mass in patients with Hodgkin’s disease and nonHodgkin’s lymphoma. Br J Haematol 2001;115:793–800. Even-Sapir E, Israel O. Gallium-67 scintigraphy: a cornerstone in functional imaging of lymphoma. Eur J Nucl Med Mol Imaging 2003;30 Suppl. 1:S65–81. Weeks JC, Yeap BY, Canellos GP, Shipp MA. Value of follow-up procedures in patients with large-cell lymphoma who achieve a complete remission. J Clin Oncol 1991;9:1196–203. Biasotti S, Garaventa A, Padovani P, Faraci M, Fioredda F, Hanau G, et al. Role of active follow-up for early diagnosis of relapse after elective end of therapies. Pediatr Blood Cancer 2005;45:781–6. Radford JA, Eardley A, Woodman C, Crowther D. Follow up policy after treatment for Hodgkin’s disease: too many clinic visits and routine tests? A review of hospital records. BMJ 1997;314:343–6. Elis A, Blickstein D, Klein O, Eliav-Ronen R, Manor Y, Lishner M. Detection of relapse in non-Hodgkin’s lymphoma: role of routine follow-up studies. Am J Hematol 2002;69:41–4. Edelman MJ, Meyers FJ, Siegel D. The utility of follow-up testing after curative cancer therapy. A critical review and economic analysis. J Gen Intern Med 1997;12:318–31. Cohen MD, Siddiqui A, Weetman R, Provisor A, Coates T. Hodgkin disease and non-Hodgkin lymphomas in children: utilization of radiological modalities. Radiology 1986;158:499–505. Dryver ET, Jernstrom H, Tompkins K, Buckstein R, Imrie KR. Follow-up of patients with Hodgkin’s disease following curative treatment: the routine CT scan is of little value. Br J Cancer 2003;89:482–6. Torrey MJ, Poen JC, Hoppe RT. Detection of relapse in early-stage Hodgkin’s disease: role of routine follow-up studies. J Clin Oncol 1997;15:1123–30. Guppy AE, Tebbutt NC, Norman A, Cunningham D. The role of surveillance CT scans in patients with diffuse large B-cell non-Hodgkin’s lymphoma. Leuk Lymphoma 2003;44:123–5. Vinnicombe SJ, Reznek RH. Computerised tomography in the staging of Hodgkin’s disease and non-Hodgkin’s lymphoma. Eur J Nucl Med Mol Imaging 2003;30 Suppl. 1:S42–55. Valk PE, Pounds TR, Tesar RD, Hopkins DM, Haseman MK. Cost-effectiveness of PET imaging in clinical oncology. Nucl Med Biol 1996;23:737–43. Joines JD, McNutt RA, Carey TS, Deyo RA, Rouhani R. Finding cancer in primary care outpatients with low back pain: a comparison of diagnostic strategies. J Gen Intern Med 2001;16:14–23. Rankin SC. Assessment of response to therapy using conventional imaging. Eur J Nucl Med Mol Imaging 2003;30 Suppl. 1:S56–64. Oh YK, Ha CS, Samuels BI, Cabanillas F, Hess MA, Cox JD. Stages I-III follicular lymphoma: role of CT of the abdomen and pelvis in follow-up studies. Radiology 1999;210:483–6. NCCN. The NCCN Clinical Practice Guidelines in Oncology. Hodgkin’s Disease. www.nccn.org: National Comprehensive Cancer Network; 2004. Report No.: V.1.2004. NCCN. The NCCN Clinical Practice Guidelines in Oncology. Non-Hodgkin’s Lymphoma. www.nccn.org: National Comprehensive Cancer Network; 2005. Report No.: v.1.2005.



Image quality and breast dose of 24 screen–film combinations for mammography 1

A D DIMAKOPOULOU, MSc, 2I A TSALAFOUTAS, 1 E N YAKOUMAKIS, PhD

PhD,

1

E K GEORGIOU,

MD, PhD

and

1

Medical Physics Department, Medical School, University of Athens, 75 Mikras Asias, 115 27, Athens and 2Medical Physics Unit, Konstantopoulio-Agia Olga Hospital, 3-5 Agias Olgas, Nea Ionia, 142 33, Athens, Greece ABSTRACT. In this study the effect of different mammographic screen–film combinations on image quality and breast dose, and the correlation between the various image quality parameters, breast dose and the sensitometric parameters of a film were investigated. Three Agfa (MR5-II, HDR, HT), two Kodak (Min-R M, Min-R 2000), one Fuji (AD-M), one Konica (CM-H) and one Ferrania (HM plus) single emulsion mammographic films were combined with three intensifying screens (Agfa HDS, Kodak Min-R 2190 and Fuji AD-MA). The film characteristics were determined by sensitometry, while the image quality and the dose to the breast of the resulting 24 screen–film combinations were assessed using a mammography quality control phantom. For each combination, three images of the phantom were acquired with optical density within three different ranges. Two observers assessed the quality of the 72 phantom images obtained, while the breast dose was calculated from the exposure data required for each image. Large differences among screen–film combinations in terms of image quality and breast dose were identified however, that, could not be correlated with the film’s sensitometric characteristics. All films presented the best resolution when combined with the HDS screen at the expense of speed, and the largest speed when combined with the AD-MA screen, without degradation of the overall image quality. However, an ideal screen–film combination presenting the best image quality with the least dose was not identified. It is also worth mentioning that the best performance for a film was not necessarily obtained when this was combined with the screen provided by the same manufacturer. The results of this study clearly demonstrate that comparison of films based on their sensitometric characteristics are of limited value for clinical practice, as their performance is strongly affected by the screens with which they are combined.

The main concern in mammography screening is the detection of features characteristic of breast disease. These features often have sizes of the order of 1 mm and differ from the normal tissue only slightly in composition, thus setting high requirements for the resolution and contrast that an imaging system must offer in order to be appropriate for mammography [1]. On the other hand, given the high radiosensitivity of the breast and the large number of women examined many times during their life, it is evident that the doses during mammography should be kept as low as possible. While digital mammography may look promising, the vast majority of mammography examinations are still carried out with screen–film systems. In recent years, most film manufacturers have presented new films and intensifying screens for mammography that reduce the dose to the breast and produce the image quality required to maintain the diagnostic sensitivity and specificity of mammography at high levels. However, while the design is the major factor in determining the performance of a film, this may be affected by the processing conditions, such as the chemicals used, their The British Journal of Radiology, February 2006


temperature and the processing time [2]. Inappropriate chemicals or a developing temperature lower than recommended may result in unacceptable mammograms and this is why some films have been modified to be less dependent on processing conditions [3, 4]. Film characteristics can be determined and monitored for changes due to processing by sensitometry. However, film performance will be dependent on the screen with which it is combined and thus for clinical practice the characteristics of the screen–film combination rather than those of film or screen separately are of interest [5]. The screen–film characteristics can be determined and monitored using an appropriate quality control (QC) phantom, with which changes in image quality due to processing or other reasons can be identified. One of the parameters routinely monitored with the QC phantom is the background or reference optical density (OD) of the mammographic images. Apart from the personal preferences of radiologists, it has been shown that for a given screen–film combination, subtle details and small contrast differences are best accentuated when the film OD is within a certain range [6–8]. For 123

A D Dimakopoulou, I A Tsalafoutas, E K Georgiou and E N Yakoumakis

this reason it has been recommended that each institution should determine the optimum OD for the screen– film used and the processing conditions specific to it [7]. In this study, eight films were combined with three intensifying screens and the resulting 24 screen–film combinations were compared in terms of image quality and breast dose. Film characteristics were determined by sensitometry, whereas the image quality and speed of the screen–film combinations were assessed using a QC phantom to obtain images within three different OD ranges. Our main objective was to investigate the effect of different screens on a certain film and search for any correlation between the image quality, breast dose and the sensitometric parameters of a film.

Materials and methods The eight single emulsion mammographic films tested in this study were the MR5-II, HDR, HT (Agfa-Gevaert N.V., Mortsel, Belgium), Min-R M, Min-R 2000 (Eastman Kodak Company, New York, NY), AD-M (Fuji Photo Film Co. Ltd, Tokyo, Japan), CM-H (Konica Corporation, Tokyo, Japan) and HM plus (Ferrania Sp A, Ferrania (SV), Italy). One box from each film type was used to avoid little differences that may exist among different film batches or films that may have been stored for different times and under different storage conditions [9]. The three intensifying screens used in this study were the HDS (Agfa), the Min-R 2190 (Kodak) and the AD-MA (Fuji). For each film a 21-step sensitometric strip was produced, using an X-Rite 334 sensitometer (X-Rite, Grandville, MI) operated in the green spectrum. All films were processed in a daylight processor (Curix Capacity, Agfa) with nominal processing time 90 s (22 s developing time) and with the developer temperature set to 36 ˚C. The developer type was the Eos Dev (Agfa) and the fixer type was the G334i (Agfa). All films were processed sequentially, immediately after exposure and on the same day, to avoid day-to-day variations in processing conditions caused by the ageing of the chemicals that may have variable effects on the characteristics of each film [9, 10]. The OD of the 21 steps of the sensitometric strips was measured using a calibrated optical densitometer (RMI 331, X-Rite). For each film the Hurter-Driffield (H&D) curve was plotted and the following sensitometric parameters were derived: OD of base plus fog (ODb+f), maximum OD (ODmax), average gradient (AG), film gamma (c) and film speed. The AG and c are the slopes of the H&D curve for ODs from 0.25+ODb+f to 2.0+ODb+f and from 1.0+ODb+f to 2.0+ODb+f, respectively. AG and c are both used as indices of film contrast, however, only c can be used to reproduce the linear part of the H&D curve. The film speed was defined as the reciprocal of the relative light exposure required to obtain an OD of 1+ODb+f. Using this definition, the higher the film speed the less exposure is needed for a given OD. In order to illustrate the expected increase in breast dose – according to sensitometry – when a film other than the fastest one is used, the sensitometric relative dose index (SRDI) was defined as the reciprocal of the relative speed value. The SDRIs were expressed as percentages of the smallest SRDI value (highest speed) that was taken as 100%. It must be noted that from preliminary sensitometric tests 124

it has been confirmed that for different sheets of the same film type processed within the same day, variations of less than ¡0.01 in ODb+f, ¡10% in speed and ¡0.1 in ODmax, AG and c should be expected. To evaluate the characteristics of screen–film combinations, a mammography QC phantom was employed (breast phantom, Model 18-222; Nuclear Associates, Division of Victoreen Inc., NY). This phantom is realistically shaped and equivalent to an average firm breast of 4.5 cm compressed thickness, consisting of 50% adipose and 50% glandular tissue. It includes 12 groups of calcium carbonate specks (simulating microcalcifications), 7 hemispheric masses composed of 75% glandular and 25% adipose equivalent tissue (simulating tumours) and a wax insert with 5 embedded nylon fibres (simulating glandular tissue fibrils). The phantom also contains a five-step stepwedge, simulating breast areas with compositions 100% adipose, 70% adipose–30% glandular, 50% adipose–50% glandular, 30% adipose– 70% glandular and 100% glandular tissue. Finally, two line-pair test targets (5–20 lp mm21 each), one parallel and one perpendicular to the anode–cathode axis and a central area where the background OD is measured, are included. A similar phantom (without the nylon fibres and with only one line-pair test target) was used by Nassivera and Nardin [11]. The phantom was exposed using a Senographe 500T mammography unit (CGR, Buc, France). All exposures were made with the Mo/Mo target filter combination, constant tube potential (28 kVp), large focal spot (0.3 mm nominal size) and without the breast compression paddle. Using manual mAs selection technique, images of the phantom were acquired until for each screen–film combination three films with OD as close as possible to the central OD of three different optical density ranges (0.70–1.10, 1.11–1.50, 1.51–2.00) were produced to account for the wide range of ODs that can be encountered within actual mammographs. Phantom images were processed in the same processor and on the same day as the sensitometric strips, so both film characteristics and screen–film performance were determined under the same processing conditions. It must be noted that as the Kodak and Fuji cassettes were not compatible with the Agfa daylight processor, the films exposed with these cassettes had to be transferred manually to the Agfa cassette in order to be fed into the processor. For each one of the resulting 72 phantom images, the OD of the central area was measured with the densitometer, as well as the OD of the areas simulating 100% adipose and 100% glandular tissue. The OD difference of these two areas can be used as an index of screen–film contrast (CI) and, according to the phantom manufacturer, it should be ¢0.28. All phantom images were examined using a viewing box especially designed for mammography, featuring adjustable brightness, masking shutters and a magnifying glass. The shutters were closed down to the phantom image size and the brightness was adjusted as necessary to obtain the best possible conditions for viewing each type of simulated lesion, while for speck groups the magnifying glass was also used. A magnifying glass supplied with the QC phantom was used to inspect the line pair object. The above details are mentioned, as The British Journal of Radiology, February 2006

Image quality and breast dose of screen–film mammography

Figure 1. The Hurter-Driffield (H&D) curves of the eight films included in this study are given. The optical densities (ODs) of the 21 steps of the sensitometric strips correspond to log relative exposure values (LogE) that range from 0 to 3, in steps of 0.15 each. In these figures, only the ODs for LogE¢ge;0.9 (steps 7 to 21) are presented in order to enhance the visibility of the differences in the linear part and the shoulder of the H&D curves.

viewing conditions are very important for interpreting mammograms or scoring phantom images [12]. Two observers scored the images independently and any disagreements were resolved by consensus. Five scores were recorded for each film: one for the speckgroups, one for masses, one for fibres and two for the two line-pair test targets. For ambiguous decisions concerning not clearly visualized structures, a 0.5 mark was assigned. In order to have a single index characterizing the screen–film performance, a total score (TS) was calculated using the following weighting coefficients: 0.4 for specks, 0.35 for masses and 0.25 for fibres. These coefficients were selected after consulting with five radiologists about the clinical importance of each simulated structure for diagnosis. Since the two scores for the line-pair test targets have no straightforward clinical relevance, their mean value was calculated for reference only (resolution score). For screen–film combination comparisons in terms of breast dose, the entrance surface air-kerma (ESAK) at the phantom surface was calculated from the mAs selected for each exposure. For the range of mAs selections

utilized in this study, the output at 28 kVp defined at the phantom entrance surface was 98¡2 mGy mAs21. Furthermore, the ESAK required to achieve a net OD of 1 was calculated by interpolation from the ODs and the ESAKs of the three films acquired for each combination. The resulting ESAKs for a net OD51 were used to derive the relative dose index (RDI), expressed as a percentage of the smallest observed value, that was considered as 100%. Using this definition, the larger the RDI, the larger the dose to the breast and the smaller the speed of a given screen–film combination. To investigate the correlation between the various image quality parameters, sensitometric parameters, OD and dose indices linear regression analysis was used. A correlation coefficient (r) larger than 0.7 was taken as an indication of good correlation. Specifically, the correlations of all the image quality scores (TS, specks, masses, fibres and resolution) with AG, c, SRDI, CI, OD and ESAK were investigated. Furthermore, the correlation of TS with resolution, the correlations of CI with AG, c, SDRI, OD and ESAK and the correlations of SDRI with RDI and ESAK were also investigated.

Results The H&D curves for all the films studied are plotted in Figure 1, while their sensitometric parameters are given in Table 1. In Figure 1, the large differences among the H&D curve shapes, the high speed of Min-R 2000, the low speed of AD-M and the non-typical but similar H&D curve shapes of HDR and HM plus should be noted. From Table 1, it can be seen that the Min-R 2000 presents the highest speed and c, while the CM-H has the highest AG. Concerning the screen–film comparisons, for the 72 phantom images evaluated in this study, speck scores ranged from 6 to 11, mass scores from 2.5 to 6, fibre scores from 2 to 4, total scores from 4.3 to 7.3, resolution scores from 10 to 14.5 lp mm21 and CI from 0.26 to 0.61. For each of the three OD ranges the respective ESAK ranges were: 2.0–4.9 mGy, 2.4–6.2 mGy and 3.1–8.8 mGy. The results for the 24 combinations studied in terms of the ESAK at the phantom surface, the background OD, the TS, the CI and the resolution score are presented in Tables 2a, 2b and 2c for the three different OD ranges, respectively. In the two last columns of each table, the screen–film RDI and the mean value of TS in the three OD ranges (TSm) are also given. From these tables it can be seen that the screen affected the image quality of a given film as well as the dose

Table 1. The results of the sensitometric evaluation of the eight mammographic films included in this study. The lowest speed was arbitrarily defined as 100%. The sensitometric relative dose index (SRDI) was defined as the reciprocal of the relative speed, considering the lowest value (highest speed) as 100% Film Parameter

MR5-II (Agfa)

HT (Agfa)

HDR (Agfa)

Min-R M (Kodak)

Min-R CM-H 2000 (Kodak) (Konica)

AD-M (Fuji)

HM plus (Ferrania)

ODb+f ODmax AG c Rel. Speed (%) SRDI (%)

0.20 3.50 2.46 2.80 115 145

0.23 3.84 3.12 4.36 121 138

0.21 3.98 3.09 3.98 143 117

0.17 3.88 3.13 4.13 133 125

0.20 4.03 2.80 4.46 167 100

0.17 3.58 3.50 4.24 100 167

0.22 4.13 2.93 3.68 143 117

0.20 3.80 3.64 4.23 126 132

Agfa-Gevaert N.V., Mortsel, Belgium; Eastman Kodak Company, New York, NY; Fuji Photo Film Co. Ltd, Tokyo, Japan; Konica Corporation, Tokyo, Japan; Ferrania USA Inc., USA. OD, optical density; AG, average gradient; c, film gamma.


125

A D Dimakopoulou, I A Tsalafoutas, E K Georgiou and E N Yakoumakis Table 2a. The results of the evaluation of the 24 film–screen combinations, from the quality control (QC) phantom images with optical density (OD) in the range 0.7–1.1. The screen–film combination relative dose index (RDI) has been expressed in terms of the entrance surface air-kerma (ESAK) required to produce a net OD of 1, considering the lowest value observed (2.19 mGy) as 100% Film

Screen

ESAK (mGy)

OD bgnd

TS

CI

Resolution (lp mm21)

Screen RDI (%)

Film TSm

MR5-II (Agfa)

HDS (Agfa) Min-R2190 (Kodak) AD-MA (Fuji) HDS (Agfa) Min-R2190 (Kodak) AD-MA (Fuji) HDS (Agfa) Min-R2190 (Kodak) AD-MA (Fuji) HDS (Agfa) Min-R2190 (Kodak) AD-MA (Fuji) HDS (Agfa) Min-R2190 (Kodak) AD-MA (Fuji) HDS (Agfa) Min-R2190 (Kodak) AD-MA (Fuji) HDS (Agfa) Min-R2190 (Kodak) AD-MA (Fuji) HDS (Agfa) Min-R2190 (Kodak) AD-MA (Fuji)

3.9 3.1 2.4 3.9 3.1 2.4 3.9 3.1 2.4 3.1 3.1 2.4 3.1 2.4 2.0 3.1 2.4 2.0 4.9 4.9 3.1 3.1 3.1 2.0

1.07 0.98 1.04 1.06 0.88 1.00 1.04 0.88 1.00 0.81 0.98 1.05 1.08 0.86 1.02 1.04 0.79 0.96 0.97 0.88 0.83 0.79 0.92 0.76

4.7 4.5 4.7 5.7 5.4 5.8 5.6 5.9 5.2 5.7 5.4 5.4 6.0 5.0 5.5 5.7 5.5 6.0 6.0 6.5 7.0 5.6 6.3 4.7

0.31 0.31 0.30 0.39 0.36 0.36 0.42 0.37 0.41 0.33 0.38 0.41 0.38 0.26 0.35 0.45 0.36 0.41 0.43 0.41 0.38 0.31 0.38 0.31

13.5 12.0 11.5 13.5 12.5 12.5 13.5 12.5 12.0 13.0 13.0 12.5 14.0 12.0 11.5 13.5 11.5 11.5 12.5 12.0 12.0 13.5 12.5 11.5

204 177 132 207 178 127 200 179 128 177 163 120 156 141 100 153 140 104 272 270 191 189 169 120

4.7 4.7 5.1 5.9 5.7 5.9 5.7 6.2 5.3 5.8 5.7 6.0 5.7 5.6 6.1 5.8 5.7 5.9 6.0 6.6 7.0 5.0 6.1 5.6

HT (Agfa) HDR (Agfa) MIN-R M (Kodak) MIN-R 2000 (Kodak) CM-H (Konica) AD-M (Fuji) HM plus (Ferrania)

Agfa-Gevaert N.V., Mortsel, Belgium; Eastman Kodak Company, New York, NY; Fuji Photo Film Co. Ltd, Tokyo, Japan; Konica Corporation, Tokyo, Japan; Ferrania USA Inc., USA. TS, total score; CI, index of screen–film contrast.

Table 2b. The results of the evaluation of the 24 film–screen combinations, from the quality control (QC) phantom images with optical density (OD) in the range 1.11–1.5. The screen–film combination relative dose index (RDI) has been expressed in terms of the entrance surface air-kerma (ESAK) required to produce a net OD of 1, considering the lowest value observed (2.19 mGy) as 100% Film

Screen

ESAK (mGy)

OD bgnd

TS

CI

Resolution (lp mm21)

Screen RDI (%)

Film TSm

MR5-II (Agfa)


4.9 4.9 3.1 4.9 3.9 3.1 4.9 4.4 3.1 3.9 3.9 3.1 3.9 3.1 2.4 3.9 3.1 2.4 6.2 6.2 4.9 4.9 3.9 3.1

1.30 1.44 1.25 1.26 1.24 1.45 1.38 1.37 1.39 1.27 1.33 1.49 1.39 1.23 1.40 1.48 1.20 1.24 1.22 1.37 1.36 1.47 1.38 1.48

5.0 4.9 5.2 6.2 5.8 6.0 5.8 6.1 5.4 6.0 6.1 6.6 5.8 5.7 6.4 6.0 5.5 5.4 5.8 6.5 6.6 5.0 6.3 6.1

0.31 0.30 0.31 0.45 0.48 0.48 0.50 0.50 0.50 0.52 0.48 0.50 0.47 0.44 0.46 0.49 0.44 0.47 0.47 0.49 0.50 0.50 0.47 0.49

12.5 12.0 12.0 13.5 12.5 12.5 13.5 13.0 13.0 13.5 13.0 12.5 14.5 13.5 11.5 14.0 13.5 12.0 14.0 13.5 13.0 14.5 12.5 12.5

204 177 132 207 178 127 200 179 128 177 163 120 156 141 100 153 140 104 272 270 191 189 169 120

4.7 4.7 5.1 5.9 5.7 5.9 5.7 6.2 5.3 5.8 5.7 6.0 5.7 5.6 6.1 5.8 5.7 5.9 6.0 6.6 7.0 5.0 6.1 5.6



126


Image quality and breast dose of screen–film mammography Table 2c. The results of the evaluation of the 24 film–screen combinations, from the quality control (QC) phantom images with optical density (OD) in the range 1.51–2.0. The screen–film combination relative dose index (RDI) has been expressed in terms of the entrance surface air-kerma (ESAK) required to produce a net OD of 1, considering the lowest value observed (2.19 mGy) as 100% Film

Screen

ESAK (mGy)

OD bgnd

TS

CI

Resolution (lp/mm)

Screen RDI (%)

Film TSm

MR5-II (Agfa)


7.8 6.2 3.9 6.2 4.9 3.9 6.2 6.2 3.9 6.2 4.9 3.9 4.9 3.9 3.1 4.9 3.9 3.1 8.8 7.8 6.2 6.2 4.9 3.9

1.76 1.71 1.55 1.74 1.59 1.84 1.73 1.94 1.77 1.82 1.58 1.75 1.72 1.54 1.72 1.79 1.59 1.71 1.72 1.56 1.78 1.88 1.63 1.89

4.3 4.9 5.2 5.8 5.7 6.0 5.6 6.4 5.4 5.7 5.6 6.1 5.4 6.2 6.5 5.8 6.2 6.2 6.0 6.7 7.3 4.4 5.9 6.0

0.29 0.32 0.32 0.55 0.52 0.52 0.54 0.56 0.49 0.52 0.49 0.51 0.61 0.50 0.57 0.50 0.48 0.50 0.55 0.52 0.55 0.48 0.51 0.47

10.0 12.0 11.5 13.0 12.0 12.0 13.0 11.5 12.0 12.5 13.0 12.5 14.5 12.5 13.0 14.0 13.5 13.0 13.5 13.0 14.0 13.0 13.5 12.0

204 177 132 207 178 127 200 179 128 177 163 120 156 141 100 153 140 104 272 270 191 189 169 120

4.7 4.7 5.1 5.9 5.7 5.9 5.7 6.2 5.3 5.8 5.7 6.0 5.7 5.6 6.1 5.8 5.7 5.9 6.0 6.6 7.0 5.0 6.1 5.6



required for obtaining a phantom image of a certain OD. Furthermore, it can be seen that for a certain screen–film combination the TS, CI and resolution were dependent on the background OD. The screen–film combinations that exhibited the largest dependency with OD were the ADMA/HM plus for the TS, the Min-R 2190/Min-R 2000 for the CI and the HDS/MR5-II for the resolution score. In order to highlight some of the major results of this study, the best screen for a given film and the best film for a given screen in terms of TS, CI, resolution and RDI are given in Tables 3 and 4, respectively, where the OD range for which the best score is obtained is also noted. It

is evident that, while for TS and CI there was variety in the screen with which a film was best combined, for resolution and RDI the best screen was common for all films. Indeed, all films presented the best resolution when combined with the HDS screen and the smallest RDI (largest speed) when combined with the AD-MA screen. The largest resolution offered by the HDS screen was at the expense of speed, while the largest speed offered by the AD-MA screen was at the expense of resolution but not at the expense of TS. In summary, Tables 3 and 4 clearly demonstrate two points that deserve special attention. First, the highest

Table 3. The best screen for a given film in terms of total score (TS), index of screen–film contrast (CI), line pairs per millimetre (lp mm21) and relative dose index (RDI). Letters in parentheses give reference to Tables 2a, 2b or 2c where each value can be found. The largest values overall for each parameter are given in bold

TS

CI lp mm21

RDI

HDS (Agfa) Min-R2190 (Kodak) AD-MA (Fuji) HDS (Agfa) Min-R2190 (Kodak) AD-MA (Fuji) HDS (Agfa) Min-R2190 (Kodak) AD-MA (Fuji) HDS (Agfa) Min-R2190 (Kodak) AD-MA (Fuji)

MR5-II (Agfa)

HT (Agfa)

HDR (Agfa)

MIN-R M (Kodak)

MIN-R 2000 CM-H (Kodak) (Konica)

AD-M (Fuji)

HM plus (Ferrania)

– – 5.2 (b,c) – 0.32 (c) 0.32 (c) 13.5 (a) – – – – 132

6.2 (b) – – 0.55 (c) – – 13.5 (a,b) – – – – 127

– 6.4 (c) – – 0.56 (c) – 13.5 (a,b) – – – – 128

– – 6.6 (b) 0.52 (b,c) – – 13.5 (b) – – – – 120

– – 6.5 (c) 0.61 (c) – – 14.5 (b,c) – – – – 100

– – 7.3 (c) 0.55 (c) – 0.55 (c) 14 (b) – 14 (b) – – 191

– 6.3 (a,b) – – 0.51 (c) – 14.5 (b) – – – – 120

– 6.2 (c) 6.2 (c) 0.5 (c) – 0.5 (c) 14 (b,c) – – – – 104

Agfa-Gevaert N.V., Mortsel, Belgium; Eastman Kodak Company, New York, NY; Fuji Photo Film Co. Ltd, Tokyo, Japan; Konica Corporation, Tokyo, Japan; Ferrania USA Inc., USA.


127

A D Dimakopoulou, I A Tsalafoutas, E K Georgiou and E N Yakoumakis Table 4. The best film for a given screen in terms of total score (TS), index of screen–film contrast (CI), line pairs per millimetre (lp mm21) and relative dose index (RDI). Letters in parentheses give reference to Tables 2a, 2b or 2c where each value can be found. The largest values overall for each parameter are given in bold

TS

CI lp mm21

RDI

HDS (Agfa) Min-R2190 (Kodak) AD-MA (Fuji) HDS (Agfa) Min-R2190 (Kodak) AD–MA (Fuji) HDS (Agfa) Min–R2190 (Kodak) AD–MA (Fuji) HDS (Agfa) Min–R2190 (Kodak) AD–MA (Fuji)

MR5-II (Agfa)

HT (Agfa)

HDR (Agfa)

MIN-R M (Kodak)

MIN-R 2000 CM-H (Kodak) (Konica)

AD-M (Fuji)

HM plus (Ferrania)

– – – – – – – – – – – –

6.2 (b) – – – – – – – – – – –

– – – – 0.56 (c) – – – – – – –

– – – – – – – – – – – –

– – – 0.61 – 0.57 14.5 13.5 – – – 100

– 6.7 (c) 7.3 (c) – – – – 13.5 (b) 14 (b) – – –

– – – – – – 14.5 (b) 13.5 (b) – – – –

(c) (c) (b,c) (b)

– – – – – – – 13.5 (b) – 153 140 –

Agfa-Gevaert N.V., Mortsel, Belgium; Eastman Kodak Company, New York, NY; Fuji Photo Film Co. Ltd, Tokyo, Japan; Konica Corporation, Tokyo, Japan; Ferrania USA Inc., USA.

TS, CI, resolution and the lowest RDI were observed for different combinations and thus an ‘‘ideal’’ combination was not identified. Second, the best performance of a given film or screen was not always obtained when combined with the screen or film, respectively, from the same manufacturer. Finally, concerning the correlation between image quality parameters, sensitometric parameters and dose, no correlation coefficient larger than 0.7 was calculated in any of the correlations tested. The largest correlation coefficient calculated was that between CI and OD, which was 0.66 but increased to 0.87 when the MR5-II combinations were excluded, demonstrating that for modern high contrast films the screen–film contrast increases with OD. It must be clarified, however, that this correlation has been assessed for ODs up to 2.0 and it is expected that for higher ODs the CI will start to decrease again as the films become saturated. The poor correlations of image quality parameters (TS, specks, masses, fibres, resolution and CI) with film contrast (c, AG) and the poor correlation of RDI with SRDI, confirmed that film performance is strongly affected by the screen. Concerning the lack of correlation of TS with CI, resolution and dose, the following remarks should be made. While combinations with high TS generally had high CI, there were many cases with high CI and low TS. There were also many combinations with high TS that, however, exhibited low resolution score and vice versa. Finally, slow combinations did not always give high TS, as expected according to the general principle that the higher the dose the lower the quantum mottle.

Discussion The major conclusion of this study was that film characteristics are modified by intensifying screens in such a significant and variable way, that comparisons among films based on the manufacturer’s specifications or sensitometry are of limited value. Indeed, a film with given technical specifications or sensitometric characteristics, when combined with different screens may exhibit improved or degraded performance. 128

Instructive of the variable effect that a screen may have on a film, it can be seen that while Min-R 2000 was the fastest film and remained the fastest when combined with the Fuji screen, the RDI ratio of CM-H and Min-R 2000 combinations with Fuji screen was 1.04 while for film only the respective ratio of SDRI was 1.32. That means that the Fuji screen spectral emission better matched the spectral sensitivity of CM-H compared with Min-R 2000. The slowest film, according to sensitometry, was the AD-M, which remained the slowest when combined with all screens. The SDRI for AD-M was 167 while its smallest RDI was 191 and it was observed when combined with the Fuji screen. Examples of the largest variations observed in score and dose when a film was combined with different screens are: the HM plus where the TS was 36% larger with the Fuji than with the Agfa screen (Table 2c), the Min-R 2000 where the CI was 46% and 22% larger with the Agfa than with the Kodak screen (Tables 2a and 2c, respectively), the Min-R 2000 where the resolution was 26% larger with the Agfa than with the Fuji screen (Table 2b) and the HT (Agfa) where 63% more dose is required with the Agfa screen than with the Fuji screen. As previously mentioned, from Tables 2–4 some conclusions may be drawn concerning the superiority of certain combinations over others in terms of image quality or speed. However, the absolute values of scores and other screen–film characteristics may be quite different on other mammographic facilities, given the strong dependence of film characteristics on processing conditions [13, 14]. This must be emphasised, as the objective of this study was not to recommend or condemn certain films or screens but to investigate the effect of screens on the performance of films. Although most screen–film comparisons in the literature have been carried out using the same processing conditions for all films, it must be noted that the general notion is that a film would perform optimally when it is processed according to the recommendations of the manufacturer. Even so, this does not annul the fact that the breast dose and image quality for a film optimally processed will again vary, depending on the screen with which it is combined, and that some films will be affected by the screen more than others. The British Journal of Radiology, February 2006

Image quality and breast dose of screen–film mammography

Even if it were assumed that the processing conditions were optimal for all films, it would again be difficult to select the best screen–film combination from those studied, as there are no established criteria about what increase in breast dose is justified by a superior image quality. For example the AD-MA/AD-M presented a TS of 7.3 (11 specks, 6 masses, 3 fibres) and an ESAK of 6.2 mGy while the AD-MA/Min-R 2000 a TS of 6.4 (9 specks, 5 masses, 4 fibres) with an ESAK of 2.4 mGy. To conclude which is the best combination, one has to decide if the 14% increase in TS could justify the 158% increase in breast dose. The same question still holds when considering that certain combinations (as the ADMA/AD-M) exhibited slightly larger TS for larger ODs but with disproportional increase in breast dose. An important remark should also be made concerning the OD of the films studied. It is obvious that the films included in Table 2a are of too low OD and few of the films included in Table 2c are of too high OD, compared with the target OD range of 1.3 to 1.8 proposed for mammography [1]. Nonetheless, certain combinations exhibited better scores in Table 2a than in Tables 2b and 2c, while most of the films of Table 2c with ODs larger than 1.8 exhibited scores similar to those of Table 2b. In clinical practice, however, given that the wide OD variations within a mammogram are not uncommon, some areas may present similar ODs with those of Table 2a or larger than 1.8 and therefore the performance of a screen–film combination within all OD ranges is of interest. In this context, comparisons based on the TSm may be considered more relevant to the clinical situation than comparisons based on the TS within only one OD range. The variability of TS with OD should always be considered when selecting the central OD setting of the automatic exposure control (AEC) system based on the results of phantom scores. Some final comments should be made concerning the method used to assess the image quality of screen–film combinations. Phantom scoring does not always represent clinical practice, as in actual mammograms the performance of a given combination will be also dependent on the breast type [15]. Furthermore, phantom scoring may be somewhat biased, as it relies on the detection of structures known to be present at specific positions [16]. Nevertheless, phantoms are considered as the best way for the objective evaluation of image quality and various models with fixed or randomly positioned details are extensively used. Caldwell et al [17] agreed on the usefulness of such phantoms for the objective evaluation of image quality and also reported that a subjective assessment of image quality is better accomplished with an anthropomorphic breast phantom than with actual mammograms, where the variability among radiologists was higher. However, they noted that no significant correlation was found between the various methods used to evaluate image quality and concluded that more work is required to obtain an index of true image quality correlated with the probability of correct diagnosis. In conclusion, image quality and dose in mammography are more strongly dependent on screen–film combination than on film or screen separately. While sensitometry remains an important tool for determining and monitoring the film characteristics [18], it is of little value when the image quality and breast dose in clinical The British Journal of Radiology, February 2006

mammograms are of concern. Therefore, any change of film or screen type in a mammographic facility should be carefully investigated with a phantom, for determining the performance of the selected screen–film combination and for adjusting the AEC system to the optimum OD range for this combination.

References 1. van Woudenberg S, Thijssen M, Young K. European protocol for the quality control of the physical and technical aspects of mammography screening. In: Perry N, Broeders M, de Wolf C, Kirkpatrick A, Tornberg S, editors. European guidelines for quality assurance in mammography screening (3rd edn). Luxembourg: Office for Official Publications of the European Communities, 2001. 2. Brink C, De Villiers JFK, Lo¨tter MG, Van Zyl M. The influence of film processing temperature and time on mammography image quality. Br J Radiol 1993;66:685–90. 3. Tabar L and Haus AG. Processing of mammographic films: technical and clinical considerations. Radiology 1989;173:65–9. 4. Kimme-Smith C, Bassett LW, Gold RH, Zheutlin J, Gornbein JA. New mammography screen/film combinations: imaging characteristics and radiation dose. AJR Am J Roentgenol 1990;154:713–9. 5. Kirkpatrick AE, Law J. A comparative study of films and screens for mammography. Br J Radiol 1987;60:73–8. 6. Robson KJ, Kotre CJ, Faulkner K. The use of a contrastdetail test object in the optimization of optical density in mammography. Br J Radiol 1995;68:277–82. 7. McParland BG, Boyd MM, Yousef KAL. Optimizing optical density of a Kodak mammography film-screen combination with standard-cycle processing. Br J Radiol 1998;71:950–3. 8. McParland BJ. A comparison of two mammography filmscreen combinations designed for standard-cycle processing. Br J Radiol 1999;72:73–5. 9. Kimme-Smith C, Bassett LW, Gold RH, Chow S. Increased radiation dose at mammography due to prolonged exposure, delayed processing and increased film darkening. Radiology 1991;178:387–91. 10. Fernandez JM, Guibelalde E. Technical note: Physical evaluation of recent Kodak films for mammography. Br J Radiol 1993;66:828–32. 11. Nassivera E, Nardin L. Daily quality control programme in mammography. Br J Radiol 1996;69:148–52. 12. Pisano ED, Britt GG, Lin Y, Schell MJ, Burns CB, Brown ME. Factors affecting phantom scores at annual mammography facility inspections by the U.S. Food and Drug Administration. Acad Radiology 2001;8:864–70. 13. Tsalafoutas IA, Dimakopoulou AD, Koulentianos ED, Serefoglou AN, Yakoumakis EN. The variation of the sensitometric characteristics of seven mammographic films with processing conditions. Br J Radiol 2004;77:666–71. 14. Kimme-Smith C, Rotschild PA, Bassett LW, Gold RH, Moler C. Mammographic Film-Processor Temperature, Development Time and Chemistry. AJR Am J Roentgenol 1989;152:35–40. 15. Meeson S, Young KC, Rust A, Wallis MG, Cooke J, Ramsdale ML. Implications of using high contrast mammography X-ray film-screen combinations. Br J Radiol 2001;74:825–35. 16. Jackson VP, Harrill CD, White SJ, Gillespie KR, Mail JT, Katz BP. Evaluation of a dual-screen, dual-emulsion mammography system. AJR Am J Roentgenol 1989;152:483–6. 17. Caldwell CB, Fishell EK, Jong RA, Weiser WJ, Yaffe MJ. Evaluation of mammographic image quality: pilot study comparing five methods. AJR Am J Roentgenol 1992;159:295–301. 18. West MS and Spelic DC. Using light sensitometry to evaluate mammography film performance. Med Phys 2000;27:854–60.

129


The effect of phantom type, beam quality, field size and field position on X-ray scattering simulated using Monte Carlo techniques G McVEY,

DPhil

Joint Department of Physics, The Royal Marsden NHS Trust, Fulham Road, London SW3 6JJ, UK ABSTRACT. Determining the amount of scatter inside and outside a diagnostic X-ray room is important for evaluating the dose to staff and the public. The amount of scatter is affected by many physical factors including beam quality and field size. However, there is little published data on patient scatter and there are large differences between the available data sets. Hence, a Monte Carlo code was developed to allow a systematic study of the factors affecting patient scatter. A voxel phantom was used to provide a realistic model of the patient. The variation of scatter with different phantom types was investigated to show the effect of patient inhomogeneities and obliquities. The effect of altering tube voltage, filtration, voltage ripple, field size and position on patient scatter was studied. A larger than expected variation in the patient scatter was observed with increasing field area due to the proximity of the field borders with the patient obliquities. The effect of the tube voltage ripple on the patient scatter was also calculated. This showed that there would be little effect on the scatter levels within X-ray rooms if ageing X-ray generators, which produce substantial voltage ripple, were replaced by X-ray tubes with modern medium frequency generators. Recommendations are made on the choice of published scatter data for X-ray room design.

Scatter is produced by all materials in a diagnostic Xray room, with the main source of scattered radiation being the interaction of X-rays with the patient [1]. However, there is a limited amount of scatter data available for use in X-ray room design. The data were obtained from measurements undertaken with tissue equivalent slab phantoms [2]; with human-shaped homogeneous phantoms [3] and with heterogeneous phantoms such as the RANDO phantom [4–6]. Since these studies used a variety of phantom types and technique parameters, there was a large variation in the scatter values reported. An alternative solution is to use Monte Carlo computer simulations of the scatter produced by a model of human anatomy. This has enabled three systematic studies of the effect of different parameters on scatter from patients undergoing chest posteroanterior (PA), lumbar spine anteroposterior (AP) and lumbar spine lateral (LAT) radiographic examinations. For the first study, four voxel phantoms (P1 to P4) have been used to simulate the patient as shown in Figure 1. These different approaches have been followed to investigate the effect of patient obliquities and inhomogeneities on scatter. The first approach was to use a voxel phantom (P1) reconstructed from CT data. It was developed by Zubal et al [7, 8] and was recently used in a Monte Carlo model to optimize image quality Current address: North Wales Medical Physics, Glan Clwyd Hospital, Bodelwyddan, Denbighshire LL18 5UJ, UK. This work was supported by a grant from Anglia and Oxford Health Authority.

130

Received 25 October 2004 Revised 7 June 2005 Accepted 15 June 2005 DOI: 10.1259/bjr/59998010 ’ 2006 The British Institute of Radiology

and patient dose in chest and lumbar spine radiography [9, 10]. Dance et al [11] showed that Zubal’s voxel phantom was representative of a patient undergoing chest and lumbar spine radiographic examinations. The second approach was to use the voxel phantom developed by Zubal and change all the voxels inside the patient contour to be soft tissue and those outside to be air (P2). The third approach was to use the voxel phantom with all the voxels within the phantom to be soft tissue (P3). The fourth voxel phantom was developed as a block of soft tissue specified by the average dimensions of Zubal’s voxel phantom (P4). The second study used the Monte Carlo code to calculate the effect of varying the imaging parameters on the scatter from the patient model (P1): the tube voltage (60–150 kV), tube filtration (2.5–7.0 mmAl) and voltage ripple (0–50%). By studying the effect of voltage ripple on patient scatter, it can be observed whether replacing an old X-ray generator, which has substantial voltage ripple, with a modern X-ray generator, which has negligible voltage ripple, will make a significant difference to the scatter levels inside and outside X-ray rooms. The third study used the Monte Carlo code to calculate the effect on the scatter from the patient model (P1) of varying the field area (25–1225 cm2) and the position of the field on the patient. This study generalized the results for the chest and lumbar spine regions so that the data may be interpreted for other X-ray examinations. The calculated scatter values obtained in this work may be used to aid the design of X-ray rooms, but they may also assist in the analysis of the doses received by staff The British Journal of Radiology, February 2006

Simulating scatter from patients

Figure 1. The simulation model used to calculate the scatter from a patient undergoing, for example, a chest posteroanterior X-ray examination. The position of the detector is shown at a scattering angle of 135 ˚. The three other phantoms used in the calculations are also shown below.

who undertake and assist with interventional radiological examinations.

Methods and materials Voxel Monte Carlo code The Monte Carlo code is similar to that used previously to study image quality and patient dose in radiographic examinations [9, 10], but was extended to simulate the scatter surrounding a voxel phantom. The program transports the photons through the voxel phantom; a collision density estimator [12] is used to provide an efficient method of calculating scatter. The model calculates the air kerma at points 1 m from the phantom surface for scattering angles between 30 ˚ and 150 ˚. Scatter ratios were determined by the air kerma at each of these points divided by the incident air kerma without backscatter. The scatter ratios are expressed as percentages. A large number of photon histories were used to calculate this parameter so its uncertainty was less than ¡1% (1 standard deviation). The patient model was a voxel phantom (P1) derived from segmented CT data [7, 8]. Each voxel belonged to 1 of 55 organs [10]. The tissue type of each organ was specified as one of average soft tissue, healthy lung, bone The British Journal of Radiology, February 2006

or bone spongiosa. The calculations used tissue densities and compositions taken from the International Commission on Radiation Units and Measurements (ICRU) Report No. 46 [13], except for bone which was taken from Kramer [14]. The patient support device, i.e. the chest stand for the chest examination or couch top for the lumbar spine examinations, was included in the voxel phantom by the addition of an extra layer of voxels. Table 1 shows the thickness and composition of the chest stand and couch top. The dimensions of the voxel phantom were 89.9 cm long, 35.6 cm wide and 21.4 cm thick. As the lower limbs were not present in the phantom, its length was determined to be equivalent to the height of the average European male in sitting position. The shoulder width and chest thickness were determined after an initial study [11, 15] which compared calculations with measurements of patient entrance air kerma. Figure 1 shows the computer model of a patient undergoing a radiographic examination for which the scatter was calculated. The model included the X-ray spectrum from the X-ray tube, the patient and the couch top or chest stand. The X-ray spectra were calculated using a Birch and Marshall [16] model. The grid and the screen–film imaging system were not included in the model. This means that the Monte Carlo model will produce significantly greater forward scatter than would 131

G McVey Table 1. The parameters used for the chest and lumbar spine imaging systems Imaging system parameters Parameter type

Chest

Lumbar spine

Tube voltage (kV) Filtration (mmAl) Target angle ( ˚ ) Voltage ripple (%) Focus film distance (cm) Focus surface distance (cm) Field size at focus surface distance (cm6cm)

141 5.7 15 0 150 127 30.6621.2

Chest stand/couch top

6.0 mm of wood

72 (for AP projection),77 (for LAT projection) 4.7 12 0 146 112 (for AP projection); 98 (for LAT projection) 27.5610.9 (for AP projection); 23.469.2 (for LAT projection) 1.2 mm Al

AP, anteroposterior; LAT, lateral.

be observed clinically if a grid and film cassette were present (or grid and image intensifier for fluoroscopic imaging systems). The scatter from the grid and film cassette would have been negligible as the patient significantly attenuates the X-ray beam. Therefore, the forward directed patient scatter calculated by the Monte Carlo model is a conservative estimate of the clinical situation. Table 1 shows the imaging system parameters. These parameters was found to provide good image quality in a recent EU clinical trial [17, 18] and were thus used as a reference system to observe the differences in patient scatter when the imaging system parameters were varied. In the first study, the effect of the patient heterogeneities and obliquities on scatter were investigated. Therefore, in addition to the patient model (P1) described above, the other three phantoms shown in Figure 1 were used to simulate scatter from a patient undergoing chest PA, lumbar spine AP and lumbar spine lateral examinations. Phantom P2 was defined with the voxels inside the patient’s surface set to average soft tissue and those outside the surface set to air. Phantom P3 was a slab phantom defined with all voxels, apart from the chest stand or couch top, set to be average soft tissue. Phantom P4 was also a slab phantom defined by the average thickness (z direction) and width (y direction) of the patient model within the field borders for each projection. Table 2 shows the dimensions of the P4 phantoms including the chest stand or couch top. This was undertaken as the shoulder width was considerably larger than the width further down the phantom’s body outline. Hence, it was interesting to determine which slab phantom (P3 or P4) scatter approximated the scatter from the patient model most closely. The patient model

Table 2. The dimensions of the homogeneous voxel phantoms (P4) defined by the average patient model dimensions Phantom dimensions (cm) Examination

Length

Width

Thickness

Chest PA Lumbar spine AP Lumbar spine LAT

89.9 89.9 89.9

22.0 24.0 17.6

16.0 18.0 24.4

PA, posteroanterior; AP, anteroposterior; LAT, lateral.

132

(P1) was used for the other two studies described in the introduction.

Validation of the patient model Sandborg et al [9] and McVey et al [10] describe the use of the voxel Monte Carlo code to simulate image quality and patient dose. As part of this work, Dance et al [11] and Sandborg et al [15] compared measurements of optical density behind phantoms and patient entrance air kerma with calculations using the Monte Carlo code for both of these situations. The good agreement obtained from the comparisons showed that the voxel phantom (P1) was representative of a patient undergoing chest and lumbar spine X-ray examinations [11, 15].

Simulation of Williams’ scattering experiment This section describes the method used to compare the scatter calculated using the voxel Monte Carlo code with the scatter measured by Williams [5]. This was carried out to validate the calculations against recent independently published values and also to check the reliability of Williams’ measured values. Williams measured the scatter from the abdominal and pelvic sections of a RANDO phantom. Therefore, a voxelized cylinder of Alderson Muscle A material was used to simulate the RANDO phantom of dimensions 50.0 cm long, 25.0 cm wide and 21.5 cm thick. Its composition and density were obtained from ICRU Report No. 44 [19]. Williams [5] measured the scatter in terms of air kerma normalized to the dose–area product (DAP). Therefore, a DAP meter and the air between it and the phantom surface were included in the voxel model. McVey [20] showed that these materials produce a significant amount of scatter. The DAP meter was modelled as a solid block of Perspex with dimensions: 16.4 cm length, 18.1 cm width and 1.7 cm thickness. An average density of 0.315 g cm23 was used for the Perspex as the DAP meter was constructed from layers of Perspex 0.2 cm thick with an air gap between them. The DAP meter was assumed to be at a distance of 26.6 cm from the X-ray focus and the focus to surface distance (FSD) was 80 cm. The incident field at the phantom surface was 22 cm long and 17.5 cm wide. The The British Journal of Radiology, February 2006


scatter was calculated at points 1 m from the centre of the phantom for scattering angles between 30 ˚ and 150 ˚.

was determined to be small, being 3.7% [20]. Thus, the scatter calculated by the voxel Monte Carlo code should be reasonably representative of the scatter levels found in clinical X-ray rooms.

Results and discussion Validation of the scatter calculations The scatter calculations were validated by comparing the scatter at points surrounding a block of solid water calculated with the voxel Monte Carlo code to the values calculated by an EGS4 Monte Carlo code for the same geometry as described by McVey [20]. Good agreement (within 2%) was shown between the values calculated by the two codes for photon energies between 20 keV and 150 keV and for tube voltages between 50 kV and 120 kV. The voxel Monte Carlo code could not be used to simulate the scatter measurements previously carried out in a clinical X-ray room as detailed by McVey and Weatherburn [1]. This was due to the geometrical limitations of the code and the size of the simulated model. McVey and Weatherburn [1] used the EGS4 Monte Carlo code to calculate the scatter from solid water blocks placed within a simulated X-ray room and showed reasonable agreement with the measured scatter. For this simulation, the percentage scatter contribution from the X-ray room walls to the total calculated scatter

Dependence of percentage scatter on phantom type Figure 2 shows the scatter from the four phantoms calculated for the chest PA, lumbar spine AP and lumbar spine LAT projections, respectively. The scatter for the patient (P1) and the contoured phantom (P2) lie between the values for the thick slab phantom (P3) and the slab phantom with average dimensions (P4) below scattering angles of 62 ˚ for the lumbar spine AP view, below 108 ˚ for the chest PA view and below 125 ˚ for the lumbar spine LAT view. Therefore, the phantom with the average dimensions (P4) can provide a conservative estimate of the scatter from a patient (P1) for all scattering angles for the chest PA and lumbar spine LAT exams and for scattering angles less than 67 ˚ and greater than 131 ˚ for the lumbar spine AP exam. By comparing Figures 2a–c, it can be seen that the chest PA examination produced the largest amount of scatter as the highest tube voltage and largest field area were employed. The lungs also attenuated the scattered X-rays less than soft tissue. The lumbar spine AP

Figure 2. The variation of the percentage scatter with different phantom types (P1 to P4) for (a) the chest posteroanterior projection, (b) the lumbar spine anteroposterior projection and (c) the lumbar spine lateral projection.


133

G McVey Table 3. Percentage scatter for different phantom types (P1 to P4) relative to the percentage scatter for the patient model (P1) Percentage scatter for a 45 ˚ scattering angle Phantom type

Chest PA exam

Lumbar spine AP exam

Lumbar spine LAT exam

Thick slab phantom (P3) Contoured phantom (P2) Patient model (P1) Slab phantom with average dimensions (P4)

0.62 1.12 1.00 1.36

0.56 1.29 1.00 1.46

0.70 1.14 1.00 1.22

Percentage scatter for a 120 ˚ scattering angle Phantom type

Chest PA exam



Thick slab phantom (P3) Contoured phantom (P2) Patient model (P1) Slab phantom with average dimensions (P4)

1.16 1.17 1.00 1.32

0.91 1.04 1.00 0.99

0.96 1.08 1.00 1.22


projection produced the least scatter in the forward direction as it had the lowest tube voltage and provided attenuation by a large thickness of soft tissue. The lumbar spine AP projection produced more scatter in the backward direction than the lumbar spine LAT projection as it had a larger field area. The largest differences between the different phantom types occurred between scattering angles of 30 ˚ and 87 ˚. The largest difference was 77% for the lumbar spine AP projection at 87 ˚. Table 3 shows the scatter for all the phantoms relative to the patient model for 45 ˚ and 120 ˚ scattering angles as examples of forward and backward directed scatter. The thick slab phantom (P3) produced the least scatter in the forward direction as it had the largest thickness and width and, therefore, greatly attenuated the scattered X-rays. The contoured phantom (P2) produced more scatter than the thick slab phantom in the forward direction. The phantom obliquity attenuated the scattered X-rays less. The contoured phantom’s thickness and width varied along its length, and in places these were larger than the phantom with average dimensions (P4). Thus, the contoured phantom produced less scatter than the phantom with average dimensions.

The scatter from the patient model (P1) was lower than from the contoured phantom (P2) in the forward direction. Bone attenuated the scattered X-rays more than soft tissue in all the examinations. In the chest examination, fewer scatter interactions occurred in lung than soft tissue as its density was lower. For different examinations, Table 3 shows that the differences between the phantom types were larger in the forward direction for the lumbar spine AP projection compared with those for the chest examination as a lower tube voltage was used for the lumbar spine AP projection. For the lumbar spine LAT examination, the differences were not as large as would be expected for the large thickness of the patient in the lateral projection. The field edge was close to the phantom boundary. Therefore, there was less tissue to attenuate the forward scattered photons. Figure 2 shows that the phantom type had less effect in the backward direction than in the forward direction for the frontal projections. The scatter was produced near the entrance surface of a phantom [20]. For the lateral projection, changes in the phantom width were more significant. The effect of the tissue inhomogeneities in the

Table 4. The variation of percentage scatter from the patient (P1) for different tube voltages normalized to the scatter values for reference imaging systems Percentage scatter relative to the reference imaging system values 120 ˚ scattering angle

45 ˚ scattering angle kV

60 70 72 77 80 90 95 102 110 130 141 150

Chest PA exam

0.68 0.76 0.83 0.95 1.00 1.03



0.64 0.91 1.00

Chest PA exam



0.90

0.87 0.97 1.00

0.94

1.23 1.51

1.00 1.05 1.17 1.22

1.06 1.15

1.00 1.04 1.10 1.12

1.97

1.36

1.23

1.18

0.89 0.92 0.94 0.99 1.00 1.01


134



backward direction was similar to their effect in the forward direction for the chest PA and lumbar spine LAT projections. For the lumbar spine AP projection, there was a larger difference between the forward and backward directed scatter as bone volume occupied a greater proportion of the irradiated volume than for the other projections.

Dependence of percentage scatter on tube voltage Table 4 shows the variation of scatter normalized to the reference system values for X-ray tube voltages between 60 kV and 150 kV for the different examinations. The largest variation was for forward directed scatter in all projections. The forward directed scatter became more penetrating with increasing tube voltage. The majority of backward scattered X-rays were produced close to the entrance surface [20]. Therefore, there was less of an effect for increasing tube voltage. The percentage scatter for the lumbar spine AP projection had the largest variation with tube voltage. This projection had the lowest reference system tube voltage of 72 kV. Therefore, increasing the tube voltage had a large effect. The percentage scatter in the forward direction increased by a factor of 2 for an increase in the tube voltage of 38 kV. The lumbar spine LAT projection had a smaller variation with tube voltage than the AP projection. The LAT projection had a reference system tube voltage of

77 kV and the field was positioned closer to the patient’s edge than the AP projection. The chest PA projection had a smaller variation with tube voltage than both the lumbar spine projections. This was due to the high reference system tube voltage of 141 kV and because the lungs attenuated the scattered X-rays less than tissue in the chest PA projection. Table 5 compares the variation of the lumbar spine AP and LAT values with those from McVey and Weatherburn [1], Trout and Kelley [3] and Williams [5]. There were large differences in the variations of forward directed scatter with tube voltage due to differences in the attenuating properties of the phantoms used. The differences in the variations were less in the backward direction as the scatter was produced close to the entrance surface of the phantom [20]. The variation of Trout’s scatter values was considerably larger than the other values. This was due to the 70 kV values being produced by a self-rectified X-ray tube which had a low beam quality. The variation of Williams’ scatter values was slightly less than the other values. This was due to the significant amount of scatter produced by the DAP meter which was independent of tube voltage [1].

Dependence of percentage scatter on tube filtration and voltage ripple Table 6 shows the variation of percentage scatter with tube filtration and voltage ripple. The filtration was

Table 5. A comparison of the variation of percentage scatter with tube voltage in the literature [1, 3, 5] Percentage scatter relative to 70 kV values for a 45 ˚ scattering angle Tube voltage (kV)

50 60 70 80 85 90 95 100 110 120 125 150

LS AP exam values

0.70 1.00 1.35

LS LAT exam values

Measured values [1]

Calculated values [1]

1.00 1.17

Trout and Kelley [3] values

Williams [5] values

0.52

0.84

1.00

1.00 1.13

1.65

1.30 1.35

2.16

1.51

4.76

1.28

6.19 8.10

1.51

Percentage scatter relative to 70 kV values for a 135 ˚ scattering angle Tube voltage (kV)

50 60 70 80 85 90 95 100 110 120 125 150

LS AP exam values

0.90 1.00 1.07

LS LAT exam values

1.00 1.10

Measured values [1]

Calculated values [1]

Trout and Kelley [3] values

Williams [5] values

0.77

0.80

0.79

0.87

1.00

1.00

1.00

1.00 1.08

1.16

1.15 1.17

1.22

1.22 1.26

1.75

1.10

2.00 2.08

1.23

1.30



135

G McVey Table 6. The variation of percentage scatter from the patient model (P1) for different filtrations and voltage ripples normalized to the reference system scatter values Percentage scatter normalized to reference values for a 45 ˚ scattering angle Chest PA exam Filtration (mmAl)



2.5

5.7

7.0

2.5

3.5

4.7

2.5

3.5

4.7

0.80 – – –

1.00 0.99 0.96 0.90

1.06 – – –

0.70 – – –

0.84 – – –

1.00 0.94 0.85 0.78

0.80 – – –

0.92 – – –

1.00 0.98 0.94 0.90

% Voltage Ripple

0 5 20 50

Percentage scatter normalized to reference values for a 120 ˚ scattering angle Chest PA exam Filtration (mmAl)



2.5

5.7

7.0

2.5

3.5

4.7

2.5

3.5

4.7

0.89 – – –

1.00 1.00 0.99 0.98

1.03 – – –

0.83 – – –

0.91 – – –

1.00 0.98 0.94 0.91

0.84 – – –

0.93 – – –

1.00 0.99 0.97 0.94

% Voltage Ripple

0 5 20 50


varied between 2.5 mmAl and 7.0 mmAl for the chest PA examination and between 2.5 mmAl and 4.7 mmAl for the lumbar spine examinations. A larger range in filtration was investigated for the chest PA examination due to the higher tube voltage employed. The voltage rectification was varied between 0% and 50% ripple for all examinations. Table 6 shows that the tube filtration affected the percentage scatter less than the tube voltage due to the smaller differences in the incident beam qualities simulated. The voltage ripple had a similar effect on the percentage scatter as the tube filtration. Therefore, changing an old X-ray generator with significant voltage ripple to a medium frequency X-ray generator will not produce significantly more scatter. The percentage scatter for the lumbar spine AP projection had the largest variation with filtration and voltage ripple. The scatter in the forward direction decreased by 15% if the filtration decreased by 1.2 mmAl or the voltage ripple decreased by 20%. The variations in percentage scatter for the lumbar spine LAT

and chest PA projections were less due to their higher beam qualities. Similar variations of scatter with different filtrations and voltage ripples were observed at 102 kV, 90 kV and 95 kV for the chest PA, lumbar spine AP and LAT projections, respectively. Therefore, the variations shown in Table 6 are applicable over a large range of tube voltages.

Dependence of percentage scatter on field area Table 7 shows the percentage scatter for a 100 cm2 square field area which was used to normalize the scatter values for the different field areas and field positions shown in Table 8. All the reference system parameters for each projection, as shown in Table 1, were employed except for the field area. Table 7 shows the percentage scatter in the lumbar spine LAT projection calculated with the field centre at two positions. First, the field centre was positioned at the same place as the reference

Table 7. The percentage scatter for the chest PA, lumbar spine AP and LAT reference imaging systems with 100 cm2 square field areas Percentage scatter Examination

Chest PA exam



Imaging system

Reference system

Reference system

Reference system

Field centred on centre of patient

0.044 0.035 0.027 0.024 0.049 0.059 0.067

0.012 0.011 0.013 0.024 0.048 0.057 0.065

0.039 0.048 0.050 0.050 0.058 0.062 0.066

0.014 0.021 0.026 0.031 0.041 0.049 0.056

Scattering angle ( ˚ )

30 45 60 87 120 135 150


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Simulating scatter from patients Table 8. Percentage scatter for different square field areas relative to the percentage scatter for the 100 cm2 square field areas given in Table 7 Percentage scatter normalized to the value for a 10 cm610 cm field size at a 45 ˚ scattering angle Exam

Chest PA exam




Imaging system

Reference system

Reference system

Reference system

Field centred on patient centre

0.25 1.00 2.20 4.07 6.90 9.45 11.64

0.21 1.00 2.86 7.55 11.92 14.52 17.38

0.20 1.00 1.61

0.20 1.00 3.38

Field size (cm6cm)

Expected values

565 10610 15615 20620 25625 30630 35635

0.25 1.00 2.25 4.00 6.25 9.00 12.25

Percentage scatter normalized to the value for a 10 cm610 cm field size at a 120 ˚ scattering angle Exam

Chest PA exam




Imaging system

Reference system

Reference system

Reference system

Field centred on patient centre

0.26 1.00 2.22 3.94 6.01 7.78 9.34

0.25 1.00 2.25 3.86 5.19 6.13 7.16

0.26 1.00 1.63

0.26 1.00 2.29

Field size (cm6cm)

Expected values

565 10610 15615 20620 25625 30630 35635

0.25 1.00 2.25 4.00 6.25 9.00 12.25


system, which was close to the patient’s lateral boundary. Second, the field centre was positioned at the centre of the patient’s width. In both cases, the scatter was calculated at points on the side of the patient where the field was off-centre. Table 7 shows that the position of field centre had a considerable effect. The percentage scatter was larger for all scattering angles with the field centred on the patient’s obliquity as there was less tissue to attenuate the scattered radiation. The scatter for the field centred on the patient obliquity was 2.3 times greater than for the field centred on a thicker part of the patient for a scattering angle of 45 ˚. Trout and Kelley [3] found the scatter for the field centred at the phantom’s edge was between 6.5 and 1.8 times larger than the scatter for the field positioned at the centre of the phantom’s width for tube voltages between 50 kV and 150 kV at a 45 ˚ scattering angle. Shielding reports [6, 21] indicate a linear relationship between scatter and field area. Table 8 shows that this relationship is not valid for the variation of patient scatter with field area. For a 25 cm 6 25 cm field area incident in the chest PA view, the normalized scatter was 10.5% higher and 3.9% lower than expected for scattering angles of 45 ˚ and 120 ˚. The forward directed scatter values for the chest PA view were closest to the expected variation with field area. The largest differences were observed for the forward directed scatter in the lumbar spine AP view. For the same field area incident in the lumbar spine AP view, the normalized scatter was 90.7% higher and 17.0% lower than expected for scattering angles of 45 ˚ and 120 ˚. The British Journal of Radiology, February 2006

The variation of scatter with field area shown in Table 8 was due to the patient obliquities. A field width of 25 cm covered the majority of the patient’s trunk. The field edges were incident on the patient obliquities and thus, the scattered X-rays were less attenuated. The scatter was therefore larger than expected in the forward direction. The patient’s obliquity produced a small reduction in the amount of scatter produced in the backward direction due to there being less tissue. There was a greater difference for the lumbar spine view than the chest PA view due to the lower tube voltage. A smaller number of field areas were investigated for the lumbar spine LAT exam due to the small width of the patient in this orientation. Table 8 shows the normalized scatter values with the field centre close to the patient obliquity (reference system position) and the field positioned at the centre of the patient’s width. With the 225 cm2 field centred medially on the patient, the normalized values were 3.38 and 2.29 for scattering angles of 45 ˚ and 120 ˚. With the field moved laterally by 3 cm, the normalized scatter values were smaller and thus, there was a larger variation if the field was centred over a thick part of the patient than if the field was centred on the patient obliquity. However, the actual scatter values tended to be larger for the field positioned at the patient obliquity than at the patient centre as there was less attenuation of the scattered X-rays (Table 7). If the scatter was calculated at points on the other side of the patient than the field centre, then the increased attenuation would substantially reduce its amount. 137

G McVey

Figure 3. The variation of the normalized percentage scatter at a 45 ˚ scattering angle with square, rectangular and equivalent

square field sizes between 25 cm2 and 900 cm2 for (a) the lumbar spine anteroposterior projection and (b) the chest posteroanterior projection.

Figure 3 shows the variation of scatter with square, rectangular and equivalent square field areas between 25 cm2 and 900 cm2 for a 45 ˚ scattering angle for the lumbar spine AP and chest PA projections. The rectangular field length was kept constant at 30 cm and the width increased from 5 cm to 30 cm. In X-ray room design it is difficult to account for the variation of field dimensions used in the clinical situation. Therefore, one method investigated was to calculate the equivalent square field area (s2) using the equivalent square field length (s) and the rectangular field dimensions (x and y) shown in Equation (1) [22]. 2xy ð1Þ s~ ðxzyÞ Figure 3a shows large differences in the variation of scatter between square and rectangular field sizes for the lumbar spine AP projection as a low tube voltage was employed. There was an 82% difference between the rectangular and square fields with areas of 300 cm2 (i.e. 10 cm 6 30 cm for the rectangular field). For the rectangular field, the scattered X-rays were more attenuated as the smaller width covers thicker parts of the patient. For the square field, the scattered X-rays were less attenuated as the field edge was closer to the patient obliquity. Figure 3a shows that the equivalent square field approximated the variation of scatter for

square fields well when the smallest rectangular field dimension was greater than or equal to 10 cm. For the smallest field dimension of 5 cm, the equivalent square field overestimated the scatter for the square field. Figure 3b shows that for the chest PA projection, the differences in the scatter between square and rectangular field sizes were less than for the lumbar spine AP view due to the lower attenuation of the lung and the higher tube voltage. For example, for field areas of 300 cm2, a difference of 11.4% was observed at a 45 ˚ scattering angle. The scatter for the rectangular fields was closer to the square field values than the equivalent square field values where the smallest rectangular field dimension was less than or equal to 10 cm. For both examinations, the variation of backward directed scatter with field area was similar for both square and rectangular fields. Therefore, the scatter value for the largest appropriate square field should be employed in X-ray room design to provide a conservative dose estimate. The scatter from large rectangular fields and their equivalent square field sizes tends to be less than the scatter from square fields.

Comparison of scatter values in the literature Table 9 shows a summary of the imaging parameters stated in the literature which were used to obtain the scatter values shown in Table 10. All the imaging system

Table 9. The parameters used for measuring scatter in the different reports in the literature [1–3, 5, 20] Image parameters Parameter type

Trout and Kelley [3]

Bomford and Burlin [2]

McVey and Weatherburn [1, 20]

Williams [5]

Tube voltages (kV) Filtration (mmAl) Generator type

50, 70, 100 0.5, 1.5, 2.5 Self-rectified and single phaseb Torso-shaped masonite

100 – –

50, 70, 100 3.5 Constant potential

MixD block

50, 70, 100a 3.4 Medium frequency, 3% ripple Solid Water block

72630620

30630622

30.5630.5610

RANDO phantom (Alderson A muscle) 50625621.5

20 cm620 cm 100

400 cm2 (circle) –

20 cm620 cm 100

22 cm617.5 cm 80

Phantom material Phantom dimensions (cm6cm6cm) Incident field size FSD (cm) a b

100 kV for calculation only. 50 kV and 70 kV were used with a self-rectified generator and 100 kV was used with a single phase generator.

138


Simulating scatter from patients Table 10. Comparison of published scatter values [1–3, 5, 20] and the scatter calculated using the voxel Monte Carlo code (MC) in this paper for tissue equivalent materials Percentage scatter values for 50 kV X-rays Source of data

30 ˚

45 ˚

60 ˚

90 ˚

120 ˚

135 ˚

150 ˚

Trout and Kelley [3] Measured values [1] Calculated values [1] Williams [5] Voxel MC calculated values

0.044 – 0.081 0.071 0.059

0.011 – 0.059 0.070 –

0.012 – 0.048 0.073 0.068

0.024 – 0.041 0.120 0.100

0.069 – 0.114 0.221 0.162

0.095 0.181 0.140 0.279 –

– – 0.222 0.314 0.321

Percentage scatter values for 70 kV X-rays Source of data

30 ˚

45 ˚

60 ˚

90 ˚

120 ˚

135 ˚

150 ˚

Trout and Kelley [3] Measured values [1] Calculated values [1] Williams [5] Voxel MC calculated values Lumbar spine AP view values (72 kV)a

0.052 – 0.123 0.084 0.080 0.072

0.021 0.137 0.090 0.083 – 0.084

0.020 – 0.068 0.086 0.090 0.096

0.035 0.078 0.052 0.137 0.126 0.119

0.091 0.195 0.143 0.253 0.196 0.185

0.120 0.237 0.181 0.322 – 0.218

– 0.252 0.269 0.358 0.363 0.244

Percentage scatter values for 100 kV X-rays Source of data

30 ˚

45 ˚

60 ˚

90 ˚

120 ˚

135 ˚

150 ˚

Trout and Kelley [3] Bomford [2] Calculated values [20] Williams [5] Voxel MC calculated values

0.127 0.026 0.180 0.108 0.106

0.100 0.039 0.133 0.106 –

0.098 0.052 0.098 0.108 0.116

0.110 0.065 0.070 0.163 0.153

0.190 0.156 0.176 0.291 0.228

0.210 0.221 0.220 0.355 –

– 0.273 0.309 0.412 0.395

AP, anteroposterior. 20 cm 6 20 cm incident field size.

a

parameters were selected to be as similar as possible for this comparison. The range of tube voltages studied was restricted to be from 50 kV to 100 kV. All scattering materials used were tissue equivalent. The phantoms used by Trout and Kelley [3] and Williams [5] had a human body contour. The voxel Monte Carlo code was used to model the experimental set up for Williams scatter measurements (as detailed above). The scatter values in Table 10 were corrected where necessary to the same units of percentage scatter. Bomford and Burlin [2] measured the percentage scatter for incident air kerma on the surface of the phantom. Thus, a backscatter factor of 1.3 [22] was used to give the values in Table 10 as the ratio of scattered air kerma to incident air kerma without backscatter. Williams [5] reported scattered air kerma divided by the DAP. Williams’ values in Table 10 were converted to be in terms of scattered air kerma divided by the incident air kerma without backscatter. These values were also increased to be equivalent to a field area of 400 cm2, which was the same field area as the other studies, instead of 385 cm2 as shown in Table 9. Table 10 shows that there was good agreement between the calculated and measured scatter values reported by McVey and Weatherburn [1] and the calculated values given in the previous section and with the measured values given by Williams [5]. These agreements give confidence in both the calculations and measurements. Both studies [1, 5] employed modern equipment, including a DAP meter, in the experimental set up. A FSD of 100 cm was used in the work of McVey and Weatherburn and a FSD of 80 cm was used in the The British Journal of Radiology, February 2006

work of Williams (Table 9). Williams’ values were considerably greater than those of McVey and Weatherburn for scatter in the backward direction. The difference in FSD produced a difference in the position of the DAP meter which resulted in large differences in the scatter in the backward direction. McVey and Weatherburn’s values were greater than Williams’ values in the forward direction due to the smaller phantom thickness. Scatter in the backward direction was less affected by changes in phantom thickness than scatter in the forward direction [20]. The patient scatter values, calculated for the lumbar spine AP view at 72 kV, were similar to Williams’ values in the forward direction and less than Williams’ values in the backward direction (Table 10) as a DAP meter was not included in the scatter calculations. Table 10 shows that there was poor agreement in the forward directed scatter values reported by Bomford and Burlin [2] and Trout and Kelley [3]. For example at a 30 ˚ scattering angle, Bomford’s values were 0.2 times smaller than Trout’s. It is difficult to understand the reason for the differences between Trout’s and Bomford’s results as their phantoms had similar thicknesses (Table 9). Bomford and Burlin had corrected their values for scatter from the surroundings and leakage from the Xray tube head. These contributions were a large proportion of the total reading as the scatter from the phantom was small in the forward direction. McVey [20] calculated that the scatter from the collimators, ceiling, floor and walls varied between 0.03% and 0.05%. This accounted for some of the differences which were between 0.045% and 0.101%, but also suggested that 139

G McVey

the masonite and MixD phantoms may have substantially different attenuating properties. The scatter values were similar in the backward direction for the two phantoms as the scatter was less dependent on the phantom thickness or density.

Conclusions Accurate determination of the scatter in X-ray rooms is important for designing shielding to meet the desired radiation protection requirements. Previous studies have used a variety of phantoms to estimate these scatter levels and, as a review of the literature has shown, there are large differences in the published scatter values. The work in this paper has determined the magnitude of scatter from patients undergoing diagnostic X-ray procedures with the imaging system parameters varied in a systematic manner to provide a comprehensive data set. The voxel Monte Carlo calculations have demonstrated that the linear relationship between scatter and field area, as used in shielding reports [6, 21], is not valid for patient irradiation. For example, the scatter was 91% larger than the expected value for increasing the area of a square field from 100 cm2 to 625 cm2 for a patient undergoing an X-ray examination in the lumbar spine region. The position of the field on the patient in relation to the calculation points also had an effect. The scatter from a patient undergoing a lumbar spine LAT exam increased by 2.3 times for the centre of the field being moved from the centre of patient’s width closer to the patient’s obliquity with the calculation points on the same side of the patient. If scatter was calculated on both sides of the patient and the field centre moved laterally, then the scatter distribution would become asymmetric i.e. the scatter would be higher on one side compared with the other. Thus, the calculation points on the same side of the patient as the lateral shift would provide a conservative estimate of the scatter. As well as X-ray room design, this work can be applied to estimate the doses received by staff who undertake interventional procedures. The Monte Carlo calculations have also demonstrated small variations in patient scatter, in particular for changing the tube voltage ripple. For example, the scatter from a patient undergoing a lumbar spine AP exam increased by 22% if an X-ray generator with a voltage ripple of 50% was replaced by a constant potential X-ray generator, whereas the scatter increased by 97% if the tube voltage increased by 38 kV. Therefore, replacing X-ray generators with substantial voltage ripple by medium frequency units would not produce sufficiently more scatter to warrant a change to the X-ray room design (using the same dose constraint). A review of published scatter values [1, 5] has suggested that there was a FSD dependence on backward directed scatter when a DAP meter was present. However, Marshall and Faulkner [4] found no FSD dependence for air kerma measured adjacent to the couch for a 90 ˚ scattering angle, i.e. at a position forward of the DAP meter. Marshall and Faulkner imply that the FSD was simply increased, leading to an increase in field size incident on the phantom, which may explain the 140

constancy in their scatter measurements. McVey and Weatherburn [1] showed that there was a large variation in backward directed scatter from the DAP meter. Therefore, further work is necessary to investigate these effects, their influence on patient scatter and their possible impact on shielding barrier calculations. For X-ray room design, the largest scatter values provided by either McVey and Weatherburn [1] or Williams [5] are recommended to provide a conservative dose estimate at any FSD. The scatter from the patient detailed in this paper may also be considered. In this case, the significant scatter from the surroundings (e.g. the DAP meter and X-ray collimators) [1] should be taken into account. The inconsistent forward scatter values given by Trout and Kelley [3] and Bomford and Burlin [2] are not recommended for use even for X-ray units which have significant voltage ripple. In this case, the scatter data from this work, McVey and Weatherburn [1] or Williams [5] should be used and modified by the ratios detailed in this paper depending on the amount of voltage ripple. The work in this paper can also be used to study the effect of changing tube voltage, filtration, voltage ripple, field area and field position on the patient scatter. For example, in X-ray room design, these factors can be used to increase the recommended scatter values or independently measured scatter values to provide a conservative dose estimate as appropriate to the clinical situation.

Acknowledgments I would like to thank Dr David Dance, Prof. Gudrun Alm Carlsson and Dr Michael Sandborg for providing the voxel Monte Carlo code which was the basis of the scatter calculations. I also acknowledge the use of the computer facilities at the Physics Department, Royal Marsden Hospital, London, where all the Monte Carlo simulations described in this paper were undertaken.

References 1. McVey G, Weatherburn H. A study of scatter in diagnostic X-ray rooms. Br J Radiol 2004;77:28–38. 2. Bomford CK, Burlin TE. The angular distribution of radiation scattered from a phantom exposed to 100– 300 kVp X-rays. Br J Radiol 1963;36:426–39. 3. Trout ED, Kelley JP. Scattered radiation from a tissue equivalent phantom for X-rays from 50 to 300 kVp. Radiology 1972;104:161–9. 4. Marshall NW, Faulkner K. The dependence of scattered radiation dose to personnel on technique factors in diagnostic radiology. Br J Radiol 1992;65:44–9. 5. Williams JR. Scatter dose estimation based on dose area product and the specification of radiation barriers. Br J Radiol 1996;69:1032–7. 6. Sutton DG, Williams JR. Radiation shielding for diagnostic X-rays. Joint report of the BIR/IPEM working party, 2000. 7. Zubal G, Harrell CR. Voxel based Monte Carlo calculations of Nuclear Medicine images and applied variance reduction techniques. Image Vision Computing 1992;10:342–8. 8. Zubal G, Harrell CR, Smith EO, Rattner Z, Gindi G, Hoffer PB. Computerised three dimensional segmented human anatomy. Med Phys 1994;21:299–302. 9. Sandborg M, McVey G, Dance DR, Alm Carlsson G. Schemes for the optimization of chest radiography using a



10.

11.

12.

13.

14.

15.

computer model of the patient and X-ray imaging system. Med Phys 2001;28:2007–19. McVey G, Sandborg M, Dance DR, Alm Carlsson G. A study and optimization of lumbar spine X-ray imaging systems. Br J Radiol 2003;76:177–88. Dance DR, McVey G, Sandborg M, Persliden J, Alm Carlsson G. Calibration and validation of a voxel phantom for use in the Monte Carlo modeling and optimization of Xray imaging systems. Proc. SPIE Medical Imaging 1999;3659:548–59. Persliden J, Alm Carlsson G. Calculation of the small-angle distribution of scattered photons in diagnostic radiology using a Monte Carlo collision density estimator. Med Phys 1986;13:13–24. International Commission on Radiological Units and Measurements. Photon, electron, proton and neutron interaction data for body tissues. ICRU Report No. 46, Bethesda, MD: ICRU, 1992. Kramer R. Determination of conversion factors between body dose and relevant radiation quantities for external X- and c-radiation. GSF Bericht-S-556, Neuherberg: GSF, 1979. Sandborg M, McVey G, Dance DR, Persliden J, Alm Carlsson G. A voxel phantom based Monte Carlo computer program for optimisation of chest and lumbar spine X-ray imaging. Radiat Prot Dosim 2000;90:105–8.


16. Birch R, Marshall M. Computation of bremsstrahlung X-ray spectra and comparison with spectra measured with a Ge(Li) detector. Phys Med Biol 1979;24:505–17. 17. Almeń A, Tingberg A, Mattsson S, Besjakov J, Kheddache S, Lanhede B, et al. The influence of different technique factors on image quality of lumbar spine radiographs as evaluated by established CEC image criteria. Br J Radiol 2000;73:1192–9. 18. Lanhede B, Ba˚th M, Kheddache S, Sund P, Bjo¨rneld L, Widell M, et al. The influence of different technique factors on image quality of chest radiographs as evaluated by modified CEC image quality criteria. Br J Radiol 2002;75:38–49. 19. International Commission on Radiation Units and Measurements. Tissue substitutes in radiation dosimetry and measurement. ICRU Report 44, Bethesda, MD: ICRU, 1989. 20. McVey GH. Monte Carlo computing applied to X-ray room design. D.Phil. Thesis, University of Oxford, 2002. 21. British Standards Institution. Recommendations for data on shielding from ionizing radiation: part 2. shielding from X-radiation. British Standard 4094, Part 2, London: BSI, 1971. 22. British Institute of Radiology. Central axis depth dose data for use in radiotherapy: 1996. BJR Supplement 25, London: British Institute of Radiology, 1996.

141


Techniques for measurement of dose width product in panoramic dental radiography P DOYLE,

MSc,

C J MARTIN,

PhD

and J ROBERTSON,

PhD

Health Physics, Department of Clinical Physics and Bio-Engineering, Gartnavel Royal Hospital, Glasgow G12 OXH, UK ABSTRACT. Dose width product (DWP) is the quantity recommended for assessment of patient dose for panoramic dental radiography. It is the product of the absorbed dose in air in the X-ray beam integrated over an exposure cycle and the width of the beam, both measured at the receiving slit. A robust method for measuring the DWP is required in order to facilitate optimization of practices and enable comparison of dose levels at different centres. In this study, three techniques for measuring the DWP have been evaluated through comparison of results from 20 orthopantomographic units. These used a small in-beam semiconductor detector and X-ray film, a pencil ionization chamber and an array of thermoluminescent dosemeters (TLDs). The mean results obtained with the three techniques agreed within ¡6%. The technique employing a pencil ionization chamber of the type used for dose assessment of CT scanners is the simplest and most reliable method. The in-beam detector and film method has larger errors both from positioning the radiation detector and from measurement of X-ray beam width, which should be the full width at half maximum obtained from a scan of the film optical density. The TLD array method was accurate, but more time consuming to carry out. The mean DWP for the units studied was 65 mGy mm and the mean dose– area product was 89 mGy cm2. The DWP for 30% of the units tested exceeded the diagnostic reference dose of 65 mGy mm, recommended by the National Radiological Protection Board.

Panoramic radiography is a technique used in dentistry to show the mandibular joints with the teeth laid out between them. The X-ray tube and film holder both rotate during the exposure. The film is exposed to a narrow X-ray beam through a secondary collimator slit, across which the film moves as the radiographic image is built up. The assessment of patient dose in panoramic radiography is difficult because of the dynamic nature of the imaging process and the narrow width of the X-ray beam. The dose quantity used is the product of the absorbed dose in air and the horizontal width of the beam, both measured at the front side of the secondary collimator slit, and integrated over a standard exposure cycle. This is referred to as the dose width product (DWP) with units of mGy mm. The DWP provides a measurement related to the total amount of radiation to which the patient is exposed. It can be derived either from the product of the peak dose at the centre of the Xray beam and the width of the beam, or from an incremental summation of the dose across the beam. The aim of this study has been to compare and evaluate results obtained from different techniques available for measuring DWP. As part of the National Radiological Protection Board (NRPB) dental X-ray protection service, Napier reported DWPs for 387 panoramic dental X-ray sets derived from a technique that employed film to assess both dose and beam width [1]. Based on results from this survey, the Dental Guidance Notes recommend that a DWP of 65 mGy mm should be adopted as a local diagnostic 142

Received 7 March 2005 Revised 7 June 2005 Accepted 15 June 2005 DOI: 10.1259/bjr/33207232 ’ 2006 The British Institute of Radiology

reference level (DRL) for a standard adult panoramic radiograph [2]. Results from this study have been compared with this value.

Method The dose measurement techniques used in this study were: (a) ‘‘In-beam’’ detector and film: Measurement of peak dose within the X-ray beam at the receiving slit using a small solid state detector and determination of the beam width using X-ray film [3]. The DWP is calculated from the product of peak dose and beam width. (b) Partial volume detector: Direct measurement of the summation of dose across the beam obtained from the partial volume irradiation of a pencil ionization chamber [4, 5]. (c) Thermoluminescent dosemeter (TLD) array: Measurement of dose at the receiving slit using a linear array of TLDs. This method can evaluate the DWP either from the incremental summation of dose across the beam or from the peak dose multiplied by the beam width [3, 6]. More details of the techniques used and the measurements made in evaluating them are given in the following paragraphs. The British Journal of Radiology, February 2006

Dose measurement for paranormic dental radiography

‘‘In-beam’’ detector and film A solid state detector that has an active width of 1.5 mm, which is marketed for measurement of the DWP, was used with an Unfors 511 Mult-O-Meter (Unfors Instruments AB, Billdal, Sweden). Measurement showed that the length of the sensitive region was approximately 4 mm. The detector was attached to the front side of the secondary collimator parallel to the slit and aligned visually with the slit. It is important that the detector is aligned accurately with the X-ray beam and is sufficiently narrow to enable it to lie entirely within the region of the dose peak in order to give an accurate result. An assessment of the spatial response across the solid state detector was made using an X-ray beam from a radiographic unit collimated by a 0.2 mm wide lead slit. The detector was moved perpendicular to the slit in 0.2 mm steps by means of a micromanipulator with a vernier scale. Images of the X-ray beam at the receiving slit of each orthopantomographic (OPT) X-ray unit were obtained by exposing Kodak T-mat L film. When this technique is employed, care is needed to avoid saturating the film. The width of the beam was obtained by measuring the film optical density with a microdensitometer (MKIII CS; Joyce-Loebl Ltd, Gateshead, UK) and determining the full width at half maximum (FWHM) (technique A1). However, a simple measurement using a ruler with a light box (technique A2) has been recommended [3] and this was also used in order to determine whether the errors involved were significant.

Partial volume detector A pencil ionization chamber commonly used for CT dose index measurements (model No. 20X5-CT with a MDH 2025 electrometer; Radcal Corporation, Monrovia, NY) was attached in front of the secondary collimator, perpendicular to the slit. The DWP was taken as the product of the partial volume irradiation reading and the active length of the chamber (100 mm) (technique B).

TLD array This technique involved measuring the dose profile at the receiving slit using an array of 34 TLDs mounted in a Perspex jig with 1 mm thick walls and lid. The TLDs used were high sensitivity LiF:Mg:Cu:P TLD-100H chips (0.38 mm thick and 3.2 mm diameter), calibrated in a 70 kVp X-ray beam against air kerma in air, measured using a 6 cm3 chamber and a Radcal 9010 electrometer. The TLDs were placed on their edge, side by side in the jig, which was then positioned in the centre of the secondary collimator perpendicular to the slit. The TLDs were read out using a Harshaw 5500 TLD reader (Qados, Sandhurst, UK). The dose that each TLD received was obtained by correcting the readout for background radiation and applying a batch calibration factor. The spacing of the TLDs in the jig was determined from a measurement of the length of the arrays and found to be 0.40 mm. The DWP was calculated from the sum of the The British Journal of Radiology, February 2006

doses received by all the TLDs d1–d34, multiplied by the spacing w (technique C1) i.e.: DWP 5 w(d1 + d2 +...+ d34) The doses recorded were plotted against position in the jig to give a profile of the dose distribution across the slit (Figure 1). The DWP was also calculated from the product of the maximum dose at the centre of the beam and the FWHM value (technique C2). Comparison of techniques C1 and C2 was used to confirm that the dose summation and the product of peak dose and FWHM gave similar results for the DWP.

Study method Detectors and TLDs were all calibrated with respect to an ionization chamber with a Keithley Triad 35050A dosemeter system, which had a calibration traceable to a national standard. Relative responses were measured with the detector free in air and lying on a steel plate with a slit overlying a cassette to simulate actual exposure conditions. Based on these measurements, results obtained with the CT chamber and the TLDs were reduced by 5% to allow for the effect of backscatter. The Unfors detector is shielded from backscatter to a greater extent because of the metal its plate to its rear. Measurements of DWP using the three techniques were made on 20 different OPT X-ray units from eight manufacturers (Table 1). The OPT units had been installed at various times over the previous 25 years and had an average age of 10 years. Successive measurements were made using each of the three techniques at the standard adult settings, typically; tube potential 70 kV, tube current 10 mA and exposure time 16 s, and values for the DWP derived for each technique. Experimental errors for the different techniques were estimated to be A1 ¡16%, A2 ¡19%, B ¡7% and C1 ¡8% and C2 ¡8%. Errors are expressed as percentages for each result, combining errors from individual components. The largest contributions were from the measurement of beam width using a ruler and the positioning of the Unfors detector. The dose–area products (DAPs) for each unit were calculated from the product of the DWP and the beam length L [6]. The mean value for the DWP derived from measurement techniques A1, B and C2 was employed. The beam length L was measured with a ruler on a light box using the film exposed for technique A.

Results The OPT units studied had a range of beam widths and examples of dose profiles obtained using TLDs for OPT units with average beam widths of 2.5 mm and 4 mm are shown in Figure 1. DWP results obtained using the different techniques are shown in Table 1. The two sets of values for the DWP obtained from the TLD data using different calculation methods, i.e. from dose summation and from the product of the peak and the FWHM, are compared in Figure 2. The ratio of the DWPs measured using the two techniques is 0.96¡0.02 (mean ¡ standard error of the mean (sem)), confirming that the 143

P Doyle, C J Martin and J Robertson

Figure 1. Dose profiles from orthopantomographic (OPT) unit with beam widths of 2.5 mm and 4 mm, measured using thermoluminescent dosemeters (TLDs), compared with the measured sensitivity profile across the width of the Unfors detector.

two methods give results which agree within experimental error. There is also reasonable agreement between the DWP results obtained using the TLDs, technique C1 and those from the pencil ionization chamber, technique B (0.91¡0.014, mean ratio ¡ sem) and the in-beam and detector method using the FWHM, technique A1 (0.85¡0.034, mean ratio ¡ sem), see Figure 3. The average DWP given by the different techniques are A1 61 mGy mm, A2 80 mGy mm, B 65 mGy mm, C1 72 mGy mm and C2 69 mGy mm. The standard deviation between the techniques A1, B and C2 averaged for the 20 units was 13%. The Unfors detector should provide a reasonably accurate measurement of the peak dose, if it is positioned at the centre of the X-ray beam. The measured sensitivity response across the width of the Unfors detector is compared with two X-ray beam profiles in Figure 1. Misalignment of the detector and the centre of the beam

by distances of 0.5 mm, 1 mm and 2 mm would give measurements lower by 2%, 16% and 59%, respectively, for a 4 mm beam width, and by 5%, 27% and 73%, respectively, for a 2.5 mm beam width. Results from technique A were more scattered than those from techniques B and C. Any misalignment between the detector and the beam would give a lower value for the peak dose and results for the DWP from technique A1 were slightly lower than for the other techniques. The FWHM measured from films using the microdensitometer on five of the OPT units selected randomly, agreed to within 3% with the FWHM derived from TLD measurements. DWP measurements calculated using the data measured by the Unfors detector multiplied by the TLD profile FWHM, rather than the film FWHM, are included in Table 1 (technique A1). This was because a few of the films were saturated at the centre of the X-ray beam and so could not be used. The average beam

Table 1. Dose width product (DWP) measurements taken on different panoramic X-ray models at the standard adult setting Code Model/commissioned

I II III IV V VI VII VIII IX X XI XII XIII XIV XV XVI XVII XVIII XIX XX

Sirona Orthophos (2002) Instrumentarium OPT (2002) Siemens Orthophos (1993) Morita Inc. Panex EC (1982) Siemens Orthophos (1999) Instrumentarium OPT (1998) Siemens Orthophos (1989) Orion Cranex DC2 (1982) Planmeca PM2002CC (1990) Siemens Orthophos (1990) Planmeca Proline (2002) Planmeca PM2002CC (1990) Siemens Palomex (1987) Planmeca Proline (2002) Morita Inc. Panex EC (1979) Planmeca Proline (2002) Siemens Palomex (1976) Soredex Cranex (1993) Yoshida Panoura (1990) Morita Inc. Panex EC (1980)

kV

68 73 69 70 74 66 66 69 68 60 68 68 60 68 70 68 65 81 85 90

mA

8 8 15 7 14 10 15 6 6 16 7 6 14 7 8 7 15 10 10 10

Time (s)

14 18 14 14 13 18 15 19 18 14 18 18 15 18 16 18 14 16 12 16

Nominal Film/screen speed

DR DR Kodak T-mat L 400 Kodak T-mat L 400 Kodak T-mat L 250 Kodak Ekta 400 Kodak T-mat L 400 Kodak T-mat L 400 Kodak T-mat L 400 Kodak T-mat L 400 Kodak T-mat L 400 Kodak T-mat L 400 Kodak T-mat L 400 Kodak T-mat L 400 Agfa Curix 250 CR Kodak T-mat L 250 Agfa HTG Ortho 250 Ceahiplus 200 Kodak T-mat L 250

Beam width DWP (mGy mm): (mm) Technique Technique A1 B

Technique C2

3.5 2.5 3.0 4.5 4.1 2.8 2.3 5.5 3.7 3.2 2.4 3.7 3.0 3.0 4.5 3.8 6.3 3.6 7.6 6.9

45 48 48 41 65 53 61 49 62 63 63 63 63 66 75 79 76 93 122 152

43 32 40 58 44 54 45 65 59 54 43 62 53 55 60 58 81 65 114 128

38 51 53 43 52 56 57 54 51 55 66 52 62 61 66 64 83 93 119 128

DR, digital radiography; CR, computed radiography.

144



Figure 2. Plot of dose width products (DWPs) derived from thermoluminescent dosemeters (TLDs) showing the DWP derived from the product of the peak dose and full width half maximum (FWHM) against the DWP from the summation of the doses for all the TLDs across the beam. The line of identity is a 45 ˚ trendline.

FWHM of the OPT units included in this study is 4.0¡0.3 mm. The average beam width measured from the film with a ruler and a light box was 20% higher than the FWHM, and the DWP results calculated using this (technique A2) were higher than those obtained using the other techniques (Table 1, Figure 3). The overestimation of the beam width was partially offset by the lower dose resulting from misalignment of the detector. Values of DAP were calculated for each unit from the product of the average DWPs derived from techniques A1, B and C1 and measurements of the slit length L (Figure 4). The average beam length L was 136¡2 mm and the average DAP was calculated to be 89¡8 mGy cm2. Values of the mean and third quartile DWP and DAP are compared with results from other studies in Table 2. The average DWPs were greater than the proposed DRL of 65 mGy mm [2] for 30% of the units. 400 speed index systems are recommended by the European Guidelines on Quality Criteria for diagnostic radiographic images [7]. The five units with the highest DWPs and DAPs all used films with speed indices of 200–250, while the two lowest both used direct digital radiography (DR), (Table 1, Figure 4).

Discussion All three techniques gave results within reasonable agreement, but the errors associated with the in-beam detector and film technique A are larger than those for techniques B and C. For technique A, a microdensitometer was required for the measurements and it was important that exposures were limited to avoid saturation of the film emulsion in order that accurate results could be obtained. Measurement of the beam width with a ruler gave a result 20% greater than the FWHM and this method is therefore not appropriate. Use of a 35 mm film scanner (PrimeFilm 1800u; Pacific Image Electronics, Torrance, CA) linked to a PC with appropriate software (e.g. Scion Image; Meyer Instruments Inc., Houston, TX) and a calibrated film test strip to allow optical densities to be determined provides an inexpensive method for film scanning if a microdensitometer is not available, although this requires further limitation to be placed on the exposure because the measurable range in optical density is more limited. Another potential source of error in technique A is the visual positioning of the detector. The active area of the detector is 1.5 mm in diameter, which is similar to the width of the dose peak (Figure 1) with seven of the units studied having beam widths of

Figure 3. Plot of dose width product (DWP) measurements using an Unfors detector (techniques A1 – with full width half maximum (FWHM) derived from thermoluminescent dosemeter (TLD) profile and A2 – FWHM measured with a ruler and film) and a pencil ionization chamber (technique B), against the DWP derived from summation of doses across the beam from TLDs (technique C1). The line of identity is a 45 ˚ trendline.


145

P Doyle, C J Martin and J Robertson

Figure 4. Bar chart showing dose–area product (DAP) values for a standard adult exposure for the 20 units studied. Data from film/screen combinations with indices of 200–250 (F250, dark) and 400 (F400, light), and from computed radiography (horizontal lines) and digital radiography (angled lines) systems are indicated by different shading.

113

139

for the summation of the dose across the beam were similar to the product of the peak dose and FWHM. The agreement is closer than that reported in a previous study [6], probably because the earlier study used TLDs that were 0.85 mm thick. As a result, a limited number would lie within the X-ray peak, and this is likely to affect the accuracy of measurements of both the peak dose and the FWHM. The average DWP from the three techniques assessed in this study is similar to the reference dose recommended by the NRPB [1] and to the mean values reported in other studies by medical physics departments [6, 8] (Table 2). In the present study, the third quartile was not significantly different from the proposed DRL [2] because there were a significant number of units with similar DAPs (Figure 4). The third quartile values in other studies tended to be higher than the DRL. This could reflect differences resulting from the sample size or distribution, measurement technique or poorer optimization. It will also be influenced by the 5% correction applied to account for backscatter in techniques B and C in the present study. DAP measurements were slightly less than results from other studies (Table 2) [5, 6, 9]. Six of the units tested in this study had a DWP greater than the DRL of 65 mGy mm. Five of these were using a film/screen combination with a nominal speed index of 200 or 250, so adoption of a 400 speed system, which could potentially reduce these doses by 40–50%, has been recommended with a proportionate reduction in exposure levels. Two of the units had DWPs that were close to the recommended suspension level of 150 mGy mm [10] and these units also had beam widths of 7–8 mm which were significantly broader than the maximum recommended value of 5 mm [2]. Investigation of the operation of the units has been recommended in order to optimize the system set ups and so reduce the exposures.

101

117

Conclusion

85

97

3 mm or less (Table 1). The detector’s active area must be positioned within 0.5 mm of the centre of the beam to keep errors to within ¡5%. A dedicated holder incorporating a phosphor screen to facilitate alignment of the detector with the X-ray beam is available from the supplier of the detector, although it was not used in this study. The partial volume chamber method (B) is the most direct and simplest of the three techniques. Errors in the technique result from the calibration of the ionization chamber and the magnitude of the backscatter. When using this and technique A, care must be taken to ensure that the length of cable attached to the detector is sufficient to account for the rotational movement involved in the scan. Method (C) using the TLD array is the most time consuming of the three techniques because of the handling and processing of the high sensitivity TLDs, which are brittle and need to be handled with care. The technique is accurate and so provides a useful method for dose comparisons, but is not recommended for routine use. It was useful for confirmation that values Table 2. Comparison of results of this study with published data Sample size

This study 20 Napier [1] 387 Isoadri and 5 Ropolo [4] Perisinakis and 6 Damilakis [5] Williams and 16 Montgomery [6] Oduko [8] 26 Tierris et al 62 [9] (male) (female) 62

DWP (mGy mm):

DAP (mGy cm2):

Mean

3rd Quartile

Mean

65 57 74

67 67 84

90

113 65

76

69

80

DWP, dose width product.

146

89

3rd Quartile

This study has measured the DWP using three different techniques. The method using a semiconductor detector and film required the slit width to be assessed The British Journal of Radiology, February 2006


from the FWHM of the exposure peak, measured using a microdensitometer, as use of a simple ruler measurement [3] gave a result 20% greater than the true one. Uncertainty in alignment of the detector with the X-ray beam of more than 0.5 mm could result in a significant error. If this technique is employed, a microdensitometer and a dedicated alignment tool are recommended. Use of a partial volume ionization chamber technique (B) provides a simple, robust method for direct measurement of the DWP, and is recommended as the technique of choice. The measurements are simple to record, avoid errors from positioning the radiation detector and do not require a measurement of beam width. The pencil type ionization chamber is also widely available and commonly used in diagnostic radiology departments for the measurement of CT dose index.

Acknowledgments The authors would like to thank Louise Lindsay and Navneet Dulai for their assistance with the DWP measurements.

References 1. Napier D. Reference doses for dental radiography. Br Dental J 1999;186:8.


2. NRPB. Guidance notes for dental practitioners on the safe use of X-ray equipment. NRPB, Department of Health, Chilton, UK: NRPB, 2001. 3. British Institute of Radiology. Assurance of quality in the diagnostic imaging department (2nd edn). London, UK: BIR, 2001:51. 4. Isoardi P, Ropolo R. Measurement of dose–width product in panoramic dental radiology. Br J Radiol 2003;76: 129–31. 5. Perisinakis K, Damilakis J, Neratzoulakis J, Gourtsoiannis N. Determination of dose–area product from panoramic radiography using a pencil ionization chamber: normalized data for the estimation of patient effective and organ doses. Med Phys 2004;31:4. 6. Williams JR, Montgomery A. Measurements of dose in panoramic dental radiology. Br J Radiol 2000;73:1002–6. 7. Commission of the European Communities. European Guidelines on Quality Criteria for diagnostic radiographic images. EUR 16260 EN. Brussels, Belgium: CEC, 1996. 8. Oduko J. Optimisation of patient dose and image quality in dental radiology– Over 65 time to retire your OPG? IPEM Meeting, York, March 2001. 9. Tierris CE, Yakoumakis EN, Bramis GN, Georgiou E. Dose area product reference levels in dental panoramic radiology. Radiat Protection Dosim 2004;111:283–7. 10. Institute of Physics and Engineering in Medicine, College of Radiographers, NRPB. Recommended Standards for the Routine Performance Testing of Diagnostic X-ray Imaging Systems, IPEM Report No. 77. IPEM: York, 1997.

147


A comparison of three-field and four-field techniques in different clinical target volumes in prostate cancer irradiation using dose volume histograms: a prospective three-dimensional analysis A HILLE,

MD,

¨ WS and C F HESS, N TO

PhD, MD

Department of Radiotherapy, University of Go¨ttingen, Go¨ttingen, Germany ABSTRACT. The purpose of the current study was to quantitatively assess differences between irradiation techniques on normal tissue exposure in different clinical target volumes (CTVs) in irradiation of prostate cancer. 14 patients with prostate cancer undergoing external beam radiotherapy were investigated. The prostate and prostate + proximal/entire seminal vesicles were delineated as CTVs. A three-field and two different four-field plans were generated and compared concerning rectum, bladder and femoral head dose–volume histograms (DVHs). The exposure of the rectum exposed to 40–60 Gy was significantly lower for all CTVs with the three-field technique compared with both four-field techniques. The exposure of the rectum to 70 Gy was significantly lower for all CTVs with the weighted four-field technique compared with the unweighted four-field and three-field techniques. The weighted four-field technique was worst in bladder dose sparing for the three CTVs. Comparing the threefield and the unweighted four-field technique for irradiation of the prostate and prostate + entire seminal vesicles, no technique provided a clear advantage or disadvantage in bladder dose sparing. For irradiation of the prostate + proximal seminal vesicles the unweighted four-field technique provided the best bladder dose sparing. Concerning the exposure of the femoral heads, the three-field technique was significantly worse for the three CTVs compared with both four-field techniques. No difference was found between the unweighted and the weighted four-field techniques. In conclusion, none of the studied techniques consistently proved superior in different CTVs in prostate cancer irradiation with respect to sparing all organs at risk. The absolute differences between the three techniques were small and the clinical relevance of these findings is uncertain.

Three-dimensional (3D) conformal radiation treatment with the use of individual multileaf collimators (MLCs) has become the standard treatment technique for localized prostate cancer [1–5]. The number of beams and their orientation vary from one department to another. The simplest techniques use three or four fields [2, 4, 6–10], others use techniques with over five fields [11, 12]. However, the published data do not indicate that more sophisticated techniques increase the therapeutic index [13–19]. It is known that rectal toxicity following external beam irradiation of prostate cancer correlates with radiation dose and the percentage of rectal volume included in the intermediate and high dose-volumes [1, 4, 10, 12, 20]. Recently, the impact of inclusion of the seminal vesicles in the clinical target volume (CTV) on rectal dose has been recognized and a risk-adapted CTV with exclusion of seminal vesicles or inclusion of the proximal 2–2.5 cm of the seminal vesicles was suggested to reduce the risk of rectal toxicity [21– 24]. Few studies compared different techniques concerning irradiation of the prostate only [3, 14, 18], the prostate + Address correspondence to: Dr Andrea Hille, Klinik fu¨r Strahlentherapie, Robert-Koch-Str. 40, 37075 Go¨ttingen, Germany.

148

Received 11 April 2005 Revised 10 June 2005 Accepted 21 June 2005 DOI: 10.1259/bjr/10206556 ’ 2006 The British Institute of Radiology

base of the seminal vesicles [19], or the prostate + entire seminal vesicles [14–16]. These studies draw differing conclusions concerning the best irradiation technique, which may partly be due to different definitions of the CTV in these studies. None of these studies investigated systematically whether there is a difference between techniques concerning irradiation of different CTVs of the prostate. The purpose of the current study was to quantitatively assess the differences between a simple three-field and two different four-field techniques on irradiated normal tissue exposure in irradiation of the prostate only, the prostate + proximal and the entire seminal vesicles. The evaluation was based on three-dimensional treatment planning including dose–volume histograms (DVH). To our knowledge, this is the first prospective systematic analysis for the effect of treatment technique on normal tissue exposure concerning three different CTVs in prostate cancer irradiation.

Methods and materials 14 consecutive patients with localized prostate cancer stage T1–2 undergoing external beam radiotherapy with The British Journal of Radiology, February 2006

Treatment techniques in prostate cancer irradiation

curative intent to 72 Gy were investigated prospectively. 3D conformal computer-based planning was carried out on CadPlan treatment planning system (Varian, Palo Alto, CA). The prostate (P), the prostate + entire seminal vesicles (PESV), or the prostate + proximal (PPSV) 2– 2.5 cm (approximately 60% in longitudinal direction) of the seminal vesicles were taken as CTV and a planning target volume (PTV) margin of 1 cm was added. The definition of the proximal seminal vesicles was taken from the literature [24]. The prostate, the entire and the proximal seminal vesicles were delineated on each axial slice on the planning computer. The external wall of the rectum was contoured. The craniocaudal rectal extension was defined as the first CT slice above the anal verge (caudal border) and the cranial limit was defined as the first slice below the sigmoid flexure. This definition is consistent with definitions reported in the literature [7, 8, 15, 25]. The external wall of the bladder was contoured. One planning CT scan (5 mm continuing, 5 mm slice) was carried out with patients in supine position and a comfortably filled bladder. Irradiation technique included individual optimization with conformal treatment planning and the use of individual blocks. Nine plans were produced for each of the 14 patients. Three different irradiation techniques using 20 MV photons were evaluated. (1) Four-field box technique with equally weighted fields (so-called unweighted four-field technique for simplification); (2) Four-field box technique with unequal weighted fields (so-called weighted four-field technique for simplification); (3) Three-field technique with one anterior and two lateral fields with 90 ˚ and 270 ˚ wedges. For technique 2 the weight of the ventral field versus dorsal field was 1.3:0.7, and 1:1 for the lateral fields; for 3 it was 1.3:0.85:0.85 with the highest weight for the anterior field. For techniques 2 and 3, minor modification of the beam weights were performed in order to homogenise the dose distribution inside the PTV. Dose was specified according to the ICRU 50 report [26]. For all techniques the reference point for dose specification was the same. Dose was specified at the centre of the treatment field in projection of the central axes. Concerning dose homogeneity, at least 95% of the PTV was covered by 95% of the prescribed dose as minimum. Field size was adjusted to reach this dose homogeneity criterion. Dose calculation included tissue density correction. To determine the amount of the rectum exposed to ionizing radiation, the percentage of the irradiated rectum to 40 Gy, 50 Gy, 60 Gy and 70 Gy were calculated by the treatment planning system. Several investigations indicate a relationship between DVHs and the development of chronic rectal toxicity [1, 6, 7, 9, 10, 20, 27–31]. The rectal contouring varies from study to study with some investigators outlining the whole rectum, others the rectal wall. Concerning the rectal borders, some studies outline the anatomic rectum, others the rectum over the length of the fields. It is known that there is a high variability of volume fractions of rectal DVHs depending on how the rectal borders are defined, and it is difficult to compare the results of different The British Journal of Radiology, February 2006

studies concerning rectal DVHs [32, 33]. Studies analysing either dose–volume relationships of the rectum with or without the craniocaudal definition which we have used in our study, or studies using an identical or similar craniocaudal definition of the whole rectum, are summarized in Table 1. The above mentioned values were chosen following the rectal dose constraints given in these publications [6–8, 27–31]. To determine the amount of the bladder exposed to ionizing radiation, the volume of V100 (defined as the percentage of bladder volume receiving 100% of the prescribed dose) and the percentage of the irradiated bladder to 40 Gy, 50 Gy, 60 Gy, 65 Gy and 70 Gy were calculated by the treatment planning system. The data on the tolerance of the bladder to radiation as a function of the irradiated volume is limited. This may be due to a large variation in the bladder DVHs when considering the modifications of the organ due to different filling [9]. The above mentioned values were chosen following the data in the literature about a relationship between bladder toxicity and the irradiated bladder volume [4, 34–37]. The incidence of acute bladder toxicity increased when more than 30% of the bladder received more than 65 Gy [4, 34, 35]. Late complications, such as bladder contracture and volume loss, are described in 5% to 10% at doses of 40 Gy delivered to the majority of the bladder, at doses of 50–65 Gy delivered to about 30% of the bladder volume and at doses of 65–75 Gy applied to below 20% of the bladder volume [36]. Emami et al estimated the TD5/5 to 65 Gy irradiated to the whole bladder, and 66% of the bladder volume irradiated to 80 Gy [37]. To determine the amount of the femoral heads exposed to ionizing radiation the volume of V50 and V100 (defined as the percentage of femoral head volumes receiving at least 50% and 100% of the prescribed dose) were calculated by the treatment planning system. The available data on the dose–effect relationship for femoral heads are also limited [37, 38]. The clearest proposal is that of Emami et al who indicated that a dose of 52 Gy can be given to the whole femoral head with a risk for chronic toxicity in 5 years to be 5%. The V100 value was chosen following Emami et al’s study. However, the dose to the whole femoral head is always lower than 52 Gy in practice. Therefore, we additionally estimated the V50 value.

Statistical analysis Analysis was performed using the program STATISTICA 6.1 (Stat Soft, Palo Alto, CA). To evaluate the statistical significance of differences, Friedman’s ANOVA was performed followed by Wilcoxon matched pairs test. Closure principle was used for multiple tests. The level chosen for significance was .0.05.

Results PTV The dose distributions in the PTV for all three CTVs were between 99.8% and 102% and typical standard 149

A Hille, N To¨ws and C F Hess Table 1. Relationship between dose–volume and late rectal toxicity reported in the literature Authors

Craniocaudal definition of the whole rectum

Rectal volume (%)

Rectal dose (Gy)

Risk for chronic rectal bleeding (%)

Toxicity grade

Used toxicity score

Median follow up

Cozzarini [6]

Anal verge to the sigmoid flexure (patients with large air/faecal content in the rectum were excluded from analysis) Anal verge to the sigmoid flexure (patients with large air/faecal content in the rectum were excluded from analysis)

¢ 52 , 63 ¢ 39 , 39 50 50 60 60 70 70 50 50 60 60 70 70 . 65

50 50 60 60 . 53 , 53 . 42 , 42 . 22 , 22 . 53 , 53 . 42 , 42 . 22 , 22 40

14 6.6 13.3 7 14.2 4.3 14 5.2 12.9 6.3 14.2 4 14 5 13 6 18

¢2

RTOG

3 years

¢2

RTOG

30 months

¢2

RTOG

2 years

¢2

RTOG

28 months

65 30 30 25 25 26 26

40 60 60 70 70 70 70

¢2

RTOG/ 6 years LENT-SOMA

¢ 57 , 57 . 42 , 42 . 25 , 25 . 30 , 30

60 60 . 60 , 60 70 70 ¢ 70 ¢ 70

Fiorino [7]

Fiorino [8]

Anal verge to the sigmoid flexure (patients with large air/faecal content in the rectum were excluded from analysis)

Greco [27]

Anal verge to the sigmoid flexure

Huang [28]

Wachter [29] Zapatero [30] Storey [31]

11 cm in length starting at 2 cm below the inferiormost aspect of the ischial tuberositas From the lower to the upper border of the fields Anus to the sigmoid flexure Rectum was identified with rectal contrast

, . , . , . ,

deviations ranged from 1% to 2.5% for all patients and all considered CTVs and techniques. No statistically significant differences were found between the three techniques for all three CTVs.

0 18 0 18 0 54 13 31 11 7.7 0 37 13 40

2 ¢2 ¢2 ¢2 3 3

RTOG/ EORTC RTOG

30 months 4 years

RTOG/ 2 years LENT-SOMA

Details are demonstrated in Table 2. The values of 40 Gy, 50 Gy, 60 Gy and 70 Gy for all three techniques are demonstrated graphically for the prostate only in Figure 1, for the PPSV in Figure 2 and for the PESV in Figure 3.

Rectum volume The median volume of the rectum was 84 cm3 (mean value 90 cm3, standard deviation 31 cm3).

Bladder volume The median bladder volume was 154 cm3 (mean value 146 cm3, standard deviation 40 cm3).

Exposure of the rectum with different techniques The exposure of the rectum to 40–60 Gy was significantly lower for all CTVs with the three-field technique compared with both four-field techniques. The exposure of the rectum to 70 Gy was significantly lower for all CTVs with the weighted four-field technique compared with the unweighted four-field and three-field techniques. The differences between the rectal volume receiving 40 Gy, 50 Gy, 60 Gy and 70 Gy, respectively, were significant for all three CTVs between the three-field technique and both four-field techniques, and between both four-field techniques. 150

Exposure of the bladder with different techniques P The bladder volume receiving 50 Gy, 60 Gy and 65 Gy, respectively, was significantly higher with the weighted four-field technique compared with both the unweighted four-field and the three-field techniques. The bladder volume receiving 40 Gy was significantly lower with the unweighted four-field technique compared with the weighted four-field technique. No significant difference was found in the bladder volume receiving 40 Gy, 50 Gy, 60 Gy, 65 Gy and 70 Gy, The British Journal of Radiology, February 2006

0.003

0.02

0.02

62 65 10 53 55 9 43 43 9 25 25 8

, 0.001

respectively, between the unweighted four-field and the three-field techniques. No significant difference was found in the bladder volume receiving 70 Gy and the proportion of the bladder volume receiving 100% (V100) of the prescribed dose between the three techniques.

PPSV The bladder volume receiving 40 Gy was significantly higher with the weighted four-field and three-field techniques compared with the unweighted four-field technique. The bladder volume receiving 50 Gy was significantly higher with the weighted four-field technique compared with the three-field technique. The bladder volume receiving 60 Gy, 65 Gy and 70 Gy, respectively, was significantly higher with the weighted four-field technique compared with both the unweighted four-field and the three-field techniques. Comparing the bladder volume receiving 60 Gy and 65 Gy, respectively, the three-field technique resulted in a significantly lower value for 60 Gy and in no significantly different value for 65 Gy compared with the unweighted four-field technique. Concerning the proportion of the bladder volume receiving 100% (V100) of the prescribed dose, no significant difference was found between the three techniques.

0.004

0.01

, 0.001

, 60 0.001 62 7 , 48 0.001 50 6 , 37 0.001 39 7 18 0.007 17 7 1, 2, 3 relates to the radiation techniques described in Methods and materials.

0.01

0.002

, 0.001

28 28 7 22 23 6 17 17 5 8 7 3 40 Gy Mean Median SD 50 Gy Mean Median SD 60 Gy Mean Median SD 70 Gy Mean Median SD

39 35 10 26 25 6 19 19 5 9 9 3

34 32 8 25 24 7 18 18 5 6 6 3

, 39 0.001 35 10 , 26 0.001 25 6 , 19 0.001 19 5 , 9 0.001 9 3

, 34 28 0.001 32 28 8 7 , 25 22 0.001 24 23 7 6 , 18 17 0.001 18 17 5 5 6 8 0.008 6 7 3 3

, 0.001

64 65 7 49 52 7 39 41 7 23 22 7

60 62 7 48 50 6 37 39 7 18 17 7

, 64 0.001 65 7 , 49 0.001 52 7 , 39 0.001 41 7 , 23 0.001 22 7

54 56 8 45 48 7 36 39 6 20 21 7

54 56 8 45 48 7 36 39 6 20 21 7

, 0.001

74 78 10 57 58 9 46 45 9 27 26 9

69 71 9 56 56 9 44 44 9 22 19 9

, 74 0.001 78 10 , 57 0.001 58 9 , 46 0.001 45 9 , 27 0.001 26 9

62 65 10 53 55 9 43 43 9 25 25 8

, 69 0.001 71 9 56 0.02 56 9 , 44 0.001 44 9 22 0.01 19 9

p-value 3 p-value 2 3 p-value 1 2 1 p-value 2 3 1

2

p-value 1

p-value 2

3

p-value

1

2

p-value 1

3

3

p-value

PESV PPSV P (%)

Table 2. Mean and median values, standard deviations (SD) and p-values for the whole rectum exposed to 40–70 Gy in different techniques and clinical target volumes (CTVs)



PESV The bladder volume receiving 40 Gy was significantly higher with the weighted four-field and the three-field techniques compared with the unweighted four-field technique. The bladder volume receiving 60 Gy was significantly higher with the weighted four-field compared with the three-field technique. The bladder volume receiving 50 Gy and 65 Gy was significantly higher with the weighted four-field technique compared with the unweighted four-field technique. The bladder volume receiving 70 Gy was significantly lower with the threefield technique compared with both the weighted fourfield and the unweighted four-field techniques. The bladder volume receiving 100% (V100) of the prescribed dose was significantly higher with the weighted four-field technique compared with both the unweighted four-field and the three-field techniques. All other values showed no significant differences between the techniques. Details for the bladder dose exposure for the three CTVs are given in Table 3.

Exposure of the femoral head with different techniques The femoral head DVHs were averaged over the left and right sides to give a single value. For all three CTVs the radiation dose to the femoral heads was below 50 Gy with all three techniques. The mean V50 and V100 values for all three CTVs were below 53% and 10% of the prescribed dose with the different techniques. Concerning the V50 values, the three-field technique was significantly worse for all CTVs compared with both four-field techniques. No differences were 151

A Hille, N To¨ws and C F Hess

Figure 1. Mean values, standard error and standard deviations (SD) for the rectum exposed to 40–70 Gy in irradiation of the prostate only with different treatment techniques.

Figure 2. Mean values, standard error and standard deviations (SD) for the rectum exposed to 40–70 Gy in irradiation of the prostate + proximal seminal vesicles with different treatment techniques.

found between the unweighted four-field and the weighted four-field techniques. Concerning the V100 values no significant difference was found for all CTVs between the different techniques.

Discussion Our study shows that none of the studied three-field and four-field techniques consistently proved superior in irradiation of the prostate, prostate + proximal seminal vesicles and prostate + entire seminal vesicles with respect 152

to sparing all organs at risk. Published data do not indicate that more sophisticated techniques increase the therapeutic index [13–19]. Techniques with more than five fields have a very high burden for daily routine treatment planning, and an optimal radiation technique should not only provide the best sparing for all organs at risk (rectum, bladder, femoral heads), but also be safely implemented without undue burden and reduce the risk of any error. For example, simple, and therefore safe, verification by portal imaging during radiation treatment is given with simple radiation fields. For these reasons, a study, investigating only simple radiation techniques was performed. The British Journal of Radiology, February 2006


Figure 3. Mean values, standard error and standard deviations (SD) for the volumes 50 Gy, 60 Gy and 65 Gy of the rectum in case of irradiation of the prostate + entire seminal vesicles with different treatment techniques.

P Bedford et al [14] concluded that for irradiation of the prostate only a four-field technique with two oblique anterior and lateral fields would be optimal for rectal sparing. Koswig et al [3] found that, for irradiation of the prostate, only, the best rectal sparing was with a six-field technique. Khoo et al [16] concluded for prostate irradiation, a three-field technique would bring the best rectal sparing with acceptable bladder and femoral head doses. They performed plans with 6 MV photons. We compared in our study a three-field technique with two different four-field techniques using 20 MV photons and the results of Khoo et al [16] concerning rectal dose sparing can be confirmed by our results. Another recently published study investigated three-field techniques versus four-field techniques in irradiation of the prostate only and found that the three-field technique using an anterior and two lateral (270 ˚ and 90 ˚) fields with 20 MV photons provides the best rectal protection [18]. This can also be confirmed by our results. The recent study mentioned found no difference between the techniques in bladder exposure and discussed that this may be due to the CTV (prostate only) [18]. Our data indicate that the three-field technique provides, for irradiation of the prostate only, the best rectal dose sparing with no significant differences in bladder dose sparing compared with the unweighted four-field technique. The weighted four-field technique was worst in bladder dose sparing. Concerning the femoral head doses, the three-field technique was worst.

PPSV Neal et al compared a three-field, four-field and sixfield technique for irradiation of the prostate + the base of seminal vesicles [19]. They found no significant differences considering the irradiated volume of the The British Journal of Radiology, February 2006

bladder and the rectum. However, they found a better sparing of the rectum with a weighted four-field technique and a better sparing of the bladder with the six-field technique. Our data indicate for irradiation of the prostate + proximal seminal vesicles the three-field technique to be optimum in rectal dose sparing. The weighted four-field technique was worst in bladder dose sparing and the unweighted four-field technique provided a better sparing of the rectum compared with the three-field technique. Concerning the femoral head doses, the three-field technique was worst.

PESV For irradiation of the prostate + seminal vesicles, several authors compared different techniques and all of these studies conclude that no single technique is superior when considering all organs at risk (rectum, bladder, femoral head). Fiorino et al compared various coplanar techniques for conformal irradiation of the prostate and seminal vesicles [15]. A three-field technique with an anteroposterior and two lateral 30 ˚ wedged fields gave the best sparing of the rectum. The bladder was best spared with a six-field technique. The mean dose of the bladder was significantly better against the three-field technique and the four-field technique. However, considering V95, no significant difference was found between the techniques. The unweighted four-field technique gave the worst sparing of the bladder for Fiorino et al. In our study, the weighted four-field technique gave the worst sparing of the bladder. Bedford et al [14] compared various four-field techniques with a three-field technique which had lateral oblique fields. They concluded that for irradiation of the prostate and prostate + seminal vesicles, four-field techniques with two oblique anterior and lateral fields with individual field wedges for the different CTVs to be 153

154 Table 3. Mean and median values, standard deviations (SD) and p-values for the radiation exposure to the bladder in different techniques and clinical target volumes (CTVs) (%)

PPSV

1

2

p-value 1

3

44 45 15 36 35 14 28 27 12 22 20 11 14 12 8 4 3 8

48 49 17 37 36 14 29 27 12 23 21 11 15 12 9 4 3 3

44 0.003 45 15 36 0.005 35 14 28 0.003 27 12 22 0.008 20 10 14 n.s. 12 8 4 n.s. 3 8

46 45 16 35 31 14 27 24 12 22 21 10 15 13 10 3 3 2

p-value 2

n.s.

n.s.

n.s.

n.s.

n.s.

n.s.

48 49 17 37 36 14 29 27 12 23 21 11 15 12 9 4 3 3

3

46 45 16 35 31 14 27 24 12 22 21 10 15 13 10 3 3 2

p-value

n.s.

0.002

0.005

0.01

n.s.

n.s.

PESV

1

2

p-value 1

3

p-value 2

3

57 57 15 45 46 14 36 36 12 30 30 11 22 22 9 12 5 15

69 68 16 46 47 14 37 36 13 32 31 11 24 25 9 12 6 15

57 0.001 57 15 45 n.s. 46 14 36 0.02 36 12 30 0.003 30 11 22 0.002 22 9 11 n.s. 3 15

65 62 17 44 44 14 36 36 13 29 29 11 20 20 9 10 5 15

69 68 16 46 n.s. 47 14 37 0.007 36 13 32 n.s. 31 11 24 n.s. 25 9 12 n.s. 6 15

65 62 17 44 44 14 36 36 13 29 29 11 20 20 9 10 5 15

0.02

p-value

n.s.

0.005

0.003

0.001

0.003

n.s.

1

2

p-value 1

3

p-value 2

3

66 68 16 46 47 14 39 40 12 33 33 11 25 26 9 24 24 19

80 81 15 48 49 14 39 41 12 33 34 10 26 27 10 27 27 21

, 66 0.001 68 16 46 0.003 47 14 39 n.s. 40 12 33 0.03 33 11 25 0.03 25 9 24 0.04 24 19

81 84 15 47 47 13 38 38 12 32 33 11 24 24 9 25 21 21

80 0.005 81 15 48 n.s. 49 14 39 n.s. 41 12 33 n.s. 34 10 26 0.04 27 10 27 n.s. 27 21

81 84 15 47 47 13 38 38 12 32 33 11 24 24 9 25 21 21

p-value

n.s. n.s.

0.01

n.s.

0.01

0.02

1, 2, 3 relates to the radiation techniques described in Methods and materials.



40 Gy Mean Median SD 50 Gy Mean Median SD 60 Gy Mean Median SD 65 Gy Mean Median SD 70 Gy Mean Median SD V 100 Mean Median SD

P


optimal for rectal sparing. However, such a technique has a very high burden for daily routine treatment planning. The simple three-field plan in this study with an anterior and two lateral fields using 6 MV photons showed a comparable level with the four-field technique in rectal dose sparing, in case of irradiation of the prostate + entire seminal vesicles. However, the dose to the superficial body and femoral heads was found to be very high [14]. Khoo [16] concluded for both the prostate and seminal vesicles irradiation a three-field technique would bring the best rectal dose sparing with acceptable bladder and femoral head doses. They performed plans with 6 MV photons also, but the superficial body dose was not mentioned. In our study, we compared a threefield technique with two different four-field techniques using 20 MV photons, and the results of Khoo et al [16] concerning rectal dose sparing can be confirmed by our results. The weighted four-field technique was worst in bladder dose sparing. Comparing the unweighted fourfield with the three-field technique in bladder dose sparing, no clear advantages or disadvantages were found. Concerning the femoral head doses, the threefield technique was worst. The studies investigating various techniques draw differing conclusions concerning the best irradiation technique. Some studies, comparing four and three-field techniques concluded the three-field technique to be best in rectal dose sparing [16, 18]. Others did not confirm these results [14, 19]. The reasons for these differing findings are unclear; PTV margins and PTV coverage which have both an impact on radiation exposure of the organs at risk were comparable among these studies and comparable with our study. The different CTVs in theses studies could have been responsible for the different findings, but our study shows for all three CTVs the best rectal dose sparing with the three-field technique. Concerning the bladder dose, the differing conclusions may be due to different bladder fillings which is known to have an impact on bladder DVHs [9]. In most of the studies, only few endpoints concerning dose volume histograms had been chosen and the investigated points were not associated with doses given in dose-constraint studies. Analysing many dose endpoints, as we have done in our study, can lead to unclear, or even conflicting results. This could be an explanation for the differing results compared with our study. Although the differences between the three techniques were small in our study they were significant, and we conclude from our data that a three-field technique provided the best rectal but the worst femoral dose sparing with inconsistent results regarding the bladder dose sparing for all three CTVs.

Estimated risk for chronic normal tissue exposure Rectum To associate the rectal DVHs in our study with an estimated risk for chronic rectal toxicity, the results were compared with studies analysing relationships between dose–volume and rectal toxicity (Table 1). Although the definitions of the rectum differ in some of these studies from our definition, and although the The British Journal of Radiology, February 2006

cut-off levels and the resulting risks for chronic rectal toxicity ¢ grade 2 also differ in these studies, we can draw cautious conclusions from our results regarding an estimated risk for chronic rectal toxicity. The obtained values from the three different techniques were close together. In patients treated with irradiation of the prostate only, the values for the rectum exposure were below 5% for chronic rectal bleeding ¢ grade 2; for irradiation of the prostate + proximal seminal vesicles between 5% and 15%; and for irradiation of the prostate + entire seminal vesicles over 15% with all three techniques. Nevertheless, the three-field technique provided the best rectal dose sparing, except for the rectal volume exposed to 70 Gy. The weighted four-field technique provided a significantly better rectal dose sparing than the unweighted four-field technique. Whether the small differences between the various techniques would have an impact on chronic rectal toxicity is uncertain. Furthermore, published data suggest increased local control with lower normal toxicity with new technologies such as intensity modulated radiation therapy (IMRT) [39–42]. IMRT allows the increase of dose in part of the prostate while continuing to protect normal tissue. However, until new technologies such as IMRT are introduced as a widespread clinical routine treatment, 3D conformal radiation therapy should be optimized to reduce toxicity while inreasing local control.

Bladder To estimate the risk for bladder toxicity we tried to compare our data with the clinical relationship between DVHs and the development of bladder toxicity, as reported in the literature [4, 34–37]. The exposed bladder volumes in our study were similar with the three different techniques. The risk of chronic bladder toxicity in our study can be estimated to be less than 5% to 10% in irradiation of the prostate only with all three techniques. In irradiation of the prostate + proximal/entire seminal vesicles, the risk for chronic toxicity can be estimated to be above 10% with all three techniques. For all three CTVs the weighted four-field technique provided the worst bladder dose sparing. Whether the small differences between the various techniques would have an impact on chronic bladder toxicity is uncertain. Femoral heads The three-field technique provided the worst radiation exposure to the femoral heads. The differences for the V50 value were significant; for the V100 value no significant difference was found between the techniques for all three CTVs. As the values for the three techniques were below 52 Gy to the whole femoral heads, the risk of chronic toxicity can be estimated to be below 5% in 5 years for the three CTVs [41].

Conclusion In conclusion, none of the studied techniques consistently proved superior in different CTVs in prostate cancer irradiation with respect to sparing all the organs 155


at risk. The absolute differences between the three techniques were small and the clinical relevance of these findings remains uncertain.

15.

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SHORT COMMUNICATION

A comparative evaluation of two head and neck immobilization devices using electronic portal imaging 1

K DONATO,

BSc (Hon),

1

K LESZCZYNSKI,

PhD, FCCPM

and 2K FLEMING,

MHSc, MRT(T)

1

Northeastern Ontario Regional Cancer Centre, Hoˆpital re´gional de Sudbury Regional Hospital, 41 Ramsey Lake Road, Sudbury, Ontario, P3E 5J1 and 2Grand River Regional Cancer Centre, 835 King Street West, Kitchener, Ontario, N2G 1G3, Canada ABSTRACT. A study was performed to compare the positioning reproducibility and the cost efficiency for two head and neck immobilization devices: the Uvex’ (Uvex Safety, Smithfield, USA) plastic mask system and the Finesse Frame with Ultraplast System’ (PLANET Medical, Svendborg, Denmark). 20 patients treated with 3D conformal radiation therapy for head and neck cancers were randomly selected (10 for each of the two different immobilization systems) and electronic portal images acquired during their course of treatment were saved and used in this study. The anatomical landmark coordinates and their shifts in the anteroposterior (AP) and craniocaudal (CC) directions with respect to the digitized simulator films for lateral fields were analysed using an inhouse developed portal image registration system. Statistically, no evidence was found to indicate that the systematic components of the displacement for the Uvex’ system and the Finesse Frame with Ultraplast System’ were different from each other or from zero. The random component of displacement was slightly smaller in the AP direction for the Uvex’ than the Ultraplast’ system (s51.9 mm and 2.9 mm, respectively, p50.007), but larger in the CC direction (s53.8 mm and 2.2 mm, respectively, p,1029). Production time and required materials for a radiation therapy department were also quantified to assess costs for each system. The overall costs per patient were estimated at $141.50 (CAD) and $82.10 for the Uvex’ and Ultraplast’ systems, respectively. The Finesse Frame with Ultraplast System’ of immobilization for head and neck cancer treatment provides a field placement reproducibility that is equal to, or greater than, that of the Uvex’ plastic mask immobilization system and, while it requires more expensive materials, the workload and consequently overall cost is greatly reduced.

Radiation treatment to the head and neck region is delivered with accurate and precise placement of prescribed portal fields. Reproducible alignment is increasingly important as we apply high-dose three-dimensional conformal radiation therapy (3D-CRT) techniques and intensity-modulated radiation therapy (IMRT) in conjunction with the need for smaller clinical target volume (CTV) margins. The consequences of field placement errors have been described in various publications; failure to treat the entire planning target volume (PTV) may be responsible for local failure, and irradiation outside of the PTV may result in normal tissue complications to important organs [1–4], such as the spinal cord or the eye, in the case of head and neck cancers. In order to increase the reproducibility of portal field placement, various immobilization devices are used to stabilize the position of the patient’s head while treatment is delivered. Previous publications have compared two or three different systems of immobilization [5–7], and have assessed treatment field position reproducibility with Address correspondence to: K Leszczynski.

158

Received 16 March 2005 Revised 27 June 2005 Accepted 8 August 2005 DOI: 10.1259/bjr/32191494 ’ 2006 The British Institute of Radiology

similar results: the standard deviations of field placement errors, s, were found to be between 1.7 mm and 3.3 mm, for both anteroposterior (AP) and craniocaudal (CC) directions. In this study, an immobilization system involving a Uvex’ (Uvex Safety, Smithfield, USA) plastic mask was compared with a low temperature thermoplastic mask system, the Finesse Frame with Ultraplast system’ (PLANET Medical, Svendborg, Denmark) using off-line electronic portal imaging. The costs of both systems in terms of production time and materials were also calculated since the clinical introduction of a low temperature thermoplastic mask appears to be less timeconsuming, less costly, and more convenient for the patient [7] and, therefore, beneficial in general.

Materials and methods Clinical setup This study retrospectively selected 20 consecutive head and neck patients treated at our centre between The British Journal of Radiology, February 2006

Short communication: An evolution of two head and neck immobilization devices Table 1. A summary of selected demographic data for patients included in both immobilization groups Demographics

Immobilization device group Uvex

Age: mean¡SD 65¡10 Sex: Females/Males2/8 Diagnosis Ca tongue – 4 distribution Ca tonsil –2 Ca hypopharynx – Ca pharynx – 1 Ca oral cavity – 1 Ca larynx –1 Weight change during treatment: Mean¡SD

4.8¡3.6

Ultraplast

60¡11 1/9 Ca supraglottis – 2 Ca floor of mouth – 2 1Ca tongue –1 Ca oropharynx – 1 Ca gingiva – 1 Ca glottis –1 Unknown primary – 2 3.4¡5.7

April 2001 and September 2001. These patients were randomly drawn from two groups representing different immobilization devices used in treatment setup; 10 patients with a Uvex’ mask system, and 10 patients with a FinesseFrame with Ultraplast System’. A summary of selected demographic data for both immobilization groups is provided in Table 1. Review of the summary confirms that there was no apparent significant demographical bias between the groups. One patient’s results were eliminated from the study due to poor quality imaging. All patients attended the Mevasim simulator (Siemens Medical Solutions, Concord, USA) where lateral portal positioning was marked on the mask for alignment purposes. A planning CT scan using Somatom Plus (Siemens Medical Solutions) was performed at 5 mm intervals for 3D conformal treatment planning on the Helax TMS system (Nucletron B.V., Veenendaal, The Netherlands), and all patients were treated on a Mevatron KD-2 or Primus linear accelerator (Siemens Medical Solutions) with 6 MV beams. The treatment beam arrangement consisted of two parallel opposed lateral fields covering the head and neck target. The nodes in the supraclavicular region were irradiated with an anterior field. Portal images were acquired and stored for daily fractions using Beamview Plus’ (Siemens Medical Solutions) video camera based electronic portal imaging. For the purpose of this study, only right lateral portal views were used. An average of 14 (range 9–19) portal images for each patient were acquired for a total retrospective analysis of 272 images.

Immobilization devices Two different thermoplastic masks and their respective immobilization accessories were evaluated in this study. To form the Uvex’ mask, two radiation therapists stabilize the patient’s head on a Timo head rest and form a plaster impression of the patient’s head and neck. One therapist later drapes the negative impression and fills it to form a positive impression. The 1/160 Uvex’ plastic is then heated in a vacuum former and then moulded around the plaster positive. The patient returns approximately 2 days later for the fitting process where the plastic mask is fitted directly onto the patient’s head and neck and The British Journal of Radiology, February 2006

mounted onto a Perspex’ (Lucite International Canada Inc., Mississauga, Canada) acrylic headboard at three fixation points on each side of the head (six fixation points in total, distributed evenly from the lower neck to the top of the head). The treatment field area is cut from the Uvex’ mask once a radiation oncologist approves the first-day portal image in order to allow for increased skin sparing. To form the Ultraplast’ mask, the Ultraplast’ material is dipped into a hot water bath (5–.30 (,26)

X-ray therapy

Various*

0–60 (26.5)

Intrauterine

0–.50 (11.4)

Multiple X-ray chest fluoroscopies Breast*, skin

0–57 (30)

Multiple X-ray chest fluoroscopies Lung, breast*

External high dose rate exposures Exposure to atomic bombings Life Span Study [53] Life Span Study [54, 55]

13–42 (24.4)

various*

Treatment of malignant disease Cervical cancer cohort [56] Leukaemia following cancer of the uterine corpus [57] Lung cancer following Hodgkin’s disease (international) [58] Childhood cancers (international) [59–61] Retinoblastoma [62]

Treatment of benign disease Childhood skin haemangioma: Stockholm [63–66] Childhood skin haemangioma: Gothenburg [67, 68] Ankylosing spondylitis [69, 70] Israel tinea capitis [71–74] New York tinea capitis [75] New York acute post-partum mastitis [76] Rochester thymic irradiation [77–79] Metropathia haemorrhagica [80]

287

Benign gynaecological disease [81, 82] Massachusetts TB fluoroscopy [79, 83] Canadian TB fluoroscopy [84, 85]

226

Ra

various*

(Continued )

Review article: What are the risks of low dose radiation?

The British Journal of Radiology, April 2006

Table 1. Epidemiological studies of the effects of exposures to external low-LET radiation

288

National Registry for Radiation Workers, UK [11]

*Sites for which statistically significant excesses were reported in the exposed group (cohort studies), or for which a higher proportion of the cases were exposed to radiation (case/ control studies). After UNSCEAR (2000).

Leukaemia, all other cancers Exposures mainly in nuclear industry Up to 47 (16.5)

Leukaemia, all other cancers Exposures in nuclear industry Up to 43 (22.2)

Maternal X-rays during pregnancy Leukaemia*, solid tumours 20 (max)

Cohort mortality 96673 workers Canada, UK, USA Cohort mortality 124743 monitored workers UK Nuclear workers [10]

Occupational exposure

Maternal X-rays during pregnancy Leukaemia*, all solid tumours* 16 (max)

Case-control 14491 cases 14491 controls UK Case-control 1342 cases 14292 controls USA Oxford Survey of Childhood Cancers [86, 87] NE USA childhood cancers [88]

Pre-natal exposure

Low-dose or low-dose-rate exposures

Study

Table 1 Epidemiological studies of the effects of exposures to external low-LET radiation (Cont.)

Type of exposure

Cancers studied Follow up, years (mean) Type of study

B F Wall, G M Kendall. A A Edwards et al

statistical power to provide direct evidence on radiation effects in humans for doses much below 10–50 mGy. This is because of the difficulty of observing a small number of additional cancers against very high background incidence rates; ever larger study populations would be required to detect ever smaller effects as doses decrease, and this is simply impractical. Moreover, epidemiology is an observational and not an experimental science. Epidemiologists make strenuous efforts to optimize study design, but it is not possible to select the exposed and reference populations on strict statistical grounds and there are likely to be residual effects of confounding factors and possible biases (e.g. selection bias), as well as other practical problems such as uncertainties in the dose estimates. UNSCEAR, in Annex I of its 2000 report [8], gives a discussion of the potential problems of bias and confounding, particularly in the low dose region where attempts are being made to resolve very small effects. The low dose region, where epidemiology is unable to produce clear evidence of risk, provides a fertile area for those who wish to argue that radiation risks have been overestimated or underestimated. The former can conclude, quite correctly, that there is ‘‘no significant evidence for an effect’’; this must, however, not be confused with there being ‘‘significant evidence for no effect’’. Furthermore, as discussed later, selection effects in epidemiological studies, in particular the ‘‘Healthy Worker Effect’’ must be allowed for. Those who believe that radiation risks are greater than the LNT extrapolation suggests can point to selected studies where the play of chance has resulted in apparently elevated risks at low doses, while ignoring studies which contradict this view. It is also true that publication bias will result in studies with significant findings reaching the literature more readily than those which are inconclusive. All this assumes that studies are well designed. Badly designed epidemiological studies, for example, with serious bias in the selection of the study populations, are available in the ‘‘grey’’ literature without peer review and are cited as evidence that radiation risks are underestimated or overestimated [14].

Radiobiological studies The national and international review bodies mentioned above also agreed that increased understanding of biological mechanisms will increasingly underpin judgements about the shape of the dose–response relationship in the low dose region. This increased understanding will come from qualitative and quantitative data from cellular and molecular studies of the biological mechanisms underlying the health effects of radiation. A comprehensive review of recent developments in this area has been undertaken by a Task Group of the International Commission on Radiological Protection [15]. Currently these studies indicate that the carcinogenic effects of radiation are caused largely by double strand breaks and complex lesions in stem cell DNA. Mechanistic modelling of radiation-induced carcinogenesis based on these radiobiological studies is still at an early stage of development. However, data on the role of gene mutations and DNA damage and repair The British Journal of Radiology, April 2006


mechanisms are now sufficiently well established to support the thesis that the risk of radiation-induced cancer at low doses rises as a simple function of dose without threshold for most types of cancer. While some experiments suggest a curve that is concave downwards and others one that is concave upwards, depending on the biological endpoint, there are sound biophysical arguments supporting the LNT model as the most appropriate general model for cancer induction. It should be appreciated that this may not be the most conservative approach and it might result in an underestimation of some radiation-induced cancer risks and an overestimation of others. Those involved in protection should also be alert to the fact that truly low dose experiments are difficult in cellular systems. Also, it remains the case that relatively few studies address directly the effects of low doses. Until the uncertainties are resolved, all the reviews concluded that the current weight of evidence on fundamental cellular processes supports the view that an increase in risk proportionate to the radiation dose is the most scientifically defensible approximation of the low-dose response.

comparison group. It is a very common finding in epidemiological studies of working populations that death rates are lower than in the general population. Consequently the Standardized Mortality Ratio (SMR) – the ratio (expressed as a percentage) of the number of deaths in the exposed group relative to the number expected based on rates for the general population – is often below 100. This is known as the Healthy Worker Effect [18, 19] and it is generally accepted that this is a reflection of the selection of the fit and healthy into employment and their retention in work. It would be a mistake to interpret low SMRs in such epidemiological studies as evidence for hormesis. For example, recent issues of an epidemiological journal described three studies of working populations. More examples could be cited, but these are sufficient to make the point. The cohorts were, respectively, workers exposed to ethylene oxide [20], workers in the petroleum industry [21] and workers exposed to formaldehyde in the garment industry [22]. The SMRs were:

N N

Could radiation stimulate beneficial adaptive responses and hormesis? It has been suggested [16] that low or moderate doses of radiation might stimulate responses, for example to DNA repair processes, which might counteract the harmful effects of the radiation damage. At its most extreme it has been suggested that these effects are so great as to confer a net benefit, at least in certain dose ranges. The general hypothesis of radiation stimulated beneficial changes is known as the Adaptive Response and the idea of net benefit is called Hormesis. The Adaptive Response was considered by the United Nations Scientific Committee on the Effects of Atomic Radiation (UNSCEAR) in its 1994 report [17] and by NCRP [9]. Explanations for an Adaptive Response tend to involve stimulation of DNA repair processes, although other mechanisms have also been suggested. However, the arguments put forward for adaptive responses and hormesis claim to have supporting evidence from epidemiological, ecological and radiobiological studies.

Epidemiological studies (the Healthy Worker Effect and unintentional bias) Various epidemiological studies have been conducted of populations exposed to low and/or protracted doses. In its 2000 report UNSCEAR [8] has undertaken a comprehensive review of these studies. Here we will focus on studies of people exposed to radiation either in the workplace or the home, because such studies have often been quoted to support claims of hormesis. The essence of the epidemiological method is to compare disease rates in two or more populations which differ qualitatively or quantitatively in their exposure to the agent under investigation. However, a common complicating factor is the existence of selection effects, typically factors which mean that those entering the exposed group tend to be more healthy than those in the The British Journal of Radiology, April 2006

N

Ethylene oxide exposed workers SMR 90 (95% CI 88– 93) Workers in the Petroleum Industry SMR 68 (95% CI 63–73) Garment workers exposed to formaldehyde SMR 92 (85% CI 88–96).

All three populations thus have significantly lower mortality than the general population. Are we to conclude that exposure to ethylene oxide, to hydrocarbons and to textile fibres or dust (in the presence of formaldehyde) are all beneficial to human health? The investigators who conducted these three studies did not draw such a conclusion, nor would epidemiologists generally. It is clear that what is being seen is the effect of selection factors which mean that working populations are healthier than the population as a whole, which includes the chronically sick and unemployed. It would be perverse to try to impose a different explanation in the case of exposures to radiation. One of the studies sometimes cited as demonstrating a beneficial effect of radiation exposure is that of mortality amongst British radiologists by Berrington et al [23]. Berrington et al observed significantly lower SMRs for radiologists employed after 1920, when occupational doses had fallen to moderately low levels, compared with other contemporary (unexposed) doctors. It has been suggested that this demonstrates a highly significant beneficial effect of radiation [24]. However, great care has to taken when attributing SMRs for different groups of people (even if they are all doctors) to a particular cause like radiation exposure, when no account can be taken of other possible confounding factors, like smoking. There is evidence in a study of mortality among doctors by Doll and Peto [25] that, during the 1950s and 1960s, GPs smoked more than other doctors and radiologists smoked less. This alone could be the cause of the lower SMRs for radiologists compared with other doctors (a large proportion of whom would be GPs), but without specific smoking habit information for the cohort of British radiologists and the control group of other doctors, it is impossible to be sure. It is because of these problems with external comparisons with other 289


populations that epidemiologists generally pay more regard to comparisons or trends in mortality (or disease incidence) within cohorts as evidence for radiation risks. Berrington et al focused on internal trends in mortality risks with time since entry into the radiology profession (as an indicator of cumulative dose), rather than relying solely on external comparisons of SMRs with other groups of doctors or the general population. Thus they found evidence for an increasing trend in risk of cancer mortality with time since first registration with a radiological society; for example, such that for those registered for more than 40 years there was a 41% excess risk. This was mostly due to those who registered between 1921 and 1954 when exposures were higher than in more recent years. As well as the problem of confounding, Brenner and Hall [26, 27] point out that a similar but much larger and more detailed study of male North American radiologists [28] showed a higher SMR for the radiologists compared with other doctors. Moreover, a study in the UK of the mortality of 20 000 NHS consultants employed between 1962 and 1979 [29] showed that SMRs for radiologists and radiotherapists were not significantly different from those for all consultants taken together, either for all causes of death or for cancer. Brenner and Hall suggest that such inconsistent results are entirely to be expected when the doses are so low (a few mSv or less per year after the 1950s), since the radiation effects are likely to be below the limit of detectability for epidemiological methods. In such a situation, most studies would be expected to show no statistically significant effects but there will be occasional ones showing slightly positive or negative results. Such results provide no evidence for health effects one way or the other, they merely rule out large risks or large benefits. A second study which has been cited [2] as supporting the hypothesis that a moderate dose rate of radiation is beneficial to health is that by Matanoski et al [30] of workers at US nuclear shipyards. This is on the basis of significantly lower SMRs seen in those workers with cumulative effective doses greater than 5 mSv than in those with lower doses, and in the latter compared with non-radiation shipyard workers. However, the authors of the study do not suggest that these results provide evidence for a beneficial effect of radiation, but instead regard selection bias as a more likely cause. Those selected to work on nuclear powered ships were given a physical examination prior to assignment and so are likely to be healthier than those working elsewhere in the shipyard. In addition, those with cumulative doses .5 mSv, mainly through being employed for longer, are further self-selected for enduring good health. Matanoski et al go on to point out that although lower SMRs are seen for all causes of death and for lung cancer, they are not seen for leukaemia where the SMR is 2.17 times higher for the radiation workers with doses .5 mSv than for those with doses ,5 mSv. Surprisingly, only an abstract of this study has appeared in the peer reviewed literature [31], but in it the authors recommended an extension of the study population and the application of more powerful methods of analysis. Report 136 of the US National Council on Radiation Protection and Measurements (NCRP) [9], which provides a detailed evaluation of the LNT model, dismisses 290

the suggestion that the nuclear shipyard worker study provides support for the beneficial effects of radiation. There have been various other studies of nuclear industry workers, reviewed by UNSCEAR [8], where the interpretation of low dose risks is not always clear, for the reasons described previously, but which do not provide any strong support for hormesis. Indeed, such studies provide some evidence of raised leukaemia risks associated with radiation exposure [10, 11]. In the UK’s National Registry for Radiation Workers [11, 32], radiation workers had lower SMRs than the general population, not because of any beneficial effect of their exposures, but simply due to the Healthy Worker Effect. Another study that has been cited as demonstrating radiation hormesis, claimed to show very low cancer rates for residents in Taiwanese office blocks built with Co60 contaminated steel [33]. However, it now seems overwhelmingly likely that this study was seriously flawed and grossly underestimated the true number of cancers in the study population [34].

Ecological studies (confounding factors) Reports of a strong negative correlation between mean natural background radiation levels and cancer mortality in different states of the USA [35] and between mean radon levels and mean lung cancer rates in different US counties [36, 37] have also been cited as further evidence for radiation hormesis [2]. However, it is well-recognized that ecological studies based on aggregated data for large geographical areas, rather than on the individual data used in case-control or cohort studies, have the potential for serious statistical problems [9]. The author of the first study [35] agreed that confounding factors such as smoking, poverty or environmental pollution could be affecting the mortality rates rather than background radiation levels [38]. However, the exceptional strength of the negative correlations between lung cancer and radon levels in the ecological studies reported by Cohen [36, 37] appears to contradict the LNT model. The fact that smoking is responsible for a large majority of lung cancers suggested that different smoking habits between the populations in the US counties might be influencing the results, but no direct data on smoking were available. Indirect evidence finally emerged when Puskin examined correlations between radon levels and a variety of cancers other than lung cancer [39]. Some of these other cancers are related to smoking and some are definitely not. Radon, of course, gives virtually all its dose to the lung so any effect of radiation, beneficial or harmful, could not be expected in cancers of any other organs. For the other smoking-related cancers, Puskin also found negative correlations with radon levels but there was no association between radon and those cancers that are not linked to smoking. He concluded that the negative correlation seen between radon levels and lung cancer in the earlier studies is largely a consequence of a negative correlation between smoking and radon levels across the US counties. Thus the results could be explained in terms of confounding by smoking without invoking any kind of beneficial effect of low level radiation exposure. The British Journal of Radiology, April 2006


Moreover, there have been numerous more recent analyses of cohort and case-control studies of residential radon concentrations and lung cancer incidence throughout the world, where smoking habit information was also available [40]. Most of them observe an excess lung cancer risk from residential radon, though the risks may not always achieve statistical significance. However, recent pooled analyses of case-control studies of radon and lung cancer from Europe [41] and from North America [42] have much more statistical power than the individual studies, and provided unequivocal evidence of the risks of domestic exposure to radon. In particular the European pooling demonstrated a risk of exposure to radon in homes down to concentrations less than 200 Bq m23 which is equivalent to an effective dose of about 5 mSv using the dose conversion convention recommended by ICRP [43].

Radiobiological studies (differences between metabolic and radiation-induced DNA damage and repair) There is general agreement that DNA lesions are continuously being produced in the body, for example by reactive oxygen species (ROS) generated by normal oxidative metabolism. The vast majority of these lesions are repaired by normal cellular processes. The total number of these endogenous DNA lesions exceeds the number produced by normal background radiation levels by several orders of magnitude. Proponents of hormesis suggest that acute low doses of radiation induce a temporary protective response against DNA damage that could counteract the everpresent endogenous DNA damage from the ROS. Some cellular and animal experiments suggest that this adaptive response to ionizing radiation appears to increase initially with dose but starts to decrease when the dose exceeds 100 mGy and disappears completely at higher acute doses. By combining such a non-linear model for a protective response with a linear, no threshold (LNT) model for radiation induced cancer, it can be argued that the net risk–dose relationship is more likely to exhibit a threshold than to be linear down to zero dose. It may even result in lower than spontaneous cancer incidence (i.e. a beneficial effect) at doses below 200 mGy [3]. However, it should be appreciated that the validity of the predictions of any mechanistic model is critically dependent on the appropriateness of the underlying assumptions. A fundamental objection that many radiobiologists have to this model is that the DNA damage caused by ROS is believed to consist mostly of base damage and single strand breaks. Ionizing radiation induces these simple DNA lesions but it also induces double strand breaks and more complex lesions [9]. It is only these latter two types of damage that are considered to be the initiating lesions in radiation-induced carcinogenesis [15]. There are also serious doubts about the validity of some of the radiobiological studies that are quoted to provide evidence for a protective effect at low doses. In experiments on chromosome aberrations in human lymphocytes, cells pre-treated with a low (‘‘adapting’’) The British Journal of Radiology, April 2006

dose of radiation apparently acquired increased resistance to a second high (‘‘challenging’’) dose. These studies have been interpreted as giving support for adaptive responses, but serious limitations have been recognized in two of them [44, 45], leading to a subsequent publication by some of the original authors and other scientists who found it very difficult to repeat the results [46]. Another study purporting to show a decreased chromosome aberration score in human lymphocytes in vitro after a single acute low dose of about 10 mGy followed by a dose-dependent increase above 50 mGy [47], could also not be repeated by other laboratories. The apparent dip in aberration yield at 10 mGy was thought to be due to an erroneous and unusually high control yield at zero dose [48]. Criticism can also be directed at two other sets of experiments [49, 50], which studied thymidine kinase activity in mouse bone marrow cells and apoptosis in mouse thymocytes, respectively. Both thymidine uptake and the percentage of apoptotic cells observed after different doses of whole body irradiation could be affected by radiation-induced perturbations in the cycling characteristics of the cells as much as by any assumed adaptive response or protective effect. The 1994 report of the UN Scientific Committee on the Effects of Atomic Radiation, UNSCEAR [17], included an annex which specifically considered evidence for adaptive responses to radiation. However, UNSCEAR decided that ‘‘it would be premature to conclude that cellular adaptive responses could convey possible beneficial effects to the organism that would outweigh the detrimental effects of exposures to low doses of radiation’’. Broadly similar conclusions were reached by NCRP [9]. In a recent draft report, an ICRP Task Group has concluded that current understanding of mechanisms and quantitative data on dose and time–dose relationships support a linear dose response at low doses with no compelling evidence for the existence of a threshold dose below which there would be no effect [15].

The impact of the LNT model on the justification and control of medical exposures One obvious implication of the LNT model is that when the doses are very low, so are the risks. There comes a point when the risks are so low they can be considered negligible, i.e. they are so small in relation to the other everyday risks that surround us that they do not need to be considered in any rational decisions about lifestyle choices. There is in a sense a dose corresponding to a threshold risk that, although not zero, is safe enough for the risks to be ignored. Above this ‘‘safe enough’’ dose there will be a range of doses where the risks are very small, but are sufficient to require some justification for allowing people to be exposed to them, in terms of an overriding benefit. As the doses and the risks increase so should the concomitant benefits, for the exposures to remain justified. This principle of justification is one of the main planks of ICRP’s recommendations regarding radiation protection for medical exposures. In recognition of the substantial potential health benefits to patients from medical exposures, ICRP does not place any restrictions on the levels of exposure that can be used 291


in diagnostic radiology. It only requires that they be justified in terms of an expected improvement in the clinical management of the patient and that all reasonable steps are taken to keep the exposures as low as possible without compromising their diagnostic efficacy (i.e. the exposures should be justified and optimized). Thus although the LNT model implies that no dose is without risk, in practice there is a dose below which the risks are considered negligible and are of no consequence in decisions regarding the radiation protection of patients. Moreover, precise quantified risk estimates are not critical for the control of medical exposures according to the justification and optimization principles. It is very difficult to quantify the benefits of diagnostic X-ray examinations in any way that is comparable with the radiation risks, so an accurate quantitative weighing of benefits against risks is usually impossible. Justification generally consists of the diagnostic radiology practitioner confirming that the exposure is clinically indicated for the patient and making a mostly subjective judgement that the expected benefits will outweigh the likely radiation risks. For such subjective judgements an approximate estimate of the risks will usually be sufficient. However, in the advice given to practitioners on radiation risks, sources of bias should still be reduced to a minimum, so that the estimated risk lies centrally within the range of uncertainty, which will unavoidably become wider as the doses become lower. Use of the LNT model at the relatively low doses typical of all diagnostic X-ray exposures (, ,100 mGy) and a DDREF of 2 to extrapolate from the effects seen in epidemiological studies, will provide sufficiently robust risk estimates for justification purposes. As well as the uncertainties in the radiation risk estimates, there are also large variations in the doses delivered to individual patients by the same type of Xray examination. The risks also depend markedly on the age and sex of the patient and might be quite different for a few individuals in the population who are genetically predisposed to cancer. Consequently, it is usually not appropriate to resolve X-ray examinations into any more than a few broad risk categories, each spanning quite a wide range of risks. This has been done in an information leaflet on the safety of X-ray examinations published by NRPB [51], where all X-ray examinations have been divided into just four risk bands each spanning a factor of ten in risk. This broad classification into negligible, minimal, very low and low risk bands is shown in Table 2 and should be sufficient for most justification purposes. Those types of examination falling into the highest risk band (still , 1 in 1000 risk of delayed cancer) are those where the potential benefits can be correspondingly high, since many of them are used in the investigation of symptoms that suggest life-threatening diseases and will easily comply with ICRP’s justification requirement. It would therefore appear that beneficial medical X-ray exposures need not be unnecessarily restricted by adoption of the ICRP recommendations for patient protection [52]. At the same time, use of the LNT model at low doses provides sufficiently reliable risk estimates to ensure that patients are being adequately protected from unnecessary medical exposures that are either unjustified or not fully optimized. 292

Table 2. X-ray examinations divided into four broad risk bands Risk banda

Risk rangeb

Typical type of X-ray examination

Negligible

, 1 in a million

Minimal

1 in a million to 1 in 100000

Very low

1 in 100000 to 1 in 10000

Low

1 in 10000 to 1 in 1000

Radiography of chest, limbs and teeth Radiography of head, neck and joints Radiography of spine, abdomen and pelvis CT, angiography, contrast studies of the alimentary, biliary and urinary tracts, and interventional radiology

a

Nomenclature according to reference [89]. Lifetime risk of cancer per examination for patients aged 30–60 years.

b

Acknowledgment We would like to thank Dr Will Atkinson for helpful discussions.

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REVIEW ARTICLE

Diagnosing a parotid lump: fine needle aspiration cytology or core biopsy? D C HOWLETT,

MRCP(UK), FRCR

Department of Radiology, Eastbourne District General Hospital, Kings Drive, Eastbourne, East Sussex BN21 2UD, UK ABSTRACT. Fine needle aspiration cytology (FNAC) has been widely adopted for the cytological diagnosis of parotid lumps. FNAC does have drawbacks, even under optimum conditions and may be associated with poor levels of diagnostic accuracy, particularly outside the specialized clinic environment. Ultrasound-guided core biopsy (USCB) is a relatively recently described technique in the parotid gland which has been well tolerated and has demonstrated a high degree of diagnostic accuracy in several studies. This article discusses the merits and pitfalls of FNAC, together with the technique of USCB and also highlights the potential advantages benefit provided by USCB in parotid diagnosis.

A broad spectrum of pathologies that present with parotid swelling and extraglandular masses can also mimic parotid lesions clinically. It is frequently difficult on clinical grounds alone to distinguish between neoplastic and non-neoplastic causes for a parotid mass and also to reliably differentiate between benign and malignant neoplasms. If an accurate pre-operative diagnosis can be achieved using a combination of imaging and cytology or histology, then many non-neoplastic lesions will not require excision. Surgery may also be avoided for certain parotid neoplasms in the elderly or unfit, e.g. Warthin’s tumour. To ensure that surgery is indicated and to allow appropriate operative planning and patient consent an accurate pre-operative diagnosis is essential. Following initial demonstration and characterization of a parotid lesion with imaging, usually ultrasound or MRI, needle biopsy is used to confirm its nature if required. Following the demise of open biopsy due to high rates of tumour seeding [1] fine-needle aspiration cytology (FNAC) has become an established technique. FNAC is most commonly performed blindly in the outpatient clinic and has a number of advantages – it is quick, safe and accurate in the hands of a skilled practitioner and high levels of diagnostic accuracy have been quoted [2–4]. This process may occur within a specialized clinic and diagnostic accuracy can be improved using ultrasound-guidance and an on-site cytologist [5]. Clinics may also be cytologist-led, although this may necessitate multiple aspirations and there is currently a severe shortage of cytology staff available with the necessary expertise, making it difficult to offer this facility outside larger centres. In the absence of ultrasound-guidance, or an on-site cytologist the diagnostic accuracy of FNAC often falls off dramatically. A recent study reported a sensitivity of only 38% in distinguishing benign from malignant disease using blind FNAC [6]. The British Journal of Radiology, April 2006

Received 21 April 2005 Revised 25 August 2005 Accepted 11 October 2005 DOI: 10.1259/bjr/74329476 ’ 2006 The British Institute of Radiology

There are also well-recognized pitfalls of FNAC even in experienced hands – particular difficulties may occur in the cytological diagnosis of pleomorphic adenoma [7, 8], Warthin’s tumour [9] and lymphoma [10]. Indeed the diagnosis of lymphoma with FNAC is not generally considered definitive, with FNAC often acting solely as a guide for the need for surgical biopsy if lymphoid proliferation is present cytologically [11]. There are specialized ancillary cytological techniques, e.g. flow cytometry and in situ hybridization, which are used to improve the diagnosis of lymphoid proliferation but these are not widely available outside larger centres. The cytological diagnosis of parotid involvement by systemic disease, e.g. sarcoidosis and Sjo¨gren’s syndrome may be difficult and it is also not possible to accurately grade or type malignant tumours or lymphomas or to distinguish in situ from invasive disease with FNAC. FNAC has a low predictive value for benign non-neoplastic lesions [3] and similarly there is a low negative predictive value if a negative FNAC result is obtained [4]. As a consequence of the diagnostic difficulties that may arise with FNAC there is a high incidence of surgical biopsy with subsequent delays in referral to the appropriate clinical team. Performing ultrasound-guided core biopsy (USCB) using a spring-loaded biopsy device does provide a recently described alternative to parotid FNAC. Core biopsy possesses an inherent advantage over FNAC in that it provides a sample of tissue for immunohistochemical analysis. This allows typing and grading of carcinomas and lymphomas and also improved differentiation of reactive nodal hyperplasia from lymphoma. A core of tissue can also be used to evaluate parotid involvement by systemic disease. An on-site cytologist is not required and USCB can be combined with an initial diagnostic ultrasound, allowing lesion characterization and compartmentalization into superficial or deep lobe. 295

D C Howlett

In two series of patients with parotid masses USCB has shown promising results with diagnostic samples obtained in all patients and reported accuracies of 100% [12] and 97% [13] when comparing core biopsy with final surgical histology. There were no reported complications of USCB and 26 out of 54 [12] and 22 out of 53 [13] patients avoided surgery as a result of core biopsy. USCB utilizes a small bore needle (18 G or 20 G) which is introduced through a small skin incision following informed written consent and infiltration with 1% lignocaine local anaesthesia. Using a high-resolution linear-array transducer for guidance, the needle-tip is positioned just adjacent to the lesion, such that following discharge the needle traverses, but does not exit deep to the lesion (Figures 1 and 2). The use of ultrasound is important as it allows avoidance of adjacent structures

during biopsy and also ensures that both the periphery and core of the lesion are sampled, with bypassing of necrotic areas, increasing diagnostic yield [12]. Some biopsy devices can allow the needle throw to be varied (Magnum gun; Bard, Covington, GA – 15–22 mm variable throw), which is useful for smaller or deeper lesions in proximity to the parotid vessels. There are important potential complications of needle biopsy including haemorrhage and facial nerve injury and also tumour seeding in the needle tract. The identification of the intraparotid vessels on ultrasound should allow the main parotid vessels and thereby the adjacent facial nerve to be avoided during biopsy. There are two reports of tumour seeding post needle biopsy [14, 15] although these occurred with large bore needles and this phenomenon is not described using smaller needle sizes. Some surgeons may choose to excise the biopsy tract at the time of operation. USCB is a little more invasive than FNAC in that it requires local anaesthesia and a skin incision. USCB also does not lend itself to the ‘‘one-stop’’ clinic setting due to longer requirements for histological reporting than a cytology aspirate. In conclusion USCB represents a recently described technique for diagnosing a parotid mass which has potential benefits over FNAC. These advantages may be particularly relevant in the district general hospital setting where ancillary cytological facilities are not routinely available. The use of USCB may help reduce the need for diagnostic surgical biopsy and thereby facilitate referral to the appropriate clinical team.

References

Figure 1. Ultrasound through the tail of the right parotid gland demonstrates a 12 mm pleomorphic adenoma overlying the angle of the mandible (M). The biopsy needle has been placed just anterior to the lesion, using a 15 mm setting on a variable-throw biopsy device.

Figure 2. The biopsy device has been discharged and the needle can be seen to traverse the mass, sampling periphery and core, but has not exited deep to the lesion.

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1. McGuirt WF, McCable BF. Significance of node biopsy before definitive treatment of cervical metastatic carcinoma. Laryngoscope 1978;88:594–7. 2. Das DK, Petkar MA, Al-Mane NM, et al. The role of fine needle aspiration cytology in the diagnosis of swellings in the salivary gland regions: a study of 712 cases. Med Princ Prac 2004;13:95–106. 3. Cohen EG, Patel SG, Lin O, et al. Fine needle aspiration biopsy of salivary gland lesions in a selected patient population. Arch Otolaryngol Head Neck Surg 2004;130:773–8. 4. Zba¨ren P, Scha¨r C, Hotz MA, Loosli H. Value of fine needle aspiration cytology of parotid gland masses. Laryngoscope 2001;111:1989–92. 5. Robinson IA, Cozens NJ. Does a joint ultrasound – guided cytology clinic optimize the cytological evaluation of head and neck masses? Clin Radiol 1999;54:312–6. 6. Balakrishnan K, Castling B, McMahan J, Imrie J, Feeley KM, Parker AJ, et al. Fine needle aspiration cytology in the management of parotid mass: a two centre retrospective study. Surgeon 2005;2:67–72. 7. Brachtel EF, Pilch BZ, Khettry V, et al. Fine needle aspiration biopsy of a cystic pleomorphic adenoma with extensive adnexa-like differentiation: differential diagnostic pitfall with mucoepidermoid carcinoma. Diagn Cytopathol 2003;28:100–3. 8. Verma K, Kapila K. Role of fine needle aspiration cytology in the diagnosis of pleomorphic adenoma. Cytopathology 2002;13:121–7. 9. Parwarni AV, Ali-Sayed Z. Diagnostic accuracy and pitfalls in the fine needle aspiration interpretation of Warthin’s tumour. Cancer 2003;99:166–71. 10. Loggins JP, Urquhart A. Pre-operative distinction of parotid lymphomas. J Am Coll Surg 2004;199:58–61.


Review article: Diagnosing a parotid lump 11. Lioe TF, Elliot H, Allen DC, Spence RA. The role of fine needle aspiration cytology (FNAC) in the investigation of superficial lymphadenopathy – uses and limitations of the technique. Cytopathology 1999;10:291–7. 12. Kesse KW, Manjaly G, Violaris N, Howlett DC. Ultrasoundguided biopsy in the evaluation of focal lesions and diffuse swelling of the parotid gland. Br J Oral Maxillofac Surg 2002;40:384–9.


13. Yung-Liang W, Siu-Cheung C, Yao-Liang C, et al. Ultrasonography-guided core-needle biopsy of parotid gland masses. Am J Neuroradiol 2004;25:1608–12. 14. Yamaguchi KT, Strong MS, Shapsey SM, Soto E. Seeding of parotid carcinoma along Vim-Silverman needle tract. J Otolaryngol 1979;81:49–52. 15. Peacock EE, Byars LT. Management of tumours of the parotid gland. N C Med J 1958;19:1–9.

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Intra-arterial MR angiography in the iliac system: initial clinical experience with 25 patients C PAETZEL, MD, N ZORGER, MD, O W HAMER, MD, J SEITZ, MD, T SCHLEICHER, MD, S FEUERBACH, MD, W R NITZ, PhD, M LENHART, MD and T HEROLD, MD Department of Diagnostic Radiology, University of Regensburg, Hospital, 93042 Regensburg, Germany ABSTRACT. The aim of this study was to evaluate intra-arterial magnetic resonance angiography (MRA) of the iliac arteries. Therefore, 25 consecutive patients (17 male, 8 female) suffering from symptomatic occlusive disease of the lower limbs were investigated prospectively. Catheter angiography was performed before MRA and served as the standard of reference. Contrast-enhanced intra-arterial MRA was performed using a 1.5 Tesla MRI system. Contrast agent (gadodiamide) was injected by a conventional pigtail-shaped angiography catheter placed in the abdominal aorta. Vascular lesions were assessed by four investigators. The degree of stenosis was compared with the findings of conventional catheter angiography. Additionally, the diagnostic quality of the MR angiograms was assessed by the investigators using a semi quantitative five-point scale. All lesions shown by catheter angiography were detected and correctly localized by intra-arterial MRA. MR angiograms exhibit a specificity of 95% and a sensitivity of 96% for stenoses of 50% or more. The diagnostic quality of the images was judged from good to excellent, on average. Intra-arterial MRA exhibits a specificity and sensitivity comparable with intravenous angiography. The image quality appears to be adequate for supporting MR-guided vascular intervention.

Fluoroscopy-guided percutaneous interventional procedures like transluminal angioplasty are well established methods for the treatment of peripheral arterial occlusive disease of the lower limbs. Although MR guidance of manoeuvres for vascular intervention provides several advantages, such as the absence of ionizing radiation, avoidance of potential hazards from iodinated contrast material and superior soft tissue contrast, it has not yet gained acceptance in the clinical routine. Specific problems with MR-guided intervention are the lack of availability of MR compatible materials (e.g. stents, guidewires), their visibility and the associated high costs. Furthermore, the precise localization of stenosis and post-interventional workup may cause difficulties. Manke et al [1] showed the feasibility of MRI-guided stent placement in iliac artery stenoses in 14 cases using intravenous magnetic resonance angiography (MRA) for pre- and post-interventional imaging. Due to the restricted total amount of applicable MR contrast agent, recurrent control examinations cannot be performed. Manke reported one stent misplacement due to patient motion during the procedure while relying on a previously acquired MRA roadmap. A newer report describes MR-guided angioplasty using intra-arterial MRA [2]. Intra-arterial MRA with a catheter positioned in the infrarenal abdominal aorta allows a reduction in the amount of contrast material applied. Based on this approach, the correct intraluminal site of the catheter can Current address for Christian Paetzel: Department of Diagnostic Radiology, Klinikum Weiden, So¨llnerstraße 16, 92637 Weiden, Germany.

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be identified, and repeated control examinations due to patient motion can be performed. There are reports regarding this technique in the literature, but sensitivity and specificity have not yet been systematically assessed [3, 4]. The aim of this study was to investigate the sensitivity, specificity and the image quality of intra-arterial MR angiograms of the iliac arteries using conventional catheter angiography as the standard of reference.

Materials and methods Patient population Over a time period of 5 months, 25 consecutive patients (17 men, 8 women; mean age 61.4 years, range 33– 82 years, standard deviation ¡8.5) suffering from peripheral arterial occlusive disease of the lower limbs underwent intra-arterial conventional catheter angiography and intra-arterial MRA of the iliac arteries. The grading of the disease according to the classification of Table 1. Patient population: Fontaine’s classification of peripheral arterial occlusive disease Grading (Fontaine’s classification)

Number of patients

I II a II b III IV

0 3 16 3 3


Intra-arterial MRA in the iliac system

Figure 1. Examination of a 53-year-old man suffering from symptomatic arterial occlusive disease of the lower limbs with rest pain. (a) Maximum intensity projection of the intra-arterial magnetic resonance angiography showing an occlusion of the left common iliac artery. The artefact of the angiography catheter is indicated by an arrow. (b,c) Corresponding conventional catheter angiography confirms the finding of the MR angiogram.

Fontaine is presented in Table 1. Written informed consent was obtained from all patients in advance. The study protocol was approved by the institutional review board.

Protocol of conventional angiography and intraarterial MRA Digital subtraction angiography (DSA) was the standard of reference in all patients. Femoral access with a The British Journal of Radiology, April 2006

5 F sheath was performed in the angiography suite. To avoid thromembolic complications due to the extended investigation time, all patients received 2500 IE heparin intravenously and 40 mg butylscopolamin. A pigtail catheter was placed in the infrarenal aorta at the level of the second lumbar vertebral body. Non-ionic contrast material with 320 mgI ml21 was injected intra-arterially (volume 20 ml; injection rate 10 ml s21; frame rate 2 s21). Imaging was performed using the 35 ˚ left and right anterior oblique position. The degree of stenosis was 299

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Figure 2. 62-year-old man with ischaemic ulcer of the right foot. (a) Intra-arterial MR angiography detects a high grade stenosis of the right external and an arteriosclerotic dilation of the left common iliac artery as effectively (b) as conventional angiography. The arrow marks the artefact of the angiography catheter.

assessed using the commercially available software on the Polytron T.O.P. (Siemens, Erlangen, Germany). Using this software to analyse the angiographic findings, the degree of stenosis was calculated automatically by comparing the geometry and density of a lesion. After conventional DSA the patients were taken to the MR-scanner with a permanently flushed pigtail-catheter. MRA images were acquired on a 1.5 T scanner (Sonata; Siemens Medical Solutions) with a gradient strength of 40 mT m21 and a slew rate of 200 T m21 s21 using a phased array receiver coil and breath-hold acquisitions (22 s). A fast-low-angle-shot (FLASH) sequence was utilized with elliptical scanning encoding, 40 partitions (75% partial Fourier), 280 mm 6 263 mm field of view, 1.5 mm partition thickness, 25 ˚ excitation angle, 266 6 512 matrix size (75% partial Fourier) with an echo time of 1.92 ms and a repetition time of 5.87 ms for a bandwidth of 200 Hz/pixel. 10 ml of gadodiamide (Omniscan; Amersham Buchler, Braunschweig, Germany) was applied via the arterial catheter with a concentration of 0.5 mmol l21, which was diluted with 50 ml of 0.9% of saline solution at an injection rate of 3.5 ml s21. The injection was followed by 20 ml of 0.9% saline solution with an injection rate of 3.5 ml s21 [4]. The application of the contrast agent was performed with an MR injection system (Spectris; Medrad, Indianola, Iowa). A maximum intensity projection (MIP) algorithm was applied to all contrast-enhanced MR angiography studies after subtraction of the unenhanced measurement by using the commercially available software on the MRI system (Magnetom Symphony, software Numaris 3.5, version VA11A; Siemens). MIPs were reconstructed in steps of 9 ˚ (range of 180 ˚). 300

Image analysis Four radiologists experienced in cardiovascular imaging assessed the MR angiograms independently in a randomized order for stenoses of the common and external iliac arteries and image quality of MRA. Observers were not aware of DSA findings when analysing contrast enhanced MRA, and vice versa. Image analysis was based on original contrastenhanced data sets, MIPs and digital subtraction angiograms. None of the observers were aware of the clinical history or the interpretations of the other observers. The degree of stenosis was defined as the ratio of the narrowest diameter (A) within the stenosis and the diameter of the nearest downstream uninvolved segment of the artery (B): 100 6 (1–A/B). The stenoses were classified into occlusions, high grade stenoses (75– 99%), moderate stenoses (50–74%), mild stenoses (25– 49%) and non-stenoses (0–24%). Additionally, the four radiologists assessed MR angiograms independently for the following: diagnostic value of the infrarenal abdominal aorta; common iliac arteries; internal and external iliac arteries, and the overall impressions were rated on a subjective scale (15excellent, no limitations; 25good, minor limitations; 35moderate, moderate limitations; 45poor, major limitations but still of diagnostic quality; 55non-diagnostic). Intermediate scores at 0.5 intervals were not allowed. The catheter angiograms were assessed by a consensus of two of the investigators. The degree of stenosis was quantified and the lesions were classified in the same manner as described for the MR angiograms. If there was no vascular lesion detected by either catheter angiography The British Journal of Radiology, April 2006

Intra-arterial MRA in the iliac system Table 2. Classification of stenoses; results of intra-arterial MR angiography compared with conventional catheter angiograms Degree of stenosis (number of lesions)

Underestimated

Correct scores

Overestimated

Occlusion (n54) 75–99% (n53) 50–74% (n512) 25–49% (n510) 0–24% (n511) All patients (n540)

0% (0/16) 0% (0/12) 6.3% (3/48) 7.5% (3/40)

100% (16/16) 100% (12/12) 89.6% (43/48) 82.5% (33/40) 79.5% (35/44) 86.9% (139/160)

0% (0/12) 4.2% (2/48) 10% (4/40) 20.5% (9/44) 9.4% (15/160)

3.8% (6/160)

or intra-arterial MRA, it was scored as a patient with one lesion with a stenosis of 0%.

Results In total, 40 vascular lesions were assessed by catheter angiography. There were 10 patients with one lesion, nine patients with two, three patients with three and three patients with none. The catheter angiography detected four occlusions (Figure 1), three high grade (Figure 2), 12 moderate and 10 mild stenoses (Table 2). All of the seven lesions with a degree of stenosis higher than 75% were correctly assessed by intra-arterial MRA. The moderate stenoses were overestimated by one of the investigators in two cases out of 48 (4.2%) and underestimated in three cases (6.3%). Mild stenoses were overestimated in 4 out of 40 (10%), no stenoses in 9 out of 44 (20.5%) cases. Mild stenoses were underestimated three times (7.5%). The lesions which were assessed incorrectly by at least one of the observers are outlined in Table 3. With respect to stenoses of more than 50%, the sensitivity and specificity were calculated at 96% and 95%, respectively. The diagnostic value of all intra-arterial MR angiograms was judged to be 1.54 (range 1–4, median 1) on average by all investigators. The absolute and relative frequency of the scores for diagnostic quality is listed in Table 4.

Table 3. Lesions scored incorrectly by at least one observer Degree of stenosis (catheter angiography)

Observer 1

Observer 2

Observer 3

Observer 4

0.19 0.20 0.22 0.22 0.24 0.29 0.33 0.42 0.48 0.49 0.51 0.58 0.59 0.69

0 0 + 0 + 0 0 + + 0 2 0 0 0

+ 0 0 + + 0 2 0 + 0 0 0 0 0

+ + 0 + 0 0 0 0 0 0 2 + 0 0

0 0 + 0 0 2 2 0 0 + 0 0 2 +

0, Correct score; 2, underestimated; +, overestimated.


Discussion MR-guided vascular interventions provide many advantages over fluoroscopy-guided techniques as mentioned above. Nevertheless, MR-guided procedures are not yet well established in the clinical routine. The reasons for this are the lack of availability of visible materials, real time imaging during the intervention and the peri-interventional check-up. Manke et al used commercially available catheters and guidewires [1] for the treatment of iliac artery stenoses by stent placement based on the safety precautions as described by Nitz et al [4]. Pre-interventional imaging was performed using intravenous MRA. Even though the usefulness of intravenous MRA for imaging the iliac arteries has been demonstrated by several authors [6–16], the restricted total amount of contrast medium of approximately 0.3 mmol kg21 body weight prohibits the investigator from performing repeated angiograms, which can lead to stent misplacement due to patient motion. To avoid this, intra-arterial MRA was used in a recent study for periinterventional imaging [2]. Thus, the contrast agent dose is reduced 50% compared with MRA with intravenous injection, allowing the investigator to perform repeated angiograms up to four times. The contrast agent dose chosen was based on a conservative level for a wide range of interindividual conditions. A further reduction in the total amount of contrast agent administered seems possible. Even if initial reports about intra-arterial MRA already exist in the literature [3, 4], there is no further experience with this method concerning its sensitivity and specificity. In this study, all of the 40 lesions assessed by catheter angiography were detected and correctly localized by intra-arterial MRA. Thus, verification of lesions is possible using intra-arterial angiography with a sensitivity of 100% in this study. The indication to treat a lesion depends on the degree of the stenosis and the clinical symptoms. The degree of stenosis was correctly classified in 86.9% (139/160); stenoses were overestimated in 9.4% (15/160) and underestimated in 3.8% (6/160) of the cases. Focusing on obstructions or stenosis of more than 50%, the sensitivity and specificity of approximately 95% is Table 4. Estimation of image quality by four investigators in 25 patients Score

1

2

3

4

5

Absolute frequency Relative frequency

58 0.58

31 0.31

10 0.10

1 0.01

0 0

1, excellent; 2, good; 3, moderate; 4, poor; 5, non-diagnostic.

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comparable with intravenous MRA [6, 10–15]. The value of contrast-enhanced intravenous MRA has been demonstrated in a study by Lenhart et al, showing that MRA can take the place of catheter angiography in the routine work-up of patients with peripheral arterial occlusive disease. Additional examinations are necessary in approximately 5% of the patients, mainly due to inadequate diagnostic quality of the vessels of the lower leg. Thus, with respect to the clinical symptoms, intravenous and intra-arterial MRA give us sufficient information for the treatment of iliac artery lesions in patients with symptomatic arterial occlusive disease of the lower limbs. The outcome of this study is comparable with the results of a recent study concerning the accuracy of intra-arterial MRA of the femorocrural arteries [16]. Using gadodiamide for intra-arterial injection represents an ‘‘off-label’’ use. There are some studies dealing with the intra-arterial application of gadolinium-complex in catheter angiography [17, 18]. The reasons for such application are allergic reactions to iodide contrast agents, renal insufficiency and hyperthyroidism. We used gadodiamide due to the low osmolality of 780 mmol kg21 compared with the other available gadolinium complexes. Similar to reports in the literature, no adverse reactions where observed during this study. The confidence of the radiologists in the diagnostic quality of the intra-arterial MR angiograms is demonstrated by their subjective assessment of the image quality. They judged 89% of the investigations as good or excellent using a five point scale. All MR angiograms were assessed as having diagnostic quality. This result indicates that no additional examinations would have been required by the readers to make decisions about further treatment.

Conclusion Lesions of the iliac arteries can be detected and localized by intra-arterial MRA. MR angiograms presented with a sensitivity and specificity comparable with intravenous MRA. Thus, intra-arterial MRA seems to support MR-guided interventional procedures.

References 1. Manke C, Nitz WR, Djavidani B, Strotzer M, Lenhart M, Vo¨lk M, et al. MR imaging-guided stent placement in iliac arterial stenoses: a feasibility study. Radiology 2001;219:527–34. 2. Paetzel C, Zorger N, Bachthaler M, Vo¨lk M, Seitz J, Herold T, et al. Feasibility of MR guided angioplasty of femoral artery stenoses using real-time imaging and intraarterial contrast-enhanced MR-angiography. Rofo 2004;176:1232–6. 3. Tello R, Mitchel PJ, Melhem ER, Witte D, Thomson KR. Interventional catheter magnetic resonance angiography with a conventional 1.5-T magnet: work in progress. Australas Radiol 1999;43:435–9.

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4. Paetzel C, Zorger N, Seitz J, Vo¨lk M, Nitz WR, Herold T, et al. Intraarterial contrast-enhanced magnetic resonance angiography of the aortoiliac system. J Vasc Interv Radiol 2004;15:981–4. 5. Nitz WR, Oppelt A, Renz W, Manke C, Lenhart M, Link J. On the heating of linear conducting structures as guidewires and catheter in interventional MR. J Magn Reson Imaging 2001;13:105–14. 6. Kai Yiu JAM, Leiner T, de Haan MW, Kessels AG, Kitslaar PJ, van Engelhoven J. Periphereal vascular tree stenoses: evaluation with moving-bed infusion-tracking MR angiography. Radiology 1998;206:683–92. 7. Quinn SF, Sheley RC, Semonsen KG, Leonardo VJ, Kryss K, Szumowski J. Aortic and lower-extremity arterial disease: evaluation with MR angiography versus conventional angiography. Radiology 1998;206:693–701. 8. Snidow JJ, Johnson MS, Harris VJ, Margosian PM, Aisen AM, Lalka SG, et al. Three-dimensional gadoliniumenhanced MR angiography for aortoiliac inflow assessment plus renal artery screening in a single breath hold. Radiology 1998;198:725–32. 9. Lenhart M, Finkenzeller T, Paetzel C, Strotzer M, Mann S, Djavidani B, et al. Contrast-enhanced MR angiography in the routine work-up of the lower extremity arteries. Rofo 2002;174:1289–95. 10. Sueyoshi E, Sakamoto I, Matsuoka Y, Ogawa Y, Hayashi H, Hashmi R, et al. Aortoiliac and lower extremity arteries: comparison of three-dimensional dynamic contrastenhanced subtraction MR angiography and conventional angiography. Radiology 1999;210:683–8. 11. Lenhart M, Herold T, Volk M, Seitz J, Manke C, Zorger N, et al. Contrast media-enhanced MR angiography of the lower extremity arteries using a dedicated peripheral vascular coil system. First clinical results. Rofo 2000;172:992–9. 12. Lenhart M, Djavidani B, Volk M, Strotzer M, Manke C, Requardt M, et al. Contrast medium-enhanced MR angiography of the pelvic and leg vessels with an automated table-feed technique. Rofo 1999;171:442–9. 13. Kreitner KF, Kalden P, Neufang A, Duber C, Krummenauer F, Kustner E, et al. Diabetes and peripheral arterial occlusive disease: prospective comparison of contrastenhanced three-dimensional MR angiography with conventional digital subtraction angiography. AJR Am J Roentgenol 2000;174:171–9. 14. Huber A, Scheidler J, Wintersperger B, Baur A, Schmidt M, Requardt M, et al. Moving-table MR angiography of the peripheral runoff vessels: comparison of body coil and dedicated phased array coil systems. AJR Am J Roentgenol 2003;180:1365–73. 15. Douek PC, Revel D, Chazel S, Falise B, Villard J, Amiel M. Fast MR angiography of the aortoiliac arteries and arteries of the lower extremity: value of bolus-enhanced, wholevolume subtraction technique. AJR Am J Roentgenol 1995;165:431–7. 16. Zorger N, Vo¨lk M, Hamer OW, Lenhart M, Seitz J, Butz B, et al. Intraarterial gadolinium-enhanced MR angiography in humans: feasibility and accuracy for detection of infrainguinal arterial stenoses before and after percutaneous angioplasty. AJR Am J Roentgenol 2005;185:867–72. 17. Schild HH, Weber W, Boeck E, Mildenberger P, Strunk H, Duber C, et al. Gadolinium-DTPA (Magnevist) as a contrast medium for arterial DSA. Rofo 1994;160:218–21. 18. Hatrick AG, Jarosz JM, Irvine AT. Gadopetate dimeglumine as an alternative contrast agent for use in interventional procedures. Clin Radiol 1997;52:948–52.



Prevalence and distribution of adnexal findings suggesting endometriosis in patients with MR diagnosis of adenomyosis T T ZACHARIA,

MD

and M J O’NEILL,

MD

Abdominal Imaging and Interventional Radiology, Massachusetts General Hospital, Harvard Medical School, USA ABSTRACT. The purpose of this investigation was to establish the prevalence and distribution of MR findings associated with pelvic endometriosis in patients with a MRI diagnosis of adenomyosis. Retrospective review of the pelvic MRI in 59 patients (age 32–54 years, mean 42 years) who met strict MRI criteria for adenomyosis was performed. T1 weighted fat saturated and T2 weighted images in these patients were reviewed for the presence or absence of T1 bright signal suggesting endometriosis in any of nine locations (uterine serosa, right and left ovary, right and left fallopian tube, right and left broad ligament, and right and left pelvic side wall). 20 (20/59) patients (34%), showed characteristic MRI features associated with endometriosis. A total of 54 sites of involvement were identified (uterine serosa n517, ovaries n514, broad ligaments n510, fallopian tubes n58, pelvic side walls n55) in 20 patients with an average of 2.7 sites per patient. Implants (n543) were more common than endometriomas (n511). Endometriomas occurred most often in the ovaries (ovaries n59, broad ligament n52) whereas implants were seen on all locations (uterine serosa n517, ovaries n55, broad ligaments n58, fallopian tubes n58, pelvic side walls n55). One third of patients with adenomyosis diagnosed by MRI also had MRI findings of endometriosis, with serosal implants being the most common finding. Imaging protocols should routinely include T1 weighted fat saturated imaging sequences in order to detect coexistent endometriosis in patents undergoing pelvic MRI for the diagnosis of adenomyosis.

Adenomyosis and endometriosis are gynaecological disorders with similar pathological appearance and proposed aetiologies. We sought to define whether these disease entities coexist. Recent studies have shown high sensitivity and specificity of MRI in the diagnosis of endometriosis [1, 2]. The detection and characterization of endometrioma is relatively easier when compared with detection of serosal endometriotic implants. A diagnosis of serosal implants can be easily missed during routine pelvic MRI evaluation. Some investigations performed recently have shown MRI to have very high sensitivity and specificity in the diagnosis of adenomyosis [3, 4]. MRI has shown added value in the evaluation of associated diseases like uterine leiomyomas when compared with transvaginal ultrasound (TVUS). Similar theories of abnormal migration and metaplasia [5–7] are postulated in the aetiology of both adenomyosis and endometriosis. The most commonly proposed aetiology for adenomyosis is direct migration of endometrial tissue into the myometrium. This mechanism of direct extension is similar to the most widely accepted Address correspondence to: T Thomas Zacharia, MD, #4225, Elmhurst Ave, Ithaca Street, Queens, New York City, NY 11373, USA. Presented at a meeting: 104th ARRS (American Roentgen Ray Society) meeting, Miami, Florida, May 2004.


Received 19 May 2005 Revised 17 July 2005 Accepted 12 August 2005 DOI: 10.1259/bjr/70121266 ’ 2006 The British Institute of Radiology

theory for the aetiology of endometriosis, i.e. reflux and retrograde extension of endometrial tissue into the peritoneal cavity through the fallopian tube. To date, no large series of patients with both disease entities has been reported. The purpose of this investigation was to establish the prevalence and distribution of MR findings associated with pelvic endometriosis in patients with a MRI diagnosis of adenomyosis.

Materials and methods Patients Retrospective review of an MRI database from 1 January 2000 to 6 January 2003 identified 63 patients (aged 32–54 years, mean 42 years) in whom the diagnosis of adenomyosis was included in the differential diagnosis. Indications for MRI included menstrual irregularities (like dysmenorrhoea, menorrhagia), pelvic pain, infertility and clinically palpable uterus. IRB (Institutional Review Board) approval and informed consent was obtained in all patients. Retrospective review by a subspecialty trained women’s imaging radiologist of the pelvic MRI exams in these patients found 59 patients who met strict MRI criteria (Reinhold’s criteria) for adenomyosis (junctional zone (JZ) .12 mm and sub endometrial T2 bright foci within abnormal JZ) [8] (Figure 1a). These 59 patients served as the 303

T T Zacharia and M J O’Neill

(a)

(b)

(c) Figure 1. 35-year old woman with posterior wall adenomyoma and coexisting right ovarian endometrioma. (a) Sagittal T2 weighted MR image (repetition time (TR) 5800/echo time (TE) 100) shows classical findings of adenomyosis involving the posterior wall of the uterus. Ill defined adenomyoma with T2 bright focus within (arrow). (b) Coronal T2 weighted MR image (TR 5000/TE 125) shows right ovarian endometrioma with T2 shading (arrow). (c) Axial T1 weighted fat saturated MR image (TR 150/ TE 1.8) shows very hyperintense right ovarian endometrioma (arrow).

investigational study group. 20 (20/59) patients had severe symptoms. Hysterectomy was performed in all these 20 patients. To serve as a control group there 304

should ideally have been a review of MR exams performed for the clinical indication of endometriosis, to evaluate co-existence of adenomyosis. However, a The British Journal of Radiology, April 2006

Endometriosis coexisting with adenomyosis

saturated T1 weighted imaging. The diagnosis of endometrial implant was made after ruling out artefacts of motion, incomplete fat saturation and bowel contents. Lesions larger that 10 mm were considered to represent endometriomas (Figure 1b,c) and lesions less than 10 mm were classified as implants [9]. The sites (n59) evaluated included uterine serosa, bilateral ovaries, bilateral fallopian tubes, bilateral broad ligaments and pelvic side walls. Fallopian tubes were distinguished from broad ligament on coronal MR images as thin tubular structures.

converse study (detection of adenomyosis in patients with MR diagnosis of endometriosis) was not performed as part of our investigation.

MRI protocols MRI was performed with 1.5 T signa scanners (General Electric Medical Systems, Milwaukee, WI). T2 weighted images were acquired with 4 mm slices with 1 mm spacing in the sagittal, coronal, and axial planes relative to the orientation of uterine cavity, using fast (turbo) spin echo sequences in all three planes. T1 weighted spin echo and gradient echo images were obtained both with and without fat saturation in all 59 patients. High resolution surface coils were used for data acquisition and exams were completed in 30–45 min.

Data analysis The reviewed data were collected on a Microsoft Excel (Redmond, DC) spreadsheet. The data entered included patient age, sex, MRI findings and sites of involvement. The percentage of patients with coexistent endometriosis and adenomyosis was calculated.

Image analysis T1 weighted fat saturated and T2 weighted images in these patients were then reviewed for the presence or absence of T1 bright/T2 dark signal suggesting endometriosis in any of nine anatomical locations. Endometrial implant has a varied appearance depending on the age of associated blood products. Diagnosis of implant was made when hyperintense lesions were detected on fat

Results Approximately one third of patients diagnosed to have adenomyosis by MRI were found to have coexisting endometriosis. 20 (20/59) patients with adenomyosis (34%, 95% confidence interval (CI) 22–47%)) had coexisting endometriosis (Figure 2).

(a)

(b)

Figure 2. 38-year-old woman with adenomyosis and serosal endometriosis. (a) Sagittal T2 weighted MR image (repetition time (TR) 5800/echo time (TE) 100) shows junctional zone thickening (arrow) consistent with adenomyosis. (b) Sagittal gradient echo T1 weighted fat saturated MR image (TR 150/TE 1.8) shows hyperintense focus (arrow) suggestive of serosal implant.

Table 1. Distribution of endometriotic lesions

Endometrioma Implants Total

Uterine serosa

Ovary

Broad Ligament

Fallopian tubes

Sidewall

Total

0 17 17

9 5 14

2 8 10

0 8 8

0 5 5

11 43 54


305

T T Zacharia and M J O’Neill

Uterine serosa and ovaries were the most common sites of involvement in our series (Table 1). Single site involvement was noted in two patients and multiple sites were involved in 18 patients. Extra ovarian involvement was noted in all patients (n520).

Sites of endometriosis The various sites of involvement studied were uterine serosa, bilateral ovaries, broad ligaments, fallopian tubes and the pelvic wall. A total of 54 sites of involvement were identified (uterine serosa n517/54, ovaries n514/ 54, broad ligaments n510/54, fallopian tubes n58/54, pelvic side walls n55/54) in 20 patients with an average of 2.7 sites per patient.

Implants Implants (n543) were more common than endometriomas (n511) and formed 79% of the lesions. Implants (lesions less than 10 mm) were most frequently seen over the uterine surface (Figure 2b). Most of these lesions involving uterine serosa were more than 2 mm thick (Figure 1), but occasionally these lesions were subtle and barely perceptible (Figure 2). Implants were seen on all locations (uterine serosa n517, ovaries n55, broad ligaments n58, fallopian tubes n58, pelvic side walls n55).

Endometriomas Endometriomas were less common than serosal implants (n511) (20.4%). Endometriomas occurred most often in the ovaries (ovaries n59, broad ligament n52). ‘‘T2 shading’’ was observed in some of these lesions (Figure 1b). The diagnosis was obvious also on T1 fat saturated sequence.

Histopathology Histopathological confirmation was obtained in 20 patients (20/59) (34%) who underwent hysterectomy. All of them had findings correlating with MRI. Seven (7/20) patients had coexisting adenomyosis and endometriosis. The presence of multiple patchy or diffuse areas of endometrial glands and/or stroma associated with surrounding muscle hypertrophy located more than halfway from the endometrial-myometrial junction was considered diagnostic of adenomyosis on histopathology. Endometriosis was diagnosed when endometrial tissue (endometrial gland and stroma) was found at histopathological examination of resected specimen in uterine serosa, ovaries, broad ligaments, fallopian tubes or pelvic sidewalls.

Discussion Common aetiological factors like abnormal migration, metaplasia and catalase are implicated in the pathogenesis of both endometriosis and adenomyosis [5–7]. Sampson’s theory of retrograde menstruation is the most 306

accepted theory postulated in the pathogenesis of endometriosis and has been supported by convincing circumstantial and experimental evidence [5]. Metaplasia of serosal epithelium is another postulated theory in the aetiology of endometriosis [7]. Adenomyosis is a fairly frequent disorder in adult women characterized by infiltration of endometrial glands and stroma within the myometrium of the uterus. Recent experimental data have supported the hypothesis that adenomyosis represents a condition of the uterine body in which the stromal cells have a primary pathogenetic role, although some contribution of accelerated epithelial down growth cannot be entirely excluded [7]. This migration of endometrial epithelium to the inner myometrium can act as an irritant focus and cause smooth muscle hyperplasia. There are a number of clinical and pathological studies, which throw light on the coexistence of both the diseases [1, 7, 10]. However, no imaging study exists which has studied the coexistence of adenomyosis and endometriosis. We sought to evaluate the patients with adenomyosis who had coexisting serosal or ovarian endometriosis on MRI. MRI findings in endometriosis has been well documented in the literature. The endometriotic lesions are typically hyperintense on T1 weighted fat saturated sequence with ‘‘shading’’ on T2 weighted imaging [11– 13]. Identification of these lesions by MRI relies on detection of haemorrhagic lesions. Signal characteristics vary according to the age of the haemorrhage. Typically, these lesions appear hyperintense on T1 weighted spin echo images and hypointense (shading) on T2 weighted fast spin echo images (T2WFSE) due to the presence of deoxyhaemoglobin and met haemoglobin. Acute haemorrhage occasionally appears hypointense on T1WSE and T2WFSE sequences. Old haemorrhage occasionally appears hyperintense on T1WSE and T2WFSE images. T1 weighted fat saturation sequence was performed in all our patients. The addition of fat saturated T1 weighted imaging improved diagnostic accuracy in the evaluation of both endometriomas and peritoneal disease by increasing lesion conspicuity and differentiating lipid containing ovarian masses from those containing blood. Recent studies have shown that administration of gadolinium will enhance detection of deep endometriotic implants. The coexistence of endometriosis and adenomyosis has got significant clinical implications. The purpose of this investigation is to prove the association between endometriosis and adenomyosis. The observation that one third of patients with adenomyosis have coexisting endometriosis should alert the radiologist to look for endometriosis, which is quite often subtle, represented by uterine serosal implants. The pelvic sidewall implants can also be easily missed. In patients presenting with infertility, the coexistence of these two conditions can result in treatment failures if both the diseases are not detected and treated. Hysterectomies can be avoided if endometriosis can be diagnosed and treated by drugs like progestin, GnRH (gonadotrophin releasing hormone) analogues and Danazol [14, 15]. If symptomatic relief is attained in this group of patients after medical management of endometriosis, hysterectomy may not be needed to treat adenomyosis. The British Journal of Radiology, April 2006

Endometriosis coexisting with adenomyosis

Our findings have important implications in deciding the MRI protocols too. Fat saturated T1 weighted imaging should be routinely included in the pelvic MRI protocol done for patients undergoing evaluation for adenomyosis. The most important drawback of our study is the lack of pathological evidence in patients who did not have the severe symptoms to undergo hysterectomy. Although MRI has high specificity in the diagnosis of both these disease entities, certain centres depend upon deep myometrial biopsy for the diagnosis of adenomyosis and laparoscopic biopsy for the diagnosis of endometriosis [16]. Another limitation is the absence of a control group. There should have been a converse study (a review of MR exams performed for endometriosis, to evaluate associated adenomyosis) and this was not performed as part of our investigation. Without rigorous statistics, this association will be difficult to prove absolutely. However, similar pathogenesis, histological appearance and presentation in a similar age group suggest that the association might be real.

3.

4.

5.

6.

7.

8.

9.

Conclusion MRI evidence of endometriosis was seen in one third of patients diagnosed to have adenomyosis on MRI. Serosal implants especially on the uterine surface were the most common observation. Endometriomas were also commonly noted. Imaging protocols should routinely include T1 weighted fat saturated imaging sequences in order to detect coexistent endometriosis in patents undergoing pelvic MRI for the diagnosis of adenomyosis.

References 1. Dumontier I, Roseau G, Vincent B, et al. Comparison of endoscopic ultrasound and magnetic resonance imaging in severe pelvic endometriosis. Gastroenterol Clin Biol 2000;24:1197–204. 2. Stratton P, Winkel C, Premkumar A, et al. Diagnostic accuracy of laparoscopy, magnetic resonance imaging, and


10. 11.

12. 13. 14.

15.

16.

histopathologic examination for the detection of endometriosis. Fertil Steril 2003;79:1078–85. Bazot M, Cortez A, Darai E, Rouger J, Chopier J, Antoine JM, et al. Ultrasonography compared with magnetic resonance imaging for the diagnosis of adenomyosis: correlation with histopathology. Hum Reprod 2001;16:2427–33. Dueholm M, Lundorf E, Hansen ES, Sorensen JS, Ledertoug S, Olesen F. Magnetic resonance imaging and transvaginal ultrasonography for the diagnosis of adenomyosis. Fertil Steril 2001;76:588–94. Sampson JA. Benign and malignant endometrial implants in the peritoneal cavity, and their relation to certain ovarian tumors. Surg Gynecol Obstet 1924;38:287–311. Ota H, Igarashi S, Sato N, Tanaka H, Tanaka T. Involvement of catalase in the endometrium of patients with endometriosis and adenomyosis. Fertil Steril 2002;78:804–9. Starzinski-Powitz A, Zeitvogel A, Schreiner A, Baumann R. In search of pathogenic mechanisms in endometriosis: the challenge for molecular cell biology. Curr Mol Med 2001;1:655–64. Reinhold C, McCarthy S, Bret PM, Mehio A, Atri M, Zakarian R, et al. Diffuse adenomyosis: comparison of endovaginal US and MR imaging with histopathologic correlation. Radiology 1996;199:151–8. Manfredi R, Valentini AL. Magnetic resonance imaging of pelvic endometriosis. Rays 1998;23:702–8. Sharpe-Timms KL. Endometrial anomalies in women with endometriosis. Ann NY Acad Sci 2001;943:131–47. Gougoutas CA, Siegelman ES, Hunt J, Outwater EK. Pelvic endometriosis: various manifestations and MR imaging findings. AJR Am J Roentgenol 2000;175:353–8. Umaria N, Olliff JF. Imaging features of pelvic endometriosis. Br J Radiol 2001;74:556–62. Glastonbury CM. The shading sign. Radiology 2002;224:199–201. Ylanen K, Laatikainen T, Lahteenmaki P, Moo-Young AJ. Subdermal progestin implant (Nestorone) in the treatment of endometriosis: clinical response to various doses. Acta Obstet Gynecol Scand 2003;82:167–72. Rotondi M, Labriola D, Rotondi M, Ammaturo FP, Amato G, Carella C, et al. Depot leuprorelin acetate versus danazol in the treatment of infertile women with symptomatic endometriosis. Eur J Gynaecol Oncol 2002;23:523–6. McLucas B, Perrella R, Adler L. Embolization for the treatment of adenomyosis. AJR Am J Roentgenol 2002;178:1028–9.

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Influence of post-treatment delay on the evaluation of the response to focused ultrasound surgery of breast cancer by dynamic contrast enhanced MRI A KHIAT,

PhD,

D GIANFELICE,

MD,

M AMARA,

MSc

and Y BOULANGER,

PhD

De´partement de Radiologie, Hoˆpital Saint-Luc du CHUM, 1058 St-Denis, Montreal, Quebec, H2X 3J4 Canada ABSTRACT. The assessment of the effectiveness of MRI-guided focused ultrasound surgery (MRIgFUS) of breast carcinomas can be performed by dynamic contrastenhanced magnetic resonance imaging (DCE-MRI) parameters which monitor the presence of residual tumour. The aim of this study was to evaluate the effect of the post-treatment delay on this assessment. DCE-MRI data were acquired immediately and 3–14 days after MRIgFUS treatment of 26 tumours (,7 days, n56; ¢ge;7 days, n520). The percentage of residual tumour was determined histologically on the resected mass and correlated with two DCE-MRI parameters: increase in signal intensity (ISI) and positive enhancement integral (PEI). No correlation could be found between DCE-MRI data acquired immediately after treatment and the percentage of residual tumour. Good correlation coefficients were found for data acquired several days after treatment (ISI, r50.749; PEI, r50.778). However, they were higher when the posttreatment time interval was 7 days or more (ISI, r50.962; PEI, r50.934). These results suggest that a post-treatment delay of 7 days is necessary for the accurate assessment of the presence of residual tumour by DCE-MRI parameters.

Magnetic resonance imaging-guided focused ultrasound surgery (MRIgFUS) is a non-invasive thermotherapeutic method allowing the elimination of different types of lesions without incision, thus reducing the risk of infection and scarring as well as diminishing the pain considerably [1]. MRIgFUS studies have been reported for patients with breast fibroadenomas [2], breast carcinomas [3–6] and uterine fibroids [7, 8]. The ultrasound radiation produced by a transducer elevates the temperature at a focal point and destroys the tissue by protein denaturation and tissue necrosis. The repeated application of the procedure leads to the ablation of multiple points covering the entire lesion or the desired part of the lesion. As the exact location of the focal points may vary depending on the tissue composition through which the focused ultrasound beam passes, monitoring the effects of the focused ultrasound is done using a noninvasive method such as MRI. Once the MRIgFUS is performed, early detection of residual disease is of crucial importance to evaluate the response. Indeed, clinical examination is compounded by necrosis and fibrosis which are often intermixed with residual tumour, breast oedema, or inflammation. Dynamic contrast-enhanced magnetic resonance imaging (DCE-MRI) allows the monitoring of the presence of tumour by demonstrating changes in the enhancement Address correspondence to: Dr Y Boulanger, PhD, De´partement de Radiologie, Hoˆpital Saint-Luc du CHUM, 1058 St-Denis, Montreál, Que´bec, H2X 3J4 Canada. Sources of support: InSightec, Inc. (Haifa, Israel and Dallas, TX) and the Canadian Institutes of Health Research (Ottawa, ON, Canada).

308

Received 7 June 2005 Revised 15 August 2005 Accepted 15 September 2005 DOI: 10.1259/bjr/23046051 ’ 2006 The British Institute of Radiology

curves following the injection of contrast agent [9, 10]. In the breast, this technique has been applied to detect the presence of tumours [11–13], to assess their grade [14, 15] and to monitor the results of different types of treatment [16–18]. Enhancing lesions on post-treatment DCE-MR images are normally indicative of the presence of significant residual tumour, but growing evidence suggests that the situation may be more complicated due to the fact that some benign processes such as oedema, fibrosis, necrosis and inflammation can mimic malignant contrast [19, 20]. Hence, the presence of contrast enhancement alone is not exclusively specific to residual cancer. An enhancing lesion indicates abnormal tissue, but the shape of the enhancement curve must be considered to identify the nature of this tissue. Indeed, inspection of the lesion morphology and enhancement profile shows that most malignancies have an irregular border and display very rapid enhancement and a distinct early washout phase in DCE-MRI. This pattern can be used to distinguish malignant masses from benign lesions or normal tissue which enhance and washout slowly [21]. Since benign phenomena due to the therapy such as necrosis, fibrosis, oedema and inflammation can often occur and be responsible for abnormal enhancement, the time interval between treatment and post-treatment must be taken into account when analysing DCE-MRI data to minimize the contribution of these processes. In the case of conventional treatments such as lumpectomy and conventional surgery, time intervals ranging from 4 weeks to 12 months have been reported to be necessary for a reliable assessment of residual tumour by DCE-MRI [19, 20, 22, 23]. The British Journal of Radiology, April 2006

Effect of post-MRIgFUS delay on DCE-MRI evaluation of residual breast cancer

In our efforts to assess the value of the MRIgFUS technique for the treatment of breast carcinomas, DCEMRI data were collected immediately and several days after treatment, and then compared with the percentage of residual tumour determined by histopathology of the excised treated mass. In our previous report, data acquired several days after treatment were presented for 17 patients which demonstrated a strong correlation between DCEMRI parameters and the percentage of residual tumour [5]. However, for a few patients the correlation was inexplicably poor. In the present study, DCE-MRI data acquired both immediately and several days after treatment were analysed for 25 patients (26 tumours) to determine the effect of the time interval between MRIgFUS and posttreatment MRI data acquisition.

Inclusion criteria were a life expectancy of at least 5 years and the possibility of receiving hormone replacement or Tamoxifen therapy. Major exclusion criteria were pregnancy or lactation, having previously undergone radiation or laser therapy, having received immunosuppressive therapy and having been treated with chemotherapy in the past 3 months. In addition, women with a history of grand mal seizures, severe cerebrovascular disease, haemolytic anaemia, or dialysis treatment; women who were receiving anticoagulation therapy; and women with breast implants were excluded. Patients were also excluded if their lesion was outside the margins of safety for ultrasound ablation, consisted only of microcalcifications, or had extensive intraductal components.

Patients and methods

MRIgFUS

Patients 25 women aged between 45 years and 87 years (mean age¡SD, 61.3 years¡11.0) fulfilling the inclusion/exclusion criteria completed the protocol between March 2000 and June 2004. Each patient was diagnosed by core breast biopsy with an invasive breast carcinoma of less than 3.5 cm in diameter (Grade 1). The tumour volumes determined by MRI ranged between 0.11 cm3 and 11.2 cm3 (Table 1). All tumour types were invasive ductal carcinomas except for one. In one patient (patient 7), two tumours were treated. All patients signed an informed consent form approved by the Scientific and Ethics Committees of our institution prior to any exam or MRIgFUS.

All MRI data were acquired with the GE 1.5 T MRI unit (Signa; GE Medical Systems, Milwaukee, WI) using a breast coil. The MRIgFUS treatments were performed by different versions of the ExAblate system developed by InSightec Inc. (Haifa, Israel and Dallas, TX). The MRIgFUS system and procedure have been described in detail previously [3]. Briefly, the patient is prone on the focused ultrasound table inside the MRI magnet bore. An acoustic ultrasound beam generated by a multielement phased-array transducer focused on a point that was thermally ablated. The location of the focal point was verified by the thermally sensitive fast spoiled gradientecho MRI sequence [3]. Multiple adjacent focal points covering the entire tumour were treated sequentially to complete the FUS ablation.

Table 1. Lesion and exam parameters for breast tumours treated by MRI-guided focused ultrasound surgery (MRIgFUS) Patient number

Tumour type

Original tumour volume (cm3)

Residual tumour volume (cm3)

Post-treatment delay (days)

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 7 25

Invasive Invasive Invasive Invasive Invasive Invasive Invasive Invasive Invasive Invasive Invasive Invasive Invasive Invasive Invasive Invasive Invasive Invasive Invasive Invasive Invasive Invasive Invasive Invasive Invasive Invasive

9.36 0.57 5.18 8.80 7.80 7.20 0.88 0.78 1.85 0.50 3.42 5.76 9.60 2.89 0.70 0.17 1.30 0.21 4.80 0.11 0.75 1.44 0.28 0.70 11.2 0.41

8.42 0.42 2.59 3.52 3.12 2.16 0.20 0.07 0.18 0.05 0.27 0.29 0.48 0.14 0.035 0.005 0.03 0.002 0 0 0 0 0 0 0 0

7 7 7 7 7 3 7 7 3 7 3 3 7 7 14 3 7 7 7 10 3 7 7 7 7 7

ductal carcinoma ductal carcinoma ductal carcinoma ductal carcinoma ductal carcinoma ductal carcinoma ductal carcinoma ductal carcinoma ductal carcinoma ductal carcinoma ductal carcinoma ductal carcinoma ductal carcinoma ductal carcinoma lobular carcinoma ductal carcinoma ductal carcinoma ductal carcinoma ductal carcinoma ductal carcinoma ductal carcinoma ductal carcinoma ductal carcinoma ductal carcinoma ductal carcinoma ductal carcinoma


309

A Khiat, D Gianfelice, M Amara and Y Boulanger

Post-treatment MRI The MR images used in this study were recorded both immediately and at a follow-up visit 3–14 days after the MRIgFUS but before conventional surgery. Among the 25 treated patients, six had their post-treatment evaluation 3 days after ultrasound ablation, 17 patients 7 days after ultrasound ablation (18 tumours), one patient 10 days and one patient 14 days after ultrasound ablation (Table 1). The timing of the post-treatment examination depended on the availability of the MRI scanner and on the scheduling of the surgery. DCE-MR images were acquired before and after injection of the MR contrast agent gadopentetate dimeglumine (Berlex Canada Inc., Lachine, Canada; 0.1 mmol kg21 body weight) using a three-dimensional fast spoiled gradient echo (FSPGR) sequence with fat saturation (repetition time (TR)56.4 ms; echo time (TE)52.4 ms; preparation time522 ms; flip angle510 ˚, slice thickness54 mm; no intersection gap; 256 6 128 points; number of images532; acquisition time53.2 min).

necrosed tumour were determined. Volumes (V) were determined by measuring the largest tumour dimensions in each axis (i.e. distances a, b, and c) and performing an ellipsoid volume calculation: V5(p/6) abc.

Statistical analysis Correlation coefficients between the DCE-MRI parameters ISI and PEI, and the percentage of residual tumour volume determined by histopathology were determined by a Pearson analysis. Correlations were considered significant at p,0.05 with r values . 0.50. Comparison of the correlation coefficients was performed using a t-test [25]. A receiver operating characteristic (ROC) curve analysis was performed for both DCE-MRI parameters. Statistical analyses were performed using the SPSS software version 9.0.1 (SPSS Inc., Chicago, IL).

Results Analysis of DCE-MRI data To calculate the most enhancing pixel (1 mm2) within the entire tumour in the DCE-MR images, a utility program was written using MATLAB programming (Mathworks, Natick, MA) operating on Linux software. From a spherical ROI centred on the original tumour centre and covering the entire tumour region in the intensity vs phase image series, a Be´zier curve was constructed by picking the control points (phase) to identify the most enhancing pixel. DCE-MRI parameters at the most enhancing pixel were evaluated by the FUNCTOOL version 2.5.36b software program (GE Medical Systems, Fremont, CA) from time-signal intensity curves. The dynamic contrast uptake at the most enhancing point was evaluated by calculating two parameters: percentage increase in signal intensity (ISI) and the positive enhancement integral (PEI). ISI corresponds to the percentage of the maximal increase in signal intensity after the injection of contrast agent (SIpost) relative to the pre-contrast signal intensity (SIpre): ISI5(SIpost2SIpre)/SIpre 6 100. PEI is the integral of the area under the enhancement curve observed after the injection of contrast agent for the time-signal intensity graph acquired for a time t: t P PEI~ SIt . Data were analysed separately for all patients 0

taking into account the time interval between the MRIgFUS and the DCE-MRI exam.

Surgery and histopathological evaluation Patients underwent a routine segmental tumour resection 3–21 days after MRIgFUS, which included resection of wide margins around the area treated by focused ultrasound. Three-dimensional macroscopic and microscopic histopathological measurements were performed on the resected mass by a pathologist immediately after surgery [3, 24]. The total tumour volume, the volume of tumour in the treated zone and the volume of 310

Processing of DCE-MRI data Our home written program allowed rapid and automatic determination of the most enhancing point in the DCE-MR images acquired at the post-treatment exam. This method was found to be reliable and exactly reproducible on repeated measurements.

Histopathological data Histopathology results of the excised mass following the MRIgFUS of breast carcinomas show a variable degree of success, as summarized in Figure 1 and Table 1. Among the 26 treated tumours, seven tumours showed no detectable residual cancer at the site of MRIgFUS, 11 tumours had residual cancer below 10%, and seven tumours showed a larger percentage of residual cancer, which ranged between 20% and 90%.

DCE-MRI results In general, complete tumour necrosis should be reflected by the absence of enhancement in the timesignal intensity DCE-MRI curves. Figure 2 presents two types of time-signal intensity curves before, immediately after and several days after MRIgFUS observed for patients 19 (Figure 2a–c) and 21 (Figure 2d–f). Before treatment, a strong enhancement is observed in all cases at the most enhancing point of the breast tumour region (Figure 2a,d). In the case of patient 19, the tumour necrosis determined by histopathology was complete (Figure 1, Table 1) and no apparent enhancement was observed 7 days after treatment (Figure 2c). However, this assessment could not have been made immediately after treatment since a slowly enhancing curve was observed (Figure 2b). When residual cancer was present, as was the case for patient 21, a significant enhancement (Figure 2f) was observed with a steep washout pattern similar to the patterns observed before treatment The British Journal of Radiology, April 2006

Effect of post-MRIgFUS delay on DCE-MRI evaluation of residual breast cancer

(Figure 2d). In that case, the enhancement patterns were very similar immediately and 3 days after MRIgFUS (Figure 2e,f).

Correlation between DCE-MRI and histopathology data

Figure 1. Graph showing the percentage of residual tumour determined by histopathology following MRI-guided focused ultrasound surgery (MRIgFUS) for 26 breast tumours in 25 patients. A variable degree of success was achieved at various stages of development of the technique.

Figure 3a,b presents the correlations between the ISI parameter and the percentage of residual tumour volume determined by histopathology for all tumours and for the 20 tumours evaluated at least 7 days after the MRIgFUS, respectively. Similar correlations are presented in Figure 3c,d between the PEI parameter and the percentage of residual tumour volume determined by histopathology for all patients and for the patients having been evaluated at least 7 days post-treatment, respectively. Strong correlations between the DCE-MRI parameters ISI (r50.749) and PEI (r50.778), and the percentage of residual tumour volume were found. When six patients who had their evaluation only 3 days after MRIgFUS were excluded from the analysis, the correlations were much stronger, the correlation coefficients for ISI increasing to 0.962 (Figure 3b) and for PEI to 0.934 (Figure 3d). Using a t-test to compare the

Figure 2. Dynamic contrast-enhanced magnetic resonance imaging (DCE-MRI) signal intensity curves as a function of image number (6.0 s/image) at the maximally enhancing pixel for patient 19 (a) before, (b) immediately after and (c) 7 days after MRIguided focused ultrasound surgery (MRIgFUS), and for patient 21 (d) before, (e) immediately after and (d) 3 days after MRIgFUS. The absence of signal enhancement in (c) is consistent with the absence of residual tumour and the progressive signal enhancement in (b), (e) and (f) is attributable to benign tissue phenomena occurring in the first few days following focused ultrasound ablation. The British Journal of Radiology, April 2006

311


Figure 3. Graphs showing the correlation between parameters calculated from the post-treatment dynamic contrast-enhanced magnetic resonance imaging (DCE-MRI) data and the percentage of residual viable cancer determined by histopathology on the excised mass for breast tumours treated by MRI-guided focused ultrasound surgery (MRIgFUS): increase in signal intensity (ISI) for (a) all patients (r50.749) and (b) patients who had their post-treatment evaluation at least 7 days after treatment (r50.962); positive enhancement integral (PEI) for (c) all patients (r50.778), and (d) patients who had their post-treatment evaluation at least 7 days after treatment (r50.934). Points corresponding to patients whose post-treatment was performed 3 days after MRIgFUS are shown as asterisks (*) in (a) and (c).

correlation coefficients for the ISI and PEI parameters for all values and for values obtained more than 7 days after treatment [25], a statistically significant difference was found for ISI (p,0.01) and an almost significant difference for PEI (p50.055). Due to the presence of enhancement in practically all cases, no correlation could be found between the ISI and PEI parameters and residual tumour volume for data acquired immediately after treatment.

ROC curve analysis Receiver operating characteristic (ROC) curves are presented in Figure 4 for the ISI parameter. When the data for all tumours were used, the area under the curve was 0.853 with a p-value of 0.005 for asymptotic significance (Figure 4a). The same analysis was made for the 20 tumours whose DCE-MRI data were recorded 7 days or more after treatment, which yielded an area under the curve of 0.986 with a p-value of 0.001 for asymptotic significance (Figure 4b). The cut off ISI value for both curves corresponded to 7% with a sensitivity of 100% and a specificity of 83%. The ROC curve analysis for the PEI parameter showed an area under the curve of 0.732 without asymptotic significance (p50.067) for all tumours, and an area under the curve of 0.785 without asymptotic significance (p50.055) for patients whose DCE-MRI exam was performed 7 days or more after treatment. The cut off value for PEI was 4% with a sensitivity of 67% and a specificity of 83%.

Discussion The calculation of the ISI and PEI parameters from DCE-MR images requires an accurate determination of 312

the most enhancing point in the tumour area. The strategy of using the automatically identified most enhancing pixel was based on the conclusions of previous work clearly demonstrating improved results with this strategy compared with using large regions of interest [26]. This was achieved by using a home written program which allowed rapid identification of this point, thus significantly reducing the time for analysis compared with the manual location of this point used in our previous report [5]. This study has also the advantage of using volumes for histological quantitation of tissue. In previous studies, diameter values were utilized. As tumour and non-tumour masses are generally not spherical, the volume calculation leads to increased accuracy. The analysis of DCE-MRI data recorded immediately after performing MRIgFUS showed enhancement patterns in practically all cases and no correlation could be found between the ISI or PEI parameters and the percentage of residual tumour determined by histopathology. These enhancements are attributable to benign phenomena such as haemorrhage, fibrosis, inflammation or oedema caused by the treatment and which prevent accurate assessment of the percentage of residual tumour by DCE-MRI. Analysis of the post-treatment DCE-MRI data acquired 3–14 days after MRIgFUS of small breast tumours revealed that the ISI and PEI parameters correlate strongly with the percentage of residual tumour determined by histopathology. The correlation coefficients for the two parameters are very similar (0.749 for ISI and 0.778 for PEI) based on data from 26 treated tumours whose volume of residual tumour determined by histopathology varied between 0 and 90%. When data for six tumours analysed less than 7 days after treatment were not considered, the correlation coefficients The British Journal of Radiology, April 2006

Effect of post-MRIgFUS delay on DCE-MRI evaluation of residual breast cancer Figure 4. Receiver operating characteristic curves for the increase in signal intensity (ISI) parameter calculated from the dynamic contrast-enhanced magnetic resonance imaging (DCE-MRI) data for (a) 26 tumours analysed 3–14 days after the MRI-guided focused ultrasound surgery (MRIgFUS) and (b) 20 tumours analysed 7–14 days after the MRIgFUS. The areas under the curves were (a) 0.853 and (b) 0.986 and the p-values for asymptotic significance were (a) 0.005 and (b) 0.001. A significant improvement is obtained when only data for patients examined 7 days or more after MRIgFUS are considered.

significantly improved, reaching 0.962 for ISI and 0.934 for PEI. These results stress the importance of the time interval between MRIgFUS therapy and post-treatment evaluation. In previous reports, the importance of the time interval between treatment and DCE-MRI was discussed for invasive treatments of breast lesions [19, 20, 22, 23]. In a study of 68 patients whose breast lesions were treated by lumpectomy, an interval of 28 days after treatment was recommended before the DCE-MRI examination [23]. When conventional surgery was performed, time intervals of 9–12 months were necessary to obtain reliable data [19, 20, 22]. As conventional surgical treatments cause a much larger perturbation of the tissue, the fact that the interval necessary is much longer than for MRIgFUS is consistent with the difference in severity between the two procedures. For some patients the correlation between DCE-MRI parameters and percentage of residual tumour volume was especially poor, such as for patient 21 whose timesignal intensity curves displayed abnormally high enhancement leading the ISI and PEI parameters to show as outliers on the correlation curves with the percentage of residual tumour. This unexpected enhancement is attributable to the fact that residual tissue is not residual tumour, but rather consists of enhancing benign tissue such as inflammation, necrosis, fibrosis or oedema at the site of MRIgFUS. It is known that the post-treatment contrast enhancement resulting from such phenomena limits the positive predictive value and specificity for the diagnosis of residual disease [23, 27]. In our case, two arguments support this explanation. The first argument is that patient 21 was evaluated only 3 days after the MRIgFUS. When the six patients who had their evaluation only 3 days after treatment were excluded from the analysis, the results were more reliable as assessed by higher correlation coefficients and larger areas under the ROC curves. The second argument in favour of non-tumour tissue for patient 21 pertains to the shape of the time-signal intensity curve at the post-treatment examination. The pattern of this curve does not correspond to the expected curve profile of a malignant tissue [21]; it displays a slow rate of enhancement in comparison with the time-signal intensity enhancement curve at the pre-treatment exam and it continues to rise over the course of the examination. Malignancies typically display an irregular border at the histological analysis; they enhance very rapidly The British Journal of Radiology, April 2006

and display a distinct early washout phase in the DCEMRI curves. For breast cancers, Figure 2a,c represents a typical pattern. Previous studies have concluded that the presence of contrast enhancement alone is not specific for residual cancer, and that the shape of the enhancement curve is a very important factor to consider [21, 28, 29]. Additional consideration of the morphology and internal architecture of the excised mass and a critical analysis of the time-signal intensity curves can be used to help distinguish residual tumours from benign processes. In a previous publication [3], histological slides were presented for a patient analysed 3 days after MRIgFUS which clearly show these phenomena. Obviously, these benign processes depend on the sensitivity of each patient, and are strongly influenced by the time interval between MRIgFUS and post-treatment evaluation. Consequently, the accuracy of DCE-MRI for predicting residual disease is affected by these processes. Comparison of the ROC curves for all patients and for the patients examined 7 days and more after MRIgFUS shows a significant improvement with an area under the curve of the ISI parameter increasing from 0.853 to an almost perfect 0.986 for the latter group. The same trend was observed for the PEI parameters although lower values were obtained. These results confirm the importance of the post-treatment time interval and suggest that the ISI parameter is a better parameter to evaluate treatment outcome. The cut-off value for the determination of residual tumour was around 7% for ISI and 4% for PEI. The main limitation of this study is the low number of patients with a high percentage of residual tumour. Although our group has reported the largest number of patients treated by MRIgFUS, the accuracy of our correlations could be improved if more patient data were available. Most of our patients showed less than 10% residual tumour. At low percentages, the accuracy of the technique is reduced. This situation is unlikely to improve as the treatment technique is increasingly successful. Nonetheless, our results demonstrate a high degree of specificity and sensitivity for the DCE-MRI parameters. In summary, our results demonstrate that DCE-MRI parameters such as ISI and PEI are both useful in determining the effectiveness of MRIgFUS in breast carcinomas, provided the measurements are performed sufficiently long after treatment, i.e. approximately 313


7 days, based on our data. Benign phenomena caused by the treatment are responsible for abnormal enhancement in the first few days. The accuracy of our analysis could be improved by increasing the sample size and by evaluating patients longitudinally at different time intervals after MRIgFUS.

13.

14.

Acknowledgments The authors thank Mr Martin Ruel, BSc, Dr Assia Belblidia, MD, and the radiology technologists at Hoˆpital Saint-Luc du CHUM for data acquisition. They are grateful to Dr Sharon Thomsen, MD, for assistance with histopathological analyses. The financial support from InSightec, Inc., Haifa, Israel and Dallas, TX and from the Canadian Institutes of Health Research, Ottawa, Canada is acknowledged.

References 1. Moonen CTW, Quesson B, Salomir R, Vimeux FC, de Zwart JA, van Vaals JJ, et al. Thermal therapies in interventional MR imaging - focused ultrasound. Neuroimaging Clin N Am 2001;11:737–47. 2. Hynynen K, Pomeroy O, Smith DN, Huber PE, McDannold NJ, Kettenbach J, et al. MR imaging-guided focused ultrasound surgery of fibroadenomas in the breast: a feasibility study. Radiology 2001;219:176–85. 3. Gianfelice D, Khiat A, Amara M, Belblidia A, Boulanger Y. MR imaging-guided focused US ablation of breast cancer: histopathologic assessment of effectiveness – initial experience. Radiology 2003;227:849–55. 4. Gianfelice D, Khiat A, Amara M, Belblidia A, Boulanger Y. Feasibility of magnetic resonance imaging-guided focused ultrasound surgery as an adjunct to tamoxifen therapy in high-risk surgical patients with breast carcinoma. J Vasc Interv Radiol 2003;14:1275–82. 5. Gianfelice D, Khiat A, Amara M, Belblidia A, Boulanger Y. MR imaging-guided focused ultrasound surgery of breast cancer: correlation of dynamic contrast-enhanced MRI with histopathologic findings. Breast Cancer Res Treat 2003;82:93–101. 6. Huber PE, Jenne JW, Rastert R, Simiantonakis, I, Sinn H-P, Strittmatter H-J, et al. A new noninvasive approach in breast cancer therapy using magnetic resonance imagingguided focused ultrasound surgery. Cancer Res 2001;61:8441–7. 7. Stewart EA, Gedroyc WM, Tempany CM, Quade BJ, Inbar Y, Ehrenstein T, et al. Focused ultrasound treatment of uterine fibroid tumors: safety and feasibility of a noninvasive thermoablative technique. Am J Obstet Gynecol 2003;189:48–54. 8. Tempany CM, Stewart EA, McDannold N, Quade BJ, Jolesz FA, Hynynen K. MR imaging-guided focused ultrasound surgery of uterine leiomyomas: a feasibility study. Radiology 2003;226:897–905. 9. Padhani AR, Husband JE. Dynamic contrast-enhanced MRI studies in oncology with an emphasis on quantification, validation and human studies. Clin Radiol 2001;56:607–20. 10. Hayes C, Padhani AR, Leach MO. Assessing changes in tumour vascular function using dynamic contrast-enhanced magnetic resonance imaging. NMR Biomed 2002;15:154–63. 11. Knopp MV, Weiss E, Sinn HP, Mattern J, Junkermann H, Radeleff J, et al. Pathophysiologic basis of contrast enhancement in breast tumors. J Magn Reson Imaging 1999;10:260–6. 12. Hulka CA, Edmister WB, Smith BL, Tan L, Sgroi DC, Campbell T, et al. Dynamic echo-planar imaging of the

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breast: experience in diagnosing breast carcinoma and correlation with tumor angiogenesis. Radiology 1997;205:837–42. Gibbs P, Liney GP, Lowry M, Kneeshaw PJ, Turnbull LW. Differentiation of benign and malignant sub-1 cm breast lesions using dynamic contrast enhanced MRI. Breast 2004;13:115–21. Daldrup H, Shames DM, Wendland M, Okuhata Y, Link TM, Rosenau W, et al. Correlation of dynamic contrastenhanced MR imaging with histologic tumor grade: comparison of macromolecular and small-molecular contrast media. AJR Am J Roentgenol 1998;171:941–9. Rydland J, Bjønerud A, Haugen O, Torheim G, Torres C, Kvistad KA, et al. New intravascular contrast agent applied to dynamic contrast enhanced MR imaging of human breast cancer. Acta Radiol 2003;44:275–83. Tsuboi N, Ogawa Y, Inomata T, Yoshida D, Yoshida S, Moriki T, et al. Changes in the findings of dynamic MRI by preoperative CAF chemotherapy for patients with breast cancer of stage II and III: pathologic correlation. Oncol Rep 1999;6:727–32. Drew PJ, Kerin MJ, Mahapatra T, Malone C, Monson JR, Turnbull LW, et al. Evaluation of response to neoadjuvant chemoradiotherapy for locally advanced breast cancer with dynamic contrast-enhanced MRI of the breast. Eur J Surg Oncol 2001;27:617–20. Martincich L, Montermurro F, De Rosa G, Marra V, Ponzone R, Cirillo S, et al. Monitoring response to primary chemotherapy in breast cancer using dynamic contrastenhanced magnetic resonance imaging. Breast Cancer Res Treat 2004;83:67–76. Soderstrom CE, Harms SE, Farrell RS Jr, Pruneda JM, Flamig DP. Detection with MR imaging of residual tumor in the breast soon after surgery. AJR Am J Roentgenol 1997;168:485–8. Viehweg P, Heinig A, Lampe D, Buchmann J, HeywangKo¨brunner SH. Retrospective analysis for evaluation of the value of contrast-enhanced MRI in patients treated with breast conservative therapy. MAGMA 1998;7:141–52. Kuhl CK, Mielcareck P, Klaschik S, Leutner C, Wardelmann E, Gieseke J, et al. Dynamic breast MR imaging: are signal intensity time course data useful for differential diagnosis of enhancing lesions? Radiology 1999;211:101–10. Heywang-Ko¨brunner SH, Schlegel A, Beck R, Wendt T, Kellner W, Lommatzsch B, et al. Contrast-enhanced MRI of the breast after limited surgery and radiation therapy. J Comput Assist Tomogr 1993;17:891–900. Frei KA, Kinkel K, Bonel HM, Lu Y, Esserman LJ, Hylton NM. MR imaging of the breast in patients with positive margins after lumpectomy: influence of the time interval between lumpectomy and MR imaging. AJR Am J Roentgenol 2000;175:1577–84. Thomsen S. Mapping of thermal injury in biologic tissues using quantitative pathologic techniques. Proc SPIE 1999;3594:82–95. Sachs L. Applied statistics: a handbook of techniques. 2nd edn. New York, NY: Springer-Verlag, 1984. Liney GP, Gibbs P, Hayes C, Leach MO, Turnbull LW. Dynamic contrast-enhanced MRI in the differentiation of breast tumors: user-defined versus semi-automated regionof-interest analysis. J Magn Reson Imaging 1999;10:945–9. Orel SG, Reynolds C, Schnall MD, Solin LJ, Fraker KL, Sullivan DC. Breast carcinoma: MR imaging before reexcisional biopsy. Radiology 1997;205:429–36. Orel SG, Schnall MD, LiVolsi VA, Troupin RH. Suspicious breast lesions: MR imaging with radiologic-pathologic correlation. Radiology 1994;190:485–93. Heywang-Ko¨brunner SH. Contrast-enhanced M.R.I. of the breast. 2nd edn. Berlin, Heidelberg: Springer-Verlag, 1995.



Introduction of grids to mobile ICU radiography in a teaching hospital D W ANDERSON,

MBBS (hons), FRANZCR

Monash Medical Centre, 246 Clayton Road, Clayton, Melbourne, Victoria, Australia 3168 ABSTRACT. The purpose of this study was to review the change in image quality before and after introducing grid use routinely to our mobile X-ray service. This was studied in the intensive care unit (ICU) setting, comparing images obtained over a 2 week period prior to and after the introduction of the change in technique. We introduced a 6:1 grid with appropriate changes in exposure factors. No other alterations were made. There were 133 patients in the preliminary group and 196 patients in the post-grid group. We found a reduction in the proportion of images that were of non-diagnostic or barely diagnostic quality. Non-diagnostic examinations were reduced from 18% to 1%. Introducing grids to our mobile service resulted in improvement in image diagnostic quality, largely by reducing the proportion of poor and unacceptable quality images. This effect does not appear to have been documented in the literature.

Received 18 March 2005 Revised 7 September 2005 Accepted 15 September 2005 DOI: 10.1259/bjr/66912494 ’ 2006 The British Institute of Radiology

Monash Medical Centre (MMC) is a large teaching hospital in suburban Melbourne, Victoria. We have recently introduced grids to our mobile ward service. Data were collected before and after the introduction of grids to assess their impact on the diagnostic quality of images obtained on the daily morning intensive care unit (ICU) film round.

N

Methods

Also contributing to image grading was a subjective assessment of lung parenchymal detail, both the ability to satisfactorily characterize the nature and extent of infiltrates as well as the ability to resolve normal parenchymal architecture in areas of non-pathological lung. The components of our scoring system are similar to those used by others in the assessment of portable radiographs [1–3]. Images were graded independently by two experienced radiologists, each with more than 2 years experience with the display system used. Images were modified for window settings by each viewer to their satisfaction with active changes made during review of each image to fully interrogate each image to the satisfaction of the reviewers. Differences in grade were resolved by consensus.

All patients examined during the morning ICU X-ray round were collected for a 2 week period prior to the introduction of a grid technique. Images were obtained using GE AMX4 mobile machine and Agfa CR system. 1 month following the changed technique, data were again collected for a 2 week period. Images were obtained using a 6:1 grid landscape or portrait grid with GE AMX4 mobile machine and Agfa CR system. Both sets of images were viewed on BARCO monitors. Data for the 4 week period following the introduction of the new technique were not collected to allow staff to become familiar with changed techniques, particularly exposure alterations. This was thought desirable for repeatable and reliable patient examinations. The images were graded for diagnostic quality using a five point grading system as follows: (1) (2) (3) (4) (5)

N N N N

Inability to follow the course of central lines, catheters, drain tubes and other medical devices. Inability to identify the tip location of the above lines, catheters and devices. Respiratory or other motion artefact of severe nature. Grid cut-off or other grid artefact of severe nature. Non-inclusion of the complete extent of the lung parenchyma.

not of diagnostic quality poor, barely adequate diagnostic quality fair, acceptable diagnostic quality good, above average diagnostic quality excellent diagnostic quality

Images were determined ‘‘not of diagnostic quality’’ if any of the following were present:

N

Inability to see the thoracic spine through the cardiac shadow.


Figure 1. Comparison chest radiograph (CXR) pre- and post-grid.

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D W Anderson

A blinded interpretation of pre- and post-grid images was not attempted as the use of a grid results in discernable grid lines, rendering overt identification of technique used.

Results The initial image set performed before grid technique (total 133 patients) produced the following results: Grade Grade Grade Grade Grade

1: 2: 3: 4: 5:

24/133 (18%) 55/133 (41%) 40/133 (30%) 13/133 (10%) 1/133 (1%)

The second image set, performed following the introduction of grid technique (total 196 patients) produced the following results: Grade Grade Grade Grade Grade

1: 2: 3: 4: 5:

Figure 3. Same patient as Figure 2 with grid.

2/196 (1%) 30/196 (15%) 104/196 (53%) 48/196 (25%) 12/196 (6%)

The difference in group size was the result of differences in patient load in the ICU for the morning radiology service. Figures 2–5 are comparative chest X-rays performed on the same day with and without grids on two separate patients.

Discussion The initial data set produced a somewhat surprising result in that nearly one in five images (18%) was not of diagnostic quality. This was largely the result of not being able to identify the position or tip of various devices. Swan-ganz catheters and fine bore central lines

Figure 4. Chest radiograph no grid.

Figure 2. Chest radiograph no grid.

Figure 5. Same patient as Figure 4 with grid.

316


Introduction of grids to mobile ICU radiography

were most often affected. There was also a contribution in many images from difficulties in reliably characterizing infiltrates. This was most notable in larger patients. In combination with barely acceptable images (category 2, 41%), our ICU mobile service was producing images of poor to unacceptable quality in nearly six out of 10 patients. Portable images had been acknowledged as of lesser quality than imaging performed in department. This had been considered a largely inevitable outcome given the compromises involved in portable imaging. The desire to investigate the impact of grids was prompted by their effect when used on a very large patient in the ICU in attempts to produce a diagnostic image. That grid use can result in a significant improvement in signal to noise ratio (SNR) at the cost of increased dose is well recognized [4–6]. Our service had not conducted a formal review of image diagnostic quality in any systematic fashion for some years, instead relying on an ad-hoc notification and correction of specific problems as they arose. Review had not been performed since the introduction of our filmless screen based CR reporting system. As such, we do not have a prior benchmark for comparison with these results. The data set following introduction of grid technique shows a significant reduction in the proportion of images deemed not of diagnostic quality (18%.1%). There was also a significant reduction in category two images (41%.15%). As a result of these improvements, 78% of images were deemed of acceptable or above average diagnostic quality (53% category 3 + 25% category 4). Review of the literature shows little has been published quantifying the change in diagnostic viewing quality of mobile ICU images performed with and without grids. Two authors have stated that portable chest radiography is significantly improved with use of a 6:1 or 8:1 grid [7, 8], but without comparative evidence. Introduction of grids has also introduced some new problems. All images contain discernable grid lines. Whilst visible, these were not thought to impede satisfactory image interpretation. Grid ‘‘cut-off’’ producing asymmetrical density was seen on approximately 30% of films. This was of variable severity but again did not appear to significantly affect image interpretability. No image was deemed category 1 as a result of grid cutoff. The finding that grid cut-off was not a significant problem is in keeping with the work of Grunert et al and Ciccotosto et al [2, 5, 9]. A grid technique results in an increased radiation dose to the patient. Rill et al stated that use of a grid resulted in a 400–600% increase in dose to their phantom for a 60– 300% improvement in SNR [4]. Increasing tube potential from 80 kVp to 100 kVp resulted in further 10% increases in SNR and phantom dose. Similar dose increases and decreases in scatter radiation were found by Floyd et al [6]. Balanced against this is the critical status of the typical ICU patient with a requirement for the most accurate diagnostic information achievable. Most patients in our ICU are older adults with multiple co-morbidities. As such, the increased dose is acceptable. No attempt was made to assess the clinical relevance of changes to subjective image quality in this study as this lay outside our scope. The British Journal of Radiology, April 2006

The intent of this article is to provide a small measure of our experience with introducing grids to our ICU mobile service and their impact on subjective radiologist assessed image quality, and the following limitations are acknowledged: (1) (2) (3)

(4)

(5)

(6)

(7)

The sample sizes are relatively small No attempt was made to correlate image quality with patient size (body mass index, BMI) No attempt was made to determine the effect of radiographer experience on image quality. MMC is a teaching hospital with staff of significantly varying experience No attempt has been made to determine if the subjective improvement in image quality has had any relevant effect on clinical decision making or patient outcome The study is unable to be performed in a blinded fashion due to image effects resulting from the intervention (grid use) The differences in sample size (patient load), may reflect further unknown confounding factors in the patients with independent effect on image quality. We made no attempt at randomization The introduction of a change in technique may have resulted in a change in the diligence of radiography staff regarding positioning, exposures and other factors

Optimizing image diagnostic quality is relevant in the ICU setting as other authors have found unexpected findings in 37–65% of patients [10–13]. Ekenmeyer et al stated changes were initiated in diagnostic approach or therapy as a result of CXR findings alone in 27% of patients [11]. Malpositioned ETT, central lines or catheters have been seen in 9–20% of patients [11, 13]. An American study from 1994 showed 12% of hospitals using grids for mobile chest radiography. It was not stated if this was routine or on a patient by patient basis [14]. Review of technique used by the other large public hospitals in our city found only one of four institutions were using grids for mobile ICU radiography and this was reserved for very large patients only. Reasons given included inconvenience and historical practice.

Conclusion Presented is our experience with introducing grids to our ICU mobile service. We found an improvement in overall image quality, largely as a result of reduced numbers of poor and unacceptable quality images. This effect has not been documented in the literature to our knowledge. From this preliminary work, further directions include a randomized study to confirm the positive impact of grids on image quality and an assessment of the impact of improved image quality with grid technique on clinical decision making in the ICU setting.

References 1. Langevin PB, Hellein V, Harms SM, Tharp C, CheungSeekit C, Lampotang S. Synchronization of radiograph film

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2.

3.

4.

5.

6.

7.

exposure with the inspiratory pause. Am J Respir Crit Care Med 1999;160:2067–71. Ciccotosto C, Storto ML, Guidotti A, Ferrante R, Bonomo L. Bedside thoracic radiography: a comparison between 3 different types of grid. Radiol Med (Torino) 1994;87:127–33. Rottenberg GT, Chin RJ, Allen CM, Maggs P, Kessel D. Portable chest radiology in intensive care: a comparison of a new dual characteristic film-screen system (Insight) incorporating a flexible grid with a standard film-screen system. Clin Radiol 1996;51:494–8. Rill LN, Brateman L, Arreola M. Evaluating radiographic parameters for mobile chest computed radiography: phantoms, image quality and effective dose. Med Phys 2003;30:2727–35. Grunert JH, Boy B, Groenwold SK, Krahn-Peters V, Hendrickx P. Use of multi-grid screens in intensive care units. Aktuelle Radiol 1995;5:293–6. Floyd CE Jr, Chotas HG, Ravin CE. Scatter-reduction characteristics of an infinity-focussed gridded radiographic cassette. Invest Radiol 1994;29:852–5. BHR 1980; HHS Publ 80–8124: 202–6.

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8. Rossi RP, Harnisch BD, Hendee WR. Evaluation of an automatic exposure control device for mobile radiography. Radiology 1982;145:823–7. 9. Grunert JH, Boy B, Busche D, Groenewold SK, Herrmann H, Krahn-Peters V, et al. Grids and high kilo-volt-peaksetting in bedside chest radiographic examinations. JBRBTR 2000;83:296–9. 10. Janower ML, Jennas-Nocera Z, Mukai J. Utility and efficacy of portable chest radiographs. AJR Am J Roentgenol 1984;142:265–7. 11. Ekenmeyer WB, Crapo RO, Calhoon S, Cannon CY, Clayton PD. Efficacy of chest radiography in a respiratory intensive care unit. A prospective study. Chest 1985;88:691–6. 12. Greenbaum DM, Marschall KE. The value of daily chest xrays in intubated patients in the medical intensive care unit. Crit Care Med 1982;10:29–30. 13. Henschke CI, Pasternak GS, Schroder S, Hart KK, Herman PG. Bedside chest radiography: diagnostic efficacy. Radiology 1983;149:23–6. 14. Wandtke JC. Bedside chest radiography. Radiology 1994;190:1–10.



Colour Doppler ultrasound patterns and clinical follow-up of incidentally found hypoechoic, vascular tumours of the spleen: evidence for a benign tumour 1

¨ RG, C GO

MD,

1

¨ RG, K GO

MD,

1

T BERT,

MD

and 2P BARTH,

MD

1

Medizinische Universita¨tsklinik, Baldingerstraße, 35043 Marburg/Lahn, Germany and 2Institut fu¨r Pathologie der Universita¨tsklinik, Baldingerstraße, 35043 Marburg/Lahn, Germany

ABSTRACT. Between January 1990 and January 2005, incidental hypoechoic, vascular tumours of the spleen were identified in 13 patients using B-mode and colour Doppler ultrasound (CDS). All lesions found were well demarcated, intrasplenically located, and ranged in size between 1 cm and 4 cm. The increased vascular pattern on CDS was confirmed in 9 of the 13 cases by contrast enhanced ultrasound (CES), while two patients showed reduced vascularity on CES. In 10 patients, lesions were confirmed by contrast enhanced CT. Histological examination was performed in three patients with the diagnosis of capillary haemangioma (n52) and hamartoma (n51). In the remaining cases, ultrasound follow-up was performed (range 4 months to 13 years) and demonstrated no evidence of tumour growth in all but one patient. During a 4 year follow-up, one lesion increased in size from 1.0 cm to 1.5 cm and in the same patient an additional 0.5 cm sized hypoechoic increased vascular lesion was also found. In the spleen a hypoechoic lesion with an increased vascular pattern incidentally found by ultrasound most likely indicates a benign tumour with capillary haemangioma/ hamartoma as the most likely diagnosis. However, it should be emphasised that in all cases a careful ultrasound follow-up is warranted.

In recent years, B-mode ultrasound patterns of various common splenic lesions including infarction [1, 2] lymphoma [3], rupture [4, 5], metastases [6], cysts [7], haemangioma [8] and other tumours [9] have been described. The poor correlation between ultrasound appearances and specific splenic pathology causes difficulty in clinical diagnosis. In liver lesions, colour Doppler ultrasound (CDS) as well as contrast enhanced ultrasound (CES) have improved diagnostic accuracy [10]. To date there are few data regarding CDS and CES patterns of splenic lesions [11–14]. This report reviews the spectrum of ultrasound findings and the clinical follow-up we have observed in 13 patients with incidentally found hypoechoic vascular splenic lesions.

Patients and methods Between January 1990 and January 2005, 13 patients were found incidentally to have focal hypoechoic splenic lesions with an increased vascular pattern on CDS at an internal medicine centre. The patient ages ranged from 2 years to 76 years. The male–female ratio was 8:5. Ultrasound and clinical data of these 13 study patients were retrospectively evaluated. The inclusion criteria for the study were:

N

Incidentally found focal splenic lesion without clinical symptoms or evidence of metastasised malignant diseases.


N N

Received 1 February 2005 Revised 11 June 2005 Accepted 17 August 2005 DOI: 10.1259/bjr/81529894 ’ 2006 The British Institute of Radiology

Hypoechoic splenic lesion using the surrounding splenic tissue as an ‘‘in vivo’’ reference. Increased vascularity on colour-Doppler ultrasound using the surrounding splenic vascularity as an ‘‘in vivo’’ reference.

The following B-mode and CDS ultrasound parameters were retrospectively evaluated: splenic size, sonomorphological pattern of the parenchymal lesions (homogeneous vs inhomogeneous vs central echoic), number (solitary vs multiple), configuration (round vs oval), margin (smooth vs irregular), maximal size of lesions and CDS patterns of vessels (diffuse vs radial). Between 1990 and 1998, patients were examined by Acuson 128 XP (Acuson, Mountain View, CA). From 1998, all new study patients and follow up examinations were performed using an Acuson Sequoia (Acuson; Siemens Medical Solutions, Mountain View, CA) equipped with a 3.5 MHz curvilinear array and 8 MHz linear array transducer. All ultrasound studies were performed by a single observer CG with more than 20 years of experience. Colour Doppler settings were optimized to achieve the greatest sensitivity for allowing detection of low flow. Since January 2004 all new study-patients and all follow-up study-patients were (re)examined by CES with a second generation contrast agent (SonoVueH; Bracco SpA, Milan, Italy). Informed consent according to legislation was obtained in each patient for CES examination and statement of the local internal review board was achieved. 319

320

US, ultrasound, CDS, colour Doppler ultrasound; CT, computed tomography; MRI, magnetic resonance imaging; CES, contrast enhanced ultrasound.

CT CT – CT MRI CT CT CT CT hypoechoic – isoechoic hypoechoic – hypoechoic hypoechoic hypoechoic hypoechoic hyperechoic, central filling – isoechoic, central filling hyperechoic, peripheral filling – hyperechoic, peripheral filling hyperechoic, peripheral filling hyperechoic, peripheral filling hyperechoic, peripheral filling diffuse radial diffuse radial diffuse radial radial radial radial solitary, central echoic solitary, central echoic solitary, homogeneous solitary, homogeneous solitary, homogeneous multiple, central echoic solitary, central echoic solitary, central echoic multiple, central echoic 5 (Figure 2) 6 7 8 9 (Figure 1) 10 (Figure 3) 11 12 13

1 1 1 1 4 3, 2.5 2 1.5 1.5, 0.5

CT CT – CT isoechoic – – hypoechoic isoechoic, central filling – – hyperechoic, peripheral filling radial radial diffuse radial 1.5 1 2 1 homogeneous inhomogeneous inhomogeneous central echoic solitary, solitary, solitary, solitary,

CDS Size (cm) B-mode ultrasound Patient

Table 1. Clinical and ultrasound findings in high vascular splenic tumours

The ultrasound data of the study patients are shown in detail in Table 1. In all patients, splenic size was normal (less than 11 cm 6 5 cm). Regarding echotexture, parenchymal lesions were homogeneous (n54), inhomogeneous (n52), and central echoic (n57) (star-like pattern). Lesions were solitary (n511), multiple (n52), round (n513) and with a smooth margin (n513). Two patients had two splenic lesions on primary diagnosis. Size of lesions ranged from 1 cm to 4 cm (Figures 1–3). CDS of increased vascular lesions showed a diffuse vessel pattern (Figures 1 and 2) or a radial (‘‘starfishlike’’) vessel pattern in the remaining 9 patients (Figure 3). In two patients (Nos. 2 and 9) surgical diagnosis was performed (one splenectomy and one tumour resection) and confirmed capillary haemangioma. In patient 1, ultrasound-guided tumour biopsy was performed and the diagnosis of hamartoma was made. All nonsurgically treated patients were subject to ultrasound follow-up investigations which were performed at 3 month intervals during the first year and then once every year (patient 3 was lost to follow-up after 2 years, patients 7 and 11 died). In patient 9, the resection specimen consisted of macroscopically normal splenic tissue surrounding a well circumscribed but unencapsulated red-brown tumour with a maximum diameter of about 3.5 cm. Histologically the lesion showed evenly distributed capillaries with inconspicuous round lumina surrounded by a thick rim of connective tissue (Figure 1c). Immunostaining (standard avidin-biotin-complex peroxidase method) for CD 34 and CD 31 demonstrated the lumina to be outlined by slightly hyperplastic endothelia (Figure 1d). These findings prompted the diagnosis of a sclerosing capillary haemangioma.

CES-arterial

Results

1 2 3 4

CES-parenchymal

Additional imaging

Follow-up/treatment

Following baseline ultrasound, CES studies were performed using an Acuson Sequoia equipped with contrast-specific, continuous-mode software operating at a low acoustic pressure. Immediately after contrast medium injection, the splenic lesions were observed for evidence of contrast uptake over a 5 min period. CES studies were analysed on the basis of review of ultrasound unit-stored clips. CES parameters were determined by only one observer (CG). For characterization of potential interobserver variability, one other completely blinded observer (KG) with at least 20 years of ultrasound experience retrospectively analysed CES data on the basis of review of ultrasound unit-stored images. A third reviewer (TB) decided in cases of disagreement (2 of 18 CES patterns). The following CES patterns of intrasplenic lesions using the normal splenic tissue as an ‘‘in vivo’’ reference were considered: extent of enhancement during the arterial phase between 1–30 s and parenchymal phase between 1 min and 5 min after injection (absent vs hypoechoic vs isoechoic vs hyperechoic vs mixed echogenicity). During the arterial phase, a distinction between a central filling and a peripheral filling was made.

13 years, biopsy 2 months, splenectomy 2 years, constant in size 4 years, increased in size, one more lesion 4 years, constant in size 8 months, constant in size 3 years, constant in size 3 years, constant in size 2 months, tumour resection 1 year, constant in size 4 months, constant in size 6 months, constant in size 6 months, constant in size

C Go¨rg, K Go¨rg, T Bert and P Barth


CDS of vascular tumours of the spleen

(a)

(c)

(b)

(d)

Figure 1. Patient No. 9. (a) B-mode ultrasound shows a well-demarcated solitary homogeneous hypoechoic splenic tumour. (b) Colour-Doppler ultrasound shows a tumour with marked diffuse intratumoral flow signals. (c) The lesion is composed of small inconspicuous round lumina surrounded by broad strands of collagen fibre-rich connective tissue (haematoxylin and eosin 6 200). (d) Immunohistochemistry (6 200) demonstrates a peculiar network of small capillaries lined by CD34 positive endothelia (sclerosing capillary haemangioma).


321


In patient 2 the spleen weight was 160 g. The cut surface showed no abnormalities except for a soft spongy dark-red nodule measuring 6 mm in diameter. Histologically, the nodule consisted of irregularly thin walled anastomosing capillaries lined by slightly hyperplastic endothelia. Immunohistochemistry (standard avidin-biotin-complex peroxidase method) revealed the endothelia to be CD 34 positive, whereas CD 8 was

negative. For these morphologic findings, diagnosis of a splenic capillary haemangioma was made. Contrast enhanced ultrasound was performed in nine cases (Table 1). During arterial phase, hyperechoic enhancement was seen in seven patients, while an isoechoic pattern was seen in two patients. Three cases showed marked central filling with contrast (Figure 2), while the remaining patients demonstrated peripheral

(a)

(b)

(c)

(d)

Figure 2. Patient No. 5. (a) B-mode ultrasound shows a small well-demarcated hypoechoic central hyperechoic splenic tumour (arrows). (b) Power-Doppler ultrasound shows a vascular tumour with marked diffuse intratumoral flow signals (arrows). (c) In the early arterial phase (15 s), contrast enhanced ultrasound (SonoVueH) shows a central contrast filling indicating the increased vascular lesion (arrow). (d) In the late parenchymal phase (4 min) contrast enhanced ultrasound (SonoVueH) shows a hypoechoic contrast enhancement of the tumour in comparison with the surrounding normal splenic tissue (arrow). (Continued)

322



(‘‘ring sign’’) filling (Figure 3). All patients showed complete filling in of the lesion by the end of the arterial phase. In the parenchymal phase, all but two lesions were hypoechoic in reference to the surrounding tissue (Figures 2 and 3). In patients 1 and 7 an isoechoic enhancement was observed during the parenchymal phase (Table 1). Additional imaging diagnostic procedures which confirmed the intrasplenic tumour were MRI (n51) and contrast enhanced CT (n510). CT revealed an increased vascular lesion in 9 out of 10 patients. Patterns of arterial Figure 3. Patient No.10. (a) B-mode ultrasound shows two well-demarcated hypoechoic central hyperechoic splenic lesions. (b) Colour-Doppler ultrasound shows two vascular tumours with marked radial intratumoral flow signals in both lesions. (c) In the early arterial phase (11 s), contrast enhanced ultrasound shows a peripheral contrast enhancement of both tumours in comparison with the surrounding normal splenic tissue (arrows). (d) In the parenchymal phase (1 min), contrast enhanced ultrasound shows a complete filling of both lesions with an isoechoic enhancement of both tumours in comparison with the surrounding normal splenic tissue. (e) In the late parenchymal phase (3 min), contrast enhanced ultrasound shows a wash out phenomenon of both lesions with a hypoechoic enhancement of both tumours in comparison to the surrounding normal splenic tissue (arrows). (Continued)

(e) Figure 2. (Cont.) (e) Contrast CT shows a hyperintense vascular splenic tumour (arrow).

(a)

(b)

(b)

(c) The British Journal of Radiology, April 2006

(d)

(e) 323


(f ) (g) Figure 3. (Cont.) (f) Contrast CT shows a hyperintense peripheral ring enhancement of the splenic tumour (arrow). (g) Contrast CT shows a complete filling of the splenic tumour (arrow).

phase enhancement on CT were concordant to those seen with CES. No ‘‘parenchymal phase’’ was examined by CT. In patient 1, no tumour was seen on CT examination. In all 11 patients without surgical treatment, ultrasound follow-up was performed and revealed stable lesions in 10 cases. In patient 4 the lesion increased in size from 1.0 cm to 1.5 cm and a second increased vascular lesion of 0.5 cm was seen. Splenectomy was recommended, but the patient refused surgery; treatment with ultrasound follow-up continued (Table 1).

Discussion The spleen is regarded as ‘‘the forgotten organ’’ of the abdomen because splenic pathology is relatively rare [15]. Secondary lesions to the spleen (i.e. lymphoma, infarction, rupture and metastasis) are more common and often demonstrate a hypoechoic pattern on B-mode ultrasound. On CDS, splenic infarction and splenic rupture are characterized by an absent vascularity. In patients with splenic lymphoma and splenic metastases a reduced vascularity is the preferred CDS pattern [2, 6, 9, 11]. On CES splenic infarction as well as rupture is characterized by an unenhanced pattern. Splenic lymphoma demonstrates irregular peripheral enhancement, while metastasis are as usually predominantly non-enhancing, typically with a rim of peripheral enhancement [13]. In all other lesions presented here CES showed a marked hyperechoic arterial phase with variable patterns of contrast filling and relatively hypoechoic enhancement during the parenchymal phase. Besides characteristic imaging patterns, clinical data and follow-up are important additives for definitive diagnosis. A clinical problem occurs with the incidentally found focal hypoechoic intrasplenic tumour on ultrasound. Bmode ultrasound as well as CT and MRI are often nonspecific and histological verification is often warranted. The differential diagnosis of incidentally found hypoechoic splenic tumours includes hamartoma and 324

haemangioma. Splenic hamartoma (also called splenoma or nodular hyperplasia of the spleen) is a rare, benign tumour which usually presents as a solitary lesion [16]. Splenic haemangioma presents histologically with two different types: the more common cavernous type and the rare capillary type. For a better interpretation of increased vascular lesions found by CDS, CES or CT, it is helpful to review angiographic studies of splenic tumours. On angiography, hamartomas have a characteristic arteriographic pattern with irregular dilated tortuous vessels, with or without aneurysmal dilatation, within a well demarcated mass, with occasional vascular lakes or arteriovenous shunting [17–21]. In contrast to the more common cavernous haemangiomas, the capillary haemangiomas appear as multiple hypervascular nodules fed by dilated intrasplenic arteries without tumour vessels and vascular lakes [22]. Histologically, hamartomas resemble normal red pulp with slit like vascular spaces lined with endothelial cells [23]. In fact, some hamartomas with a distinctive lobular pattern have been reclassified as red pulp capillary haemangiomas [24]. On B-mode ultrasound, cavernous haemangiomas are described with a predominantly hyperechoic pattern and absent or reduced Doppler flow signals on CDS [9, 11, 13, 25]. Capillary haemangiomas as well as hamartomas present with a hypoechoic pattern on B-mode ultrasound [26, 27]. CES showed an increased vascular pattern in both the arterial and parenchymal phase with multiple radial blood-flow signals in the early arterial phase in a patient with a hamartoma [25]. In another case report, splenic hamartoma was markedly enhanced on CES [28]. In our series of hypoechoic increased vascular splenic lesions with clinical follow-up, the diagnosis of a benign splenic tumour was most likely. It should be noted that in one case, ultrasound follow-up examination revealed a mild increase in size of the vascular lesion was seen and an additional lesion found. In nine patients the increased vascular pattern was confirmed by contrast CT The British Journal of Radiology, April 2006


[29]. One lesion was not seen by CT (patient 1). In the same patient, CES also showed an isoechoic enhancement during the arterial and parenchymal phase. Catalano et al [30] reported about two patients with small capillary angiomas clearly visible in the conventional ultrasound and isoechoic in comparison with splenic parenchyma in all contrast phases of the CT examination (and therefore undetectable). In all other lesions presented here, CES showed a marked hyperechoic arterial phase with variable patterns of contrast filling and a hypoechoic enhancement during the parenchymal phase. We suggest that with an incidentally found splenic tumour the hypoechoic increased vascular pattern on ultrasound indicates a benign vascular tumour. Capillary haemangioma/hamartoma seems to be the most likely diagnosis, but in all cases a careful ultrasound follow-up is warranted.

13.

14.

15.

16. 17.

18.

References 1. Go¨rg C, Schwerk WB. Splenic infarction: sonographic patterns, diagnosis, follow-up, and complications. Radiology 1990;194:803–7. 2. Go¨rg C, Zugmaier G. Chronic recurring infarction of the spleen: sonographic patterns and complications. Ultraschall in Med 2003;24:245–9. 3. Go¨rg C, Weide R, Schwerk WB. Malignant splenic lymphoma: sonographic patterns, diagnosis, and followup. Clin Radiol 1997;52:535–40. 4. McKenney KL, Nunez DB, McKenney MG, Asher J, Zelnick K, Shipshak D. Sonography as the primary screening technique for blunt abdominal trauma: experience with 899 patients. AJR Am J Roentgenol 1998;170:979–84. 5. Go¨rg C, Co¨lle J, Go¨rg K, Prinz H, Zugmaier G. Spontaneous rupture of the spleen: ultrasound patterns, diagnosis, and follow-up. Br J Radiol 2003;76:704–11. 6. Go¨rg C, Schwerk WB. Sonographic findings of splenic metastases. Imaging/Bildgebung 1991;58:26–8. 7. Dachman AH, Ros PR, Murari PJ, Olmstedt WW, Lichtenstein JE. Non parasitic splenic cysts: a report of 52 cases with radiologic-pathologic correlation. AJR Am J Roentgenol 1986;147:537–42. 8. Ros PR, Moser RP, Dachman AH, Murari PJ, Olmstedt WW. Hemangioma of the spleen: radiologic-pathologic correlation in 10 cases. Radiology 1987;162:73–8. 9. Go¨rg C. The spleen. In: Meire M, Cosgrove D, Dewbury K, Farrant P, editors. Abdominal and general ultrasound, second edition. London, Edinburgh, New York: Churchill Livingstone 2001;17:379–445. 10. Wermke W, Gassmann B. Tumor diagnostics of the liver with echo enhancers: colour atlas. Berlin, Heidelberg, New York: Springer Verlag, 1998. 11. Go¨rg C, Schwerk WB. Color Doppler imaging of focal splenic masses. Eur J Radiol 1994;18:214–9. 12. Catalano O, Lobianco R, Sandomenico F. Splenic trauma: evaluation with contrast-specific sonography and a


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second-generation contrast medium. J Ultrasound Med 2003;22:467–70. Peddu P, Shah M, Sidhu PS. Splenic abnormalities: a comparative review of ultrasound, microbubble-enhanced ultrasound and computed tomography. Clin Radiol 2004;59:777–92. Catalano O, Sandomenico F, Matarazzo I, Siani A. Pictoral essay: sonography of the spleen. AJR Am J Roentgenol 2005;184:1150–6. Mortele´ KJ, Merjo PJ, Kunnen M. Tumoral pathology of the spleen. In: DeSchepper AM, Vanhoenacker F, editors. Medical imaging of the spleen. Berlin, Heidelberg, New York: Springer Verlag, 2000:101–23.. Steinberg JJ, Suhrland MJ, Valensi OJ. The association of splenoma with disease. Lab Invest 1985;52:65A. Kishikawa T, Numaguchi Y, Watanabe K, Matsuura K. Angiographic diagnosis of benign and malignant splenic tumors. AJR Am J Roentgenol 1978;130:339–44. Wexter L, Abrams ML. Hamartoma of the spleen. Angiographic observation. AJR Am J Roentgenol 1964;92:1150–5. Komaki G, Gombas OF. Angiographic demonstration of a calcified splenic hamartoma. Radiology 1976;121:77–8. Ro¨sch J. Tumours of the spleen: the value of selective arterography. Clin Radiol 1966;17:183–90. Teates CD, Seale DC, Allen MS. Hamartoma of the spleen. AJR Am J Roentgenol 1972;116:419–22. Tada S, Shin M, Takashima T, Noguchi M, Nishro I. Diffuse capillary hemangiomatosis of the spleen as a course of portal hypertension. Radiology 1972;104:63–4. Silverman ML, LiVolsi VA. Splenic hamartoma. Am J Clin Pathol 1978;70:224. Krishnan J, Danon AD, Frizzera G. Use of anti-factor VIIIrelated antigen (F8) and Q Bend 10 (CD34) antibodies helps classify the benign vascular lesions of the spleen. Mod Pathol 1993;6:94A. Niizawa M, Ishida H, Morikawa P, Naganuma H, Masamune O. Color Doppler sonography in a case of splenic hemangioma: value of compressing the tumor. AJR Am J Roentgenol 1991;157:965–6. Tang S, Shimizu T, Kikuchi Y, Shinya S, Kishimoto R, Fujioka Y, et al. Color Doppler sonographic findings in splenic hamartoma. J Clin Ultrasound 2000;28:249–53. Fujii T, Obara T, Shudo R, Maguchi H, Saitoh Y, Ura H, et al. Splenic hamartoma associated with thrombocytopenia. J Gastroenterol 1997;32:114. Chou Y-H, Chiou H-J, Tui C-M, Chiou S-Y, Hsia C-Y, Tsay S-H. Splenic hamartoma: presentation on contrastenhanced sonography. J Clin Ultrasound 2004;32:425–8. Abbott RM, Levy AD, Aguilera NS, Gorospe I, Thompson WM. From the archives of the AFIP: primary vascular neoplasms of the spleen: radiologic-pathologic correlation. Radiographics 2004;24:1137–63. Catalano O, Lobianco R, Sandomenico F, D’E`lia G, Siani A. Real-time contrast-enhancement ultrasound of the spleen: examination technique and preliminary clinical experience. Radio Med (Torino) 2003;106:338–56.

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A review of current local dose–area product levels for paediatric fluoroscopy in a tertiary referral centre compared with national standards. Why are they so different? M P HIORNS, A SAINI and P J MARSDEN Radiology, Great Ormond Street Hospital, Great Ormond Street, London WC1N 3JH, UK ABSTRACT. A prospective single centre study has been performed to assess dose–area product (DAP) values in children having fluoroscopic examinations and to revise local diagnostic reference levels (DRLs). DAP measurements for 2658 examinations performed in a dedicated fluoroscopy room over a period of 21 months were analysed. Data for the eight most commonly performed examinations (2215 cases) are presented. DAPs (75th centile) for upper gastrointestinal studies and micturating cystograms are substantially lower (by a factor of between 6 5 and 6 25) than the current national reference doses (NRDs), with some of the median values being 50 times lower. The small DAP values in all examinations demonstrate the substantial reduction in dose and consequent risk that can be achieved when both equipment performance and operator technique are optimized. Whilst we recognize that different institutions will have differing practices, it is important that practitioners are aware of the range of DAPs achievable and that NRDs do not necessarily represent best practice, and may falsely reassure.

The risks of ionizing radiation are higher in the paediatric population than in adults [1, 2]. It is therefore particularly important that the dose–area product (DAP) used in imaging children should be as low as practicable whilst providing the clinician with diagnostic information [3]. The actual doses achieved will vary greatly depending on the age, sex, body mass, body thickness and cooperation of the child. The type of equipment and its use by the operator will also affect dose levels. The most recent national guidelines published by the National Radiological Protection Board (NRPB) in 2002 which include paediatric procedures for the first time are based on a review of dose measurements collected over a 5 year period (1996–2000). DAPs were measured from 3671 paediatric fluoroscopic examinations in up to 29 different rooms [4]. From the results, recommended NRDs could only be defined for three paediatric fluoroscopic examinations; micturating cystograms, barium swallows and barium meals. To our knowledge, very little published data exists on the current range of doses being delivered throughout the country. The objectives of this study were to revise locally established DRLs and to compare performance with the current national guidelines.

Methods Data were collected prospectively on consecutive patients from September 2001 to May 2003. For each examination performed the patients name, hospital number, date, kV, mAs, screening time, DAP, examination type, and radiologist’s name were recorded. All studies were performed on a Siemens Polystar digital 326

Received 5 October 2004 Revised 8 August 2005 Accepted 15 August 2005 DOI: 10.1259/bjr/36530782 ’ 2006 The British Institute of Radiology

unit (installed 1999; Siemens, Erlangen, Germany). The DAPs were recorded in cGycm2 using a PTW Diamentor M4 DAP meter, which is specifically sensitive enough for paediatric work (PTW, Freiberg, Germany) with a high resolution of 0.01 cGycm2. The DAP meter is calibrated by our radiation physicists on a yearly basis and a tolerance of ¡3% is considered acceptable. Overall uncertainty is within ¡25% as recommended in the National Protocol [5]. Other routine quality assurance is completed every 3 months. The exact technique was at the discretion of the radiologist. For the period surveyed, a total of nine consultant radiologists and several specialist registrars, of varying experience, performed the examinations. Views obtained were tailored to the individual clinical question but basic views were obtained in all patients, these would include a mixture of grabbed images and spot exposures. Typically, both an upper gastrointestinal (GI) series and a micturating cystogram would comprise between six and 10 images.

Results Data were collected in 2658 cases on an individual basis, but were then grouped by age for specific examinations. Only one study were abandoned, during the period and all other studies produced images of diagnostic quality. All the data were analysed, but the data for the eight most common examinations are presented (2215 cases). As patient dose has been shown to increase with body size [6], patients were assigned into one of three groups; 0–12 months, 1–7 years inclusive and 8 years plus. The case distribution for The British Journal of Radiology, April 2006

DAP levels for paediatric fluoroscopy

Figure 1. Case distribution of examinations in our institution. GI, gastrointestinal; MCU, micturating cystourethrography; FT, follow through; IVU, intravenous urogram; NJT, nasojejunal tube.

our institution is shown (Figure 1) and data for the most common examinations are listed (Table 1). The DAP range, mean DAP, standard deviation, median DAP, 75th percentile and 90th percentile were all calculated. These values were used to establish revised DRLs for our institution (Table 2). The 75th percentile was used for the DRL (bold type), but we have also added median values in our local reference chart as this gives a good indication of overall practice and is less skewed by outliers. Radiologists are quickly able to see if their DAPs are similar to those of their colleagues. Some of our examination groups contained only small numbers, but as a large referral centre this is a reflection that very few of these examinations are done in any centre. In these cases we include the figures as a guide only.

Discussion In the UK, the adoption of the European Council Directive 97/43 Euratom [7] through the Ionising Radiation (Medical Exposure) Regulations [3] has made the establishment of DRLs and their update through clinical audit a mandatory requirement. National diagnostic reference levels have been established for a variety of mainly adult examinations as a guide to what is thought to represent ‘‘good and normal practice’’. Individual hospitals can then compare their DAP performance and modify their practice accordingly. The low numbers of patients and their distribution over the range of different ages and sizes limits the number of paediatric DRLs currently available. Nevertheless, it is important that hospitals in which a relatively low number of studies are performed continue to assess local DAP levels and have some guide to those used elsewhere. The NRDs as recommended by the NRPB are based on the 75th percentile of their survey data and are shown in Table 3 [4]. The NRPB stratified its data into five standard sizes of children corresponding to newborn The British Journal of Radiology, April 2006

babies, 1 year, 5 years, 10 years and 15 years old and children were allocated according to the nearest milestone. Dose levels, either in the form of entrance surface doses (ESD) per radiograph or DAP per examination were adjusted for body thickness or both the height and weight to give a value (in cGycm2) corresponding to the nearest standard size child. It can be seen that for all types of examination there are only small differences between the DAPs for the 1 year and 5 year old groups. The NRPB has therefore recommended that the same NRD be applied for both age groups. The establishment of reference doses in paediatric radiology as a function of patient size has been investigated for patients of the same age groups as used by the NRPB and it has been shown for plain film radiography in the trunk that, although there may be a significant difference in size between adjacent reference ages, normalization factors for any intermediate size were unlikely to exceed a factor of two [8, 9]. The practical disadvantage of applying a wide variety of standard sizes for dose assessment is more evident in non-specialist hospitals where patient numbers are small and long periods of data collection would be necessary. Since body size affects the dose received in any radiological procedure it would seem ideal to assign patients into narrow age groups that will not have an unduly large range of body weight and size [10]. Whilst we realise that the use of wider age groups may result in a larger range of doses, and obscure the effects of examination technique and equipment performance, the age groups used in this study have been designed for simplicity and for consistency with previous publications. Despite the use of a single dedicated fluoroscopy suite, there is a wide variation in measured DAPs within the individual age groups of all the examinations in our study. Such variation is similar to that found by both the NRPB and the Council of the European Communities (CEC), but in these studies the differences were largely attributed to variations in equipment and film–screen combinations [4, 11]. In our study, the range reflects the wider distribution of patient size, the technical difficulties of imaging uncooperative children and the variation of operator technique and experience (including training specialist registrars). We have found that some operators are consistently producing diagnostic quality studies at DAP levels significantly lower than others. In our institution barium meals and barium swallows are combined as one procedure, the upper GI series. We have compared our values with those given for both barium meals and barium swallows (Table 3) and have found that our closest equivalent DAPs for 75th percentile are between 9 and 25 times lower (UGI in patients 15 years, and ,12 months, respectively) and for a micturating cystogram are 5 to 18 times lower (MCU in patients of 8 years and older, and ,12 months, respectively). Reviewing our range of DAPs shows that, for example, in UGIs in children up to 1 year, even our 90th centile value (14.8 cGycm2) is 13 times less that the NRPB 75th centile (with the total range being 0.06– 41.9 cGycm2). The mean DAP results from the present study are also substantially less than similar data from previous studies, including another specialist centre [12, 13] (Table 4). A further study published in 2000 [8], 327

M P Hiorns, A Saini and P J Marsden Table 1. Paediatric fluoroscopic dose–area products (DAPs) recorded over a period of 21 months Examination

Upper GI series 0–1 1–7 8+ Dysphagia swallow 0–1 1–7 8+ Micturating cystourethrography 0–1 1–7 8+ Patalal screening 0–1 1–7 8+ Barium follow through 0–1 1–7 8+ Contrast enema 0–1 1–7 8+ Barium enema 0–1 1–7 8+ Intravenous urogrphy 0–1 1–7 8+

Mean DAP (cGycm2)

Median DAP (cGycm2)

75th Percentile (cGycm2)

90th percentile (cGycm2)

DAP range (cGycm2)

235 376 197

0.06–41.9 0.09–117.6 0.18–203.5

6.4 9.5 24.7

8.6 11.4 26

3.6 6.3 18.7

7.6 11.5 31.6

14.8 21 54.6

116 246 84

0.3–39.2 0.1–80.6 1.3–76.8

12.4 13 18.9

8.6 12.7 16.2

10 9.3 13.5

15.7 16.6 22.8

23.9 27.5 40.5

165 94 36

0.02–41.9 0.7–48.8 4–645.9

3.8 8.9 44.2

5.3 15.8 105.4

1.8 4.3 21.8

4.9 10 41.6

9.7 17.3 67.1

5 118 146

1.9–5.3 0.5–31.5 1.3–27.2

4.4 5.5 6.9

4 5.3 17.3

2 4 4.1

5.3 6.3 6.7

8.7 10.6 11

18 85 101

0.2–73.9 0.1–42.9 0.2–241.2

14.7 9.9 31.7

18.6 9.5 42.4

8.1 6.7 18.1

17.9 13.6 38.7

33.4 25 63

59 25 2

0.1–42.9 0.7–50.5 6.3–178.6

5 10 92.5

7.5 10.1 121.8

2.2 7.4 92.5

5.8 12.3 135.3

12.9 15.9 161.4

21 11 3

0.7–32.8 2.8–76.3 20.1–49

11.2 15 30.5

9.1 23.1 16

9.5 5.1 22.5

16.4 8.9 35.8

24.1 40.9 43.7

3 47 22

3–7.7 0.9–39.2 7–135.5

5.9 10.2 38.5

2.5 10.4 30.4

6.9 6.4 35.3

7.3 13.8 47.3

7.54 26.9 64.8

collecting DAP data from 12 European hospitals, investigated doses for micturating cystograms in patients between neonate and 15 years and obtained third quartile doses approximately 12 times higher than the figures we have obtained for all age groups. A study in Finland [14] considered a total of 217 patients having a wide range (12 types) of fluoroscopic examinations and derived third quartile DAP values for three of these examination types, which again are approximately 12 times higher than our figures for comparable examinations and age groups. However, it is important to note that these studies were conducted in the 1990s using conventional fluoroscopic equipment. The large differences in doses demonstrated are almost certainly due to differences in equipment and operator technique. The use of automatic post processing of data in digital systems prior to image display does not provide the operator with feedback on exposure levels and as such has the potential to increase dose. In our experience this is generally not the case and we have found that the use of low frame rates and last image hold ability is advantageous. A similar trend was observed by the NRPB who found that in 8 of 16 types of examination in adults, the use of digital spot imaging reduced mean 328

Standard deviation

Number of studies

DAP values by between 20% and 50% compared with conventional techniques but was not significant in the remaining 8 types of examination [4]. We recommend that non-specialist X-ray departments should assign responsibility for paediatric imaging to a select group of interested radiologists and radiographers, thus reducing any ‘‘learning curve’’ effects. Optimizing examination technique is of utmost importance. Pulsed fluoroscopy is an effective method of reducing dose and this feature is now standard on modern machines. Our fluoroscopy unit defaults to 15 pulses s21 but by selecting 3 pulses s21 the dose is immediately reduced by a factor of five assuming the screening time is unchanged, although the image will be more ‘‘steppy’’ due to misregistration. (On our unit the decrease in dose is directly proportional to the decrease in pulse rate, but on some units the pulse length increases slightly as the pulse rate decreases, which would therefore result in a less marked overall fall in DAP). Our current practice is to perform most screening at 3 pulses s21, increasing the rate if the child is very mobile or uncooperative. Images tend to be grabbed from the digital system, reserving occasional full exposures for delineation of fine detail, in difficult cases, or if there are unexpected findings. We The British Journal of Radiology, April 2006

DAP levels for paediatric fluoroscopy Table 2. Revised DRLs used in our institution. The 75th percentile values are demonstrated in bold. The values below represent the median dose–area product (DAP) for each examination. Children over the age of 8 years having contrast or barium enemas have been included together due to the low numbers in each group DAP (cGycm2) DAP (cGycm2) DAP (cGycm2) Age 0–1

Upper GI series

8 3.6 Dysphagia 16 swallow 10 Micturating 5 cystourethrogram 1.8 Patalal screening 5 2 Barium follow 18 through 8.1 Contrast enema 6 2.2 Barium enema 16 9.5 Intravenous 7 urography 6.9

Age 1–7

Age 8+

12 6.3 17 9.3 10 4.3 6 4 14 6.7 13 7.4 9 5.1 14 6.4

32 18.7 23 13.5 42 21.8 7 4.1 39 18.1 49 22.5 47 35.3

choose to use an overcouch tube whilst maximizing the distance between the tube and the image intensifier, with the table as low and close as possible to the image intensifier. Whilst this arrangement may slightly increase the dose we have found that paediatric patients find this arrangement much less frightening and this, combined with easier access to the child for the operator and the holders, reduces the length of the examination thereby reducing dose overall. Coning to a small field of view is achieved by the operator using a light beam diaphragm for guidance and we would consider this essential for a paediatric population. We use a low attenuation carbon fibre table. A removable grid is available, and is generally only used on patients over the age of 8 years Table 3. Recommended national reference doses (NRDs) for complete examinations on paediatric patients based on the NRPB 2000 review. Recommended values for 1 year and 5 year old sizes are given in brackets. Comparisons are given with the nearest local equivalent dose reference levels (DRLs) Examination

Barium meal

Standard age (years)

0 1 5 10 15 Barium swallow 0 1 5 10 15 Micturating 0 cystourethrogram 1 5 10 15

Current NRD (cGycm2)

Local DRL equivalent (cGycm2)

70 200 (200) 200 (200) 450 720 80 160 (150) 130 (150) 270 460 40 90 (100) 110 (100) 210 470

8 8 12 32 32 8 8 12 32 32 5 5 10 42 42


unless a younger child was particularly large for their age. Younger patients do not require a grid and this should not be used. We would not advocate the use of a unit with a fixed grid for use on children. We use 0.3 mm of added copper filtration and this is left permanently in place. Operators are encouraged to use the median values as appropriate ‘‘targets’’ during an examination. DAP meters are particularly useful for assessing and comparing the dose from screening procedures as ‘‘dose 6 area’’ provides a more useful indication of overall patient exposure than measurements of entrance surface dose at different locations [13]. Fluoroscopic screening time is of limited use as a measurement of dose as it makes no allowance for the influence of dose rate or field size and to the contribution of any spot/grabbed images. NRDs are limited in having to take into account the variation in equipment and film–screen combinations and as such give an idea of national practice, which is not necessarily the best practice. The use of local audit and critical review to ensure best practice should not be underestimated. Cook and co-workers [15] described reductions in ESD of greater than 50% for routine paediatric examinations. Whilst third quartile data is useful in setting National Reference Doses, and hence National DRLs, their application to establishing local DRLs is of limited value. The Joint Working Party of the IPEM/BIR/RCR/NRPB/ CoR [16] promotes the use of the mean of the distribution of room mean doses within an organization as the local DRL. However, an alternative method of setting a local DRL in situations were the sample size is small may be to display data as a histogram of DAPs for set age ranges and to exclude outliers beyond the limit of the curve using a mathematical ‘‘best fit’’ or agreed mathematical model. The DRL may then be a better reflection of ‘‘good and normal practice’’. It would be practical to choose a DAP based on the shape of the curve that would reliably include ‘‘normal practice’’ but would flag up the high dose outliers. This method would depend on having sufficient values to plot a meaningful histogram and the development of a reliable mathematical model. For everyday working practice, established median values may give a better impression of typical local DAP values and are a useful way for an operator to compare their cases against those of their colleagues. We display median and third quartile values for all tests in our fluoroscopy room as a guide.

Conclusions We have collected data for 2658 paediatric fluoroscopic examinations. Analysis of the data, and the derivation of third quartile DRLs, shows that our practice uses very much lower DAPs (and therefore DRLs) than the NRDs published by the NRPB and those published elsewhere in the literature. The variation between data in our study and the NRDs suggests that if the NRDs are used as a sole guide, many institutions will be falsely reassured and may be using greater doses than necessary. Only strict attention to technique and critical review of local reference levels will ensure best practice. NRDs or national DRLs, however they are determined, can only reflect ‘‘good and normal practice’’ 329

M P Hiorns, A Saini and P J Marsden Table 4. Mean DAP values are compared with those from previous studies Present study

Aberdeen data [10]

Newcastle data [12, 13]

Age range (years)

Examination

Mean DAP (cGycm2)

Age range (years)

Examination

Mean DAP (cGycm2)

Age range (years)

Examination

Mean DAP (cGycm2)

0–1

UGI MCU UGI MCU UGI MCU

6.4 3.8 9.5 8.9 24.7 44.2

0–1

BM MCU BM MCU BM MCU

29 17 38 15 73 55

0–1

BM MCU BM MCU BM MCU

130 168 223 262 546 433

1–7 8+

1–5 6–15

1–5 6–15

DAP, dose–area product; UGI, upper gastrointestinal series; BM, barium meal; MCU, micturating cystourethrography.

if current DAP data is submitted regularly to the NRPB or other national/international bodies.

References 1. National Radiation Protection Board. Occupational, public and medical exposure, Documents of the NRPB, Vol 4, No. 2. Chilton, UK: National Radiation Protection Board, 1993. 2. United Nations Scientific Committee on the Effects of Atomic Radiation. Sources, effects and risks of ionising radiation, UNSCEAR 2000 Report, Vol. II: effects. New York, NY: United Nations, 2000. 3. Department of Health. The Ionising Radiation (Medical Exposure) Regulations 2000. London: Department of Health, 2000. 4. National Radiation Protection Board. Doses to patients from medical X-ray Examinations in the UK – 2000 Review, NRPB-W14. Chilton, UK: National Radiation Protection Board, 2002. 5. National Protocol for Patient Dose Measurement in Diagnostic Radiology. IPSM/NRPB/CoR. Chilton, UK: National Radiological Protection Board, 1992. 6. Montgomery A, Martin CJ. A study of the application of paediatric reference levels. Br J Radiol 2000;73:1083–90. 7. Council Directive 97/43 Euratom of 30 June 1997 on health promotion of individuals against the dangers of ionising radiation in relation to medical exposure. Official Journal of the European Communities, 1997.

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8. Hart D, Wall BF, Shrimpton PC, Dance DR. The establishment of reference doses in paediatric radiology as a function of patient size. Radiat Prot Dosim 2000;90:235–8. 9. Hart D, Wall BF, Shrimpton PC, Bungay DR, Dance DR. Reference doses and patient size in paediatric radiology. NRPB – R318 November 2000. 10. Martin CJ, Farquhar B, Stockdale E, Macdonald S. A study of the relationship between patient dose and size in paediatric radiology. Br J Radiol 1994;67:864–71. 11. Council of the European Communities. European guidelines on quality criteria for diagnostic radiographic images in paediatrics, EUR 16261. Luxembourg: Office for Official Publications of the European Communities, 1996. 12. Chapple CL, Faulkner K, Lee REJ, Hunter EW. Results of a survey of doses to paediatric patients undergoing radiological examinations. Br J Radiol 1992;65:225–31. 13. Chapple CL, Faulkner K, Lee REJ, Hunter EW. Radiation doses to paediatric patients undergoing less common radiological procedures involving fluoroscopy. Br J Radiol 1993;66:823–7. 14. Servomaa A, Komppa T, Heikkila M, Parviainen T. Patient doses in paediatric fluoroscopic examinations in Finland. Radiat Prot Dosim 2000;90:239–43. 15. Cook JV, Kryriou JC, Pettet A, Fitzgerald MC, Shah K, Pablot SM. Key factors in the optimisation of paediatric xray practice. Br J Radiol 2001;74:1032–40. 16. Guidance on the Establishment and Use of Diagnostic Reference Levels for Medical X-Ray Examinations. IPEM Report 88. York, UK: Institute of Physics and Engineering in Medicine, 2004.



Peripheral dose from uniform dynamic multileaf collimation fields: implications for sliding window intensity-modulated radiotherapy 1

D S SHARMA, MSc, DipRP, 2ANIMESH, MSc, DipRP, 1S S DESHPANDE, MSc, DipRP, 3R D PHURAILATPAM, 1 4 4 MSc, DipRP, D D DESHPANDE, PhD, S K SHRIVASTAVA, MD and K A DINSHAW, FRCR Departments of 1Medical Physics and 4Radiation Oncology, Tata Memorial Hospital, Dr. Ernest Borges Marg, Parel, Mumbai, 400 012, India, 2Kirloskar Theratronics, Mumbai and 3Clinical Research Centre, ACTREC, New Mumbai, India ABSTRACT. The increase in the number of monitor units in sliding window intensitymodulated radiotherapy, compared with conventional techniques for the same target dose, may lead to an increase in peripheral dose (PD). PD from a linear accelerator was measured for 6 MV X-ray using 0.6 cm3 ionization chamber inserted at 5 cm depth into a 35 cm 6 35 cm 6 105 cm plastic water phantom. Measurements were made for field sizes of 6 cm 6 6 cm, 10 cm 6 10 cm and 14 cm 6 14 cm, shaped in both static and dynamic multileaf collimation (DMLC) mode, employing strip fields of fixed width 0.5 cm, 1.0 cm, 1.5 cm, and 2.0 cm, respectively. The effect of collimator rotation and depth of measurement on peripheral dose was investigated for 10 cm 6 10 cm field. Dynamic fields require 2 to 14 times the number of monitor units than does a static open field for the same dose at the isocentre, depending on strip field width and field size. Peripheral dose resulting from dynamic fields manifests two distinct regions showing a crest and trough within 30 cm from the field edge and a steady exponential fall beyond 30 cm. All dynamic fields were found to deliver a higher PD compared with the corresponding static open fields, being highest for smallest strip field width and largest field size; also, the percentage increase observed was highest at the largest outof-field distance. For 6 cm 6 6 cm field, dynamic fields with 0.5 cm and 2 cm strip field width deliver PDs 8 and 2 times higher than that of the static open field. The corresponding factors for 14 cm 6 14 cm field were 15 and 6, respectively. The factors by which PD for DMLC fields increase, relative to jaws-shaped static fields for out-offield distance beyond 30 cm, are almost the same as the corresponding increases in the number of monitor units. Reductions of 20% and 40% in PD were observed when the measurements were done at a depth of 10 cm and 15 cm, respectively. When the multileaf collimator executes in-plane (collimator 90 ˚) motion, peripheral dose decreases by as much as a factor of 3 compared with cross-plane data. The knowledge of PD from DMLC field is necessary to estimate the increase in whole-body dose and the likelihood of radiation induced secondary malignancy. Absorbed dose outside the primary radiation field (peripheral dose; PD) is of clinical interest in estimating out-of-field organ dose and subsequent long-term radiation sequelae. Potential effects of PD such as cataract formation, gonadal dysfunction and infertility, and damage to fetus, and their threshold doses, have been summarized by Fraass and van de Geijn [1]. The potential damage to the fetus from PD sometimes creates clinical dilemmas during decision making in radiation therapy of pregnant patients [2–4]. Prior investigators have confirmed that increase in the number of monitor units (MU) in tomotherapy, compared with a conventional technique for the same tumour dose, results in a higher whole-body dose [5– 7]. For a dose of 70 Gy to a head and neck tumour, whole-body dose has been shown to increase from between 0.2 Sv and 0.25 Sv for conventional therapy to nearly 2 Sv for tomotherapy; this in turn may increase The British Journal of Radiology, April 2006

Received 16 June 2004 Revised 31 May 2005 Accepted 19 July 2005 DOI: 10.1259/bjr/16208090 ’ 2006 The British Institute of Radiology

the risk of radiation-induced secondary malignancies by a factor of eight [8]. The increase in MU is a function of complexity of intensity-modulation, its delivery technique and collimator design of the treatment machine. Therefore, PD needs to be measured separately for the technique employed on the treatment machine. This study was designed to measure PD from 6 MV X-rays employing static and dynamic multileaf collimation (DMLC), which is the basis for dynamic intensitymodulated radiotherapy (IMRT). Dependence of PD on out-of-field distance, strip field width, field size, direction of multileaf collimator (MLC) motion and depth of measurement were investigated.

Methods and materials Peripheral dose was measured for 6 MV X-ray using a linear accelerator (LA) (Clinac 2100 C/D; Varian 331

D S Sharma, Animesh, S S Deshpande et al

Associates, Palo Alto, CA) equipped with an MLC consisting of 26 leaf pairs, each leaf projecting to a 10 mm leaf width at the isocentre distance. All measurements were carried out using a 0.6 cm3 Farmer type ionization chamber (PTW, Friedberg, Germany) inserted at 5 cm depth into a 35 cm 6 35 cm 6 105 cm plastic water phantom (Nuclear Associates, USA) under isocentric conditions. Field sizes of 6 cm 6 6 cm, 10 cm 6 10 cm and 14 cm 6 14 cm defined by conventional jaws were simulated in DMLC field mode, wherein strip fields of constant width 0.5 cm, 1 cm, 1.5 cm and 2 cm created by the opposing MLCs were moved with constant speed from one bank to the other in a rectilinear fashion perpendicular to the radial axis of the LA. Thus for every open field, six sets of PDs were measured – two from static open fields shaped by jaws and MLCs and four from DMLC. When the collimator is set at 0 ˚ (IEC 61217) [9], and MLCs are used either in static or dynamic mode, the upper (Y) jaws define the superior and inferior borders and the lower (X) jaws are positioned 0.8 cm distal to the most retracted leaf position on each side of the field. The detail performance characteristic of this MLC in dynamic dose delivery has been described elsewhere [10–12]. The number of MU required to deliver 1 Gy at 5 cm depth on the central axis was found by matching meter readings from the different DMLC fields to that from the corresponding, static jaw-shaped open field. These MUs were used for the subsequent PD measurement along a longitudinal axis to an out-of-field distance up to 60 cm.

The factor by which planned MU increases for different DMLC fields over that of the corresponding static fields for the same dose is defined to be MU multiplication factor. The same measurements were repeated for 10 cm 6 10 cm field size with collimator at 90 ˚, wherein the MLC executes in-plane motion. Variation in PD with depth was investigated for 10 cm and 15 cm depth under isocentric set up using 10 cm 6 10 cm field size having 1 cm strip field width.

Results The MU multiplication factor for different DMLC fields was found to increase with decreasing strip field width and increasing field size (Table 1). The measured PD data for 6 MV X-ray at 5 cm depth and for 0 ˚ collimator angle are shown in Figure 1 as a function of out-of-field distance. Each family of six curves represents the PD data for jaws and MLC shaped static field and four DMLC fields having different strip field widths. The dose at the depth of maximum dose (Dmax) on the central axis of every jaws-shaped static field was normalized to 100%. All the PD data were expressed as a percentage of the respective Dmax. PD data of all DMLC fields were found to manifest two distinct regions.

PD , 30 cm Table 1. Monitor unit (MU) multiplication factor (ratio of MU from dynamic multileaf collimation (DMLC) fields to that from correspondingopen field) for 6 MV X-ray for different DMLC fields having effective field sizes of 6 cm66 cm, 10 cm610 cm and 14 cm614 cm Strip field width (cm)

MU multiplication factor for different DMLC fields Field size 6 cm66 cm

Field size 10 cm610 cm

Field size 14 cm614 cm

0.5 1.0 1.5 2.0

6.96 3.98 2.59 1.89

10.93 6.73 4.70 3.53

13.75 8.94 6.49 4.96

The region extending up to 30 cm out-of-field distance, shows crests and troughs of varying amplitude; this is prominent for 6 cm 6 6 cm (Figure 1a) and is found to grow progressively smaller for larger fields (Figure 1b).

PD . 30 cm In this region PD curves from DMLC fields of different strip field width are seen running parallel and maintaining constant slope. The factor by which PD for different DMLC fields increase over jaws-shaped static field is observed to be almost the same as that of the

Figure 1. Peripheral dose in phantom from 6 MV X-ray for 0 ˚ collimator and field sizes (a) 6 cm 6 6 cm; (b) 14 cm 6 14 cm at 5 cm depth, normalized to 100% on the central axis at depth of maximum dose of static open field. MLC, multilead collimator. 332


Peripheral dose from DMLC fields

corresponding increase in MU. Although the absolute value of PD is comparatively small, the percentage increase in PD is higher for this region, showing maximum increase at 60 cm out-of-field distance. As expected, PD of all DMLC fields was found to be higher than the corresponding static open fields. For the same field size the smaller strip field width is seen delivering higher PD compared with larger strip field width. Also for the same strip field width, PD consistently increases with increasing field sizes. For 6 cm 6 6 cm field, DMLC with 0.5 cm and 2 cm strip field width deliver a maximum of 8 and 2 times higher PD than that of static open fields (Figure 1a). The corresponding factor for 14 cm 6 14 cm field is 15 and 6, respectively (Figure 1b). DMLC field with 0.5 cm strip field width delivers 1.5 and 3 times higher PD than that of 2 cm strip field width at 1 cm and 60 cm out-of-field distance, respectively. As an exception, jaws-shaped static fields may deliver more PD than certain DMLC fields with larger strip field width. This exceptional phenomenon is observed for outof-field distances extending up to 15 cm for 6 cm 6 6 cm and up to 7 cm for 14 cm 6 14 cm. Also, PD data of MLC shaped static field is found to be less than that of jaw-shaped static fields for out-of-field distances up to 30 cm, beyond which MLC shaped static fields show a slight increase in PD over jaws-shaped fields. PD data for static and DMLC fields measured with 90 ˚ collimation shows a reduction in value compared with data measured with 0 ˚ collimation. The magnitude of reduction is seen to depend on out-of-field distance and strip field width. For 10 cm 6 10 cm field, maximum reduction in PD data for jaws and MLC shaped static field are found to be by a factor of 2 and 2.5, respectively (Figure 2). For DMLC fields, a maximum reduction factor of 3 is observed at 30 cm out-of-field distance for the smallest strip field width of 0.5 cm (Figure 3). Similar characteristics were observed for larger strip field widths, albeit with progressively lower PD. MLC shaped static and DMLC fields show no change in PD at 20 cm from the field edge for these two collimator orientations. PD measured under isocentric condition at 5 cm, 10 cm and 15 cm depth using 10 cm 6 10 cm fields and 1 cm strip field width are shown in Figure 4. A maximum reduction of 20% and 40% at 60 cm from the field edge is observed for 10 cm and 15 cm depth, respectively, as compared with the data measured at 5 cm depth.

Figure 2. Composite peripheral dose (PD) distribution of 6 MV X-ray measured at 5 cm depth with 0 ˚ and 90 ˚ collimator angle from 10 cm 6 10 cm static field shaped by jaws and multileaf collimator (MLC), normalized to 100% on the central axis at depth of maximum dose of static open field.

Figure 3. Composite peripheral dose (PD) distribution of 6 MV X-ray measured at 5 cm depth with 0 ˚ and 90 ˚ collimator angles from 10 cm 6 10 cm field simulated in dynamic multileaf collimation (DMLC) mode using 0.5 cm sweeping gap width, normalized to 100% on the central axis at depth of maximum dose of static open field.

Discussion Pre-measured PD data from static fields [1, 2, 13–17], and an empirical relationship [18], are available as a guideline to evaluate approximate out-of-field organ dose from conventional treatment. The qualitative and quantitative behaviour of our jaw-shaped open field (10 cm 6 10 cm, 0 ˚ collimation and 5 cm depth) PD data is in agreement with other findings [2, 17]. But beyond 30 cm out-of field distance, our PD data for MLC shaped static field disagrees with that of Mutic et al [17] who have reported reduction in PD at all out-of-field distances. However, when the collimator is rotated to 90 ˚ our data agrees with the findings of Mutic et al [17]. The British Journal of Radiology, April 2006

Figure 4. Peripheral dose (PD) in phantom from 6 MV X-rays measured at 5 cm, 10 cm and 15 cm depth with target to chamber distance of 100 cm. Effective field size of 10 cm 6 10 cm having 1 cm sweeping gap width and 0 ˚ collimator is used. PD is normalized to 100% on the central axis at depth of maximum dose of static open field. 333

D S Sharma, Animesh, S S Deshpande et al

Although IMRT offers significant advantages for dose conformality to irregular target volumes with sharp dose gradient beyond the target, it also increases MU/Gy and hence whole body dose. This may be a concern for longterm radiation sequelae in patients treated with this technique [8, 19]. Whole-body dose data are available from tomotherapy [5–8]. However, the same data cannot be applied to sliding window IMRT technique as the collimator scatter and transmission, head leakage and internal scatter, which are the main components of PD, depend on the collimator design of treatment machine and beam delivery technique. Our study was carried out to find PD from different DMLC fields which closely simulated sliding window IMRT. When static fields are simulated in DMLC mode, PD was found to increase by a factor of 2 to 15, similar to the increase in MU for DMLC fields over corresponding static open fields. Increase in MU due to the use of a universal wedge have also shown to increase PD by a factor of 2–4 compared with PD resulting from unwedged beam [14–16]. Even though the intensity of radiation is not modulated in our dynamic fields, different strip field width can be thought to represent the different complexity of intensity profile required in IMRT. As the degree of modulation increases, i.e. peaks and valleys in the intensity profile are sought to be closely spaced, these highly modulated beams require narrow MLC openings leading to sharp increase in MU. With sliding window IMRT, higher MU becomes an absolute corollary and the planned MU tends to be larger in cases where the fluence distribution is more modulated. Even though PD from DMLC field is generally observed to be more than that of jaw-shaped static field, a slight reduction in PD is observed near the field edge for DMLC field with larger sweeping gap width. This may be attributed to reduction in collimator scatter and transmission of DMLC fields offered by the stationary portion of the MLC. The increase in MU for smaller sweeping gap widths for the same field and dose, increases both scatter and leakage dose leading to the increase in PD. Besides internal scatter and leakage dose, PD in the proximal region (5–30 cm) seems to be a complex interplay of leaf scatter, leaf end design, MLC placement in treatment head, MLC motion and beam-on time. Similar to our crest and trough effect appearing within 30 cm of the field edge, Greene et al [20] have reported a sharp peak at 20 cm from the edge of 10 cm 6 10 cm field due to maximum scatter coming from the face of the beam-defining jaws, which disappears for large field sizes as the point that sees maximum scatter from the face of the jaws shifts into the primary beam itself. In our case, this situation never realises completely, even for large DMLC fields, probably because scatter radiation arises both from the face of jaws and from the rounded end face of the sweeping MLCs. The crest and trough effect appearing prominently for smaller dynamic fields may be clinically significant especially when a critical organ is lying close to the primary field. If an organ at risk is lying at the distance of 15 cm from the field edge (field size 6 cm 6 6 cm) it would received 4.5 times more dose than the case where this effect is not taken into account. Moreover, it is not known if commercially available treatment planning 334

systems take into account this pattern of PD during optimization and dose computation. Further study is required to segregate the different components of scatter and leakage. Inclusion of PD characteristics in dose calculation algorithms is suggested. Most published data for static field shows small variation in PD with depth of measurement [1, 2, 15]. However, similar to the finding of Tatcher et al, our measurement for DMLC field shows decrease in PD with depth of measurement. The largest field studied is 14 cm 6 14 cm as limited by the maximum range of travel of individual leaves relative to the carriage. While normalizing the dose at depth of dose maximum to 100%, central axis depth dose characteristics of the DMLC field of a particular size is assumed to remain same as that for a static field. The uncertainty associated with low dose measurement especially at 60 cm with lower MU generally for the static fields are not considered in the presented data. For open fields where the MU is small, MU was scaled by a factor of 2–3 to get good signal at largest out-of-field distance. The reproducibility of meter reading for DMLC fields was observed to be less than 2% during the measurement. As the IMRT field employed in the clinical situations consists of a mixture of multiple strip fields having different gap width (beamlet size), validation of the phantom measured PD from uniform DMLC field needs to be carried out using the clinically employed fluence.

References 1. Fraass BA, van de Geijn J. Peripheral dose from megavoltage beams. Med Phys 1983;10:809–18. 2. Stovall M, Blackwell CR, Cundiff J, Novack DH, Palta JR, Wagner LK, et al. Fetal dose from radiotherapy with photon beams: Report of AAPM Radiation Therapy Task Group No. 36. Med Phys 1995;22:63–82. 3. Sneed PK, Albright NW, Wara WM, Prados MD, Wilson CB. Fetal dose estimates for radiotherapy of brain tumors during pregnancy. Int J Radiat Oncol Biol Phys 1995;32:823–30. 4. Sharma DS, Rakesh J, Tambe CM, Animesh, Deshpande DD. Effect of tertiary multileaf collimator (MLC) on foetal dose during three-dimensional conformal radiation therapy (3DCRT) of a brain tumour during pregnancy. Radiother Oncol 2004;70:49–54. 5. Followill D, Geis P, Boyer A. Estimates of whole-body equivalent dose produced by beam intensity modulated conformal therapy. Int J Radiat Oncol Biol Phys 1997;38:667–72. 6. Mutic S, Low DA. Whole-body dose from tomotherapy delivery. Int J Radiat Oncol Biol Phys 1998;42:229–32. 7. Meeks SL, Paulino AC, Pennington EC, Simon JH, Skwarchuk MW, Buatti JM. In vivo determination of extra-target doses received from serial tomotherapy. Radiother Oncol 2002;63:217–22. 8. Verellen D, Vanhavere F. Risk assessment of radiationinduced malignancies based on whole-body equivalent dose estimates for IMRT treatment in the head and neck region. Radiother Oncol 1999;53:199–203. 9. International Electrotechnical Commission, IEC Report 61217: Radiotherapy Equipment – Co-ordinates, Movements and Scales. IEC, Geneva, Switzerland, 1997. 10. Chui CS, Spirou S, LoSasso T. Testing of dynamic multileaf collimation. Med Phys 1996;23:635–41.


Peripheral dose from DMLC fields 11. LoSasso T, Chui CS, Ling C. Physics and dosimetric aspects of a multileaf collimation system used in the dynamic mode for implementing intensity modulated radiotherapy. Med Phys 1998;25:1919–27. 12. Low DA, Sohn JW, Klein EE, Markman J, Mutic S, Dempsey J. Characterization of a commercial multileaf collimator used for intensity modulated radiation therapy. Med Phys 2001;28:752–6. 13. Greene D, Chu GL, Thomas DW, Taylor RJ. Dose levels outside radiotherapy beams. Br J Radiol 1983;56:543–50. 14. Sherazi S, Kase KR. Measurements of dose from secondary radiation outside a treatment field: effects of wedges and blocks. Int J Radiat Oncol Biol Phys 1985;11:2171–6. 15. McParland BJ. Peripheral dose of two linear accelerators employing universal wedges. Br J Radiol 1990;63:295–8.


16. Mutic S, Esthappan J, Klein EE. Peripheral dose distributions for a linear accelerator equipped with a secondary multileaf collimator and universal wedge. J Appl Clin Med Phys 2002;3:302–9. 17. Mutic S, Klein EE. A reduction in the AAPM TG-36 reported peripheral dose distributions with tertiary multileaf collimation. Int J Radiat Oncol Biol Phys 1999;44:947–53. 18. Mazonakis M, Damilakis J, Varveris H, Theoharopoulos N, Gourtsoyiannis N. A method of estimating fetal dose during brain radiation therapy. Int J Radiat Oncol Biol Phys 1999;44:455–9. 19. Hall EJ, Wuu CS. Radiation-induced second cancers: The impact of 3DCRT and IMRT. Int J Radiat Oncol Biol Phys 2003;56:83–8. 20. Greene D, Karup PGG, Sims C, Taylor RJ. Dose levels outside radiotherapy beams. Br J Radiol 1985;58:453–6.

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The DXL Calscan heel densitometer: evaluation and diagnostic thresholds J A THORPE,

MSc, BSc

and S A STEEL,

MSc, BSc

Centre for Metabolic Bone Disease, Royal Hull Hospitals NHS Trust, Hull Royal Infirmary, Anlaby Road, Hull HU3 2RW, UK ABSTRACT. The DXL Calscan (Demetech AB) is a new dual energy X-ray absorptiometry device for determining heel bone mineral density (BMD). The system is based on the standard technique of dual energy X-ray absorptiometry (DXA), using a fan beam configuration, but introduces an additional laser measurement of heel thickness intended to improve accuracy. We have examined the utility, in vitro and in vivo performance of the DXL Calscan and established triage thresholds based on the UK’s National Osteoporosis Society guidelines on peripheral densitometry. The Calscan proved convenient, easy to use and was stable over time and within a range of operating temperatures. Short-term in vitro precision as %CV, with phantom repositioning, was 0.75% and long term precision 0.73%. Precision in vivo, determined from duplicate right heel scans of 67 subjects, was 1.19%. Effective radiation dose to the patient was ,0.1 mSv per scan. 140 white females (70 osteoporotic and 70 nonosteoporotic), aged 55–70 years underwent scans of both heels. Subjects were defined as osteoporotic or non-osteoporotic on the basis of axial DXA (spine L2–L4 and total hip). Triage thresholds for reassurance-referral or referral-treatment were 0.391 g cm22 and 0.306 g cm22 for non-dominant and 0.395 g cm22, 0.294 g cm22 for dominant heel, respectively. The non-dominant heel proved slightly superior to the dominant for triage purposes. Of the seven non-osteoporotic subjects misclassified as osteoporotic by Calscan of either heel, six had severe axial osteopenia. If operated by trained personnel and used in appropriate populations exhibiting risk factors, the Calscan is well suited for use in the management of post-menopausal osteoporosis.

Measurement of bone mineral density (BMD) by dual energy X-ray absorptiometry (DXA) is now well established as the method of choice for osteoporosis assessment [1–3]. BMD assessment of the lumbar spine and hip by DXA represents the current gold standard due to the greater associated morbidity and mortality of fractures at these two sites, superior fracture prediction [4, 5] and response to treatment [6]. In addition to axial assessment, there are a variety of DXA devices available for measuring BMD in the forearm, heel and hand. The DXL Calscan (Demetech AB, Solna, Sweden) is a new peripheral device for calcaneal BMD assessment, based on fan beam DXA. The Calscan also incorporates a laser measurement of heel thickness to improve the accuracy of calcaneal BMD. Standard DXA assumes a two compartment model of tissue masses, the first bone and the second a composite of lean and adipose tissue at an assumed constant ratio. This assumed ratio does not allow for fluctuations in lean and adipose tissue proportions that have been demonstrated to occur at the spine [7] and are likely to occur at the calcaneus or elsewhere [8, 9]. This leads to calcaneal DXA providing a precise but potentially inaccurate estimate of BMD, with the degree of inaccuracy dependent on body mass index [8]. This study has been funded in part by Demetech AB, Solna, Sweden. Additional funding for J Thorpe provided by the local osteoporosis charity, OSPREY.

336

Received 16 March 2005 Revised 8 August 2005 Accepted 15 August 2005 DOI: 10.1259/bjr/22191429 ’ 2006 The British Institute of Radiology

These inhomogeneities can be corrected by solving the BMD equation as a three component model of bone, lean and adipose tissue. Swanpalmer [10, 11] described how a third X-ray energy could achieve this, but concluded that a significantly higher photon count (and hence scan times) would be required to maintain an acceptable degree of precision. Jonson [12] deduced that if the combined width of all three components were known, e.g. the width of the heel, the ratio of soft to lean tissue could then be derived and corrected for. The laser heel width measurement on the Calscan provides this additional dimension allowing the derivation of BMD from a three component model, whilst theoretically maintaining DXA precision [13]. As with other peripheral DXA (pDXA) devices [14], the Calscan is smaller, portable, cheaper and has a lower radiation dose than axial densitometers. However, pDXA results at the calcaneus cannot be interpreted using the WHO definition [15] and do not correlate perfectly with bone density at either spine or hip. The imperfect correlation can lead to pDXA misclassifying subjects to the opposite diagnostic group to which they would have been classified by axial DXA [16–18], particularly for subjects with BMD scores close to diagnostic thresholds. Hence considerable debate remains over how such peripheral devices might best be employed in the clinical setting. The UK-based National Osteoporosis Society (NOS) has recently stated The British Journal of Radiology, April 2006

Evaluation and diagnostic thresholds of the DXL Calscan

that current evidence supports the use of peripheral devices in a triage rather than diagnostic role, and has established a method for determining the required triage thresholds [19, 20]. The aim of this study was to determine the in vitro and in vivo operating conditions of the Calscan and to establish triage referral thresholds based on the NOS guidelines.

Materials and methods The DXL Calscan bone densitometer (Figure 1) utilizes a fan beam, dual energy X-ray source and a solid state detector to perform a scan of the heel. A region of interest is positioned automatically by the software to derive BMD. A concomitant measure of heel thickness is obtained using the laser reflection to correct for variations in the soft to lean tissue ratio. The Calscan, at 25 kg in weight and 80 cm long by 43 cm wide and 33 cm tall, is relatively compact and easy to transport and includes wheels at one end and a carry handle. As with all X-ray equipment, the Calscan is potentially subject to changes in tube temperature after performing an acquisition, and as a portable device it may also be subject to fluctuations in performance due to environmental changes. To counter this, the software (version 1.3.1) requires a warm-up acquisition when the device is switched on and a 4 min cooling down period after each acquisition.

of 22.1. Long term in vitro precision was determined using daily single phantom scans acquired over a period of 6 months as part of routine quality assurance. Sixteen tests were conducted to assess the effects of ambient temperature, device movement or tube heating on accuracy or precision. For each test the device was disconnected, moved to a different room, warmed up and a phantom scan acquired as soon as the warm-up was complete. The device was given a further 30 min to stabilize, then a second phantom scan was acquired and the device powered down and allowed to cool for 30 min before beginning the next test. Temperature was measured on an alcohol room thermometer throughout. Electrical, laser safety and radiological protection surveys were also carried out.

In vivo methodology Subjects Females attending for routine BMD of spine and hip were approached for participation in this study. These were referred on the basis of agreed local risk criteria which are broadly in agreement with those of the Royal College of Physicians [1]. The study was approved by the local research Ethics Committee and informed consent was obtained. All subjects were white and between the ages of 55 years and 70 years (Table 1). A total of 140 women were recruited.

In vitro methodology Phantom based studies were conducted to test the effect of temperature, stability following relocation and to determine in vitro precision. The DXL-Calscan comes with a manufacturer-supplied phantom, made from shaped pieces of aluminium depicting the calcaneus, embedded in acrylic. Short term precision was determined as percent coefficient of variation (%CV) from 30 scans with phantom repositioning and 30 without. The device was operated according to the manufacturer’s instructions. All measurements were taken on the same day. Accuracy was calculated from comparison with the phantom’s stated BMD of 0.347 g cm22, established from a central reference machine, corresponding to a T-score

Axial DXA assessment Subjects underwent BMD of lumbar spine and hip using a GE-Lunar Prodigy (GE-Lunar, Madison, WI) as part of their routine examination and clinical management was determined on the basis of the results. In our centre, DXA of the right hip is performed unless contraindicated. For the purposes of this study, if the lower of either L2–L4 spine or total hip BMD T-score values was below 22.5, the subject was classified as osteoporotic. Otherwise they were classified as nonosteoporotic. When lumbar vertebrae showed clear signs of degenerative changes, the individual vertebrae affected were excluded from the lumbar spine results. Four subjects had individual vertebrae excluded – one osteoporotic and two non-osteoporotic subjects with degenerative changes of L4 and one non-osteoporotic subject with changes at L3. For eight other subjects (four osteoporotic, four non-osteoporotic by final diagnosis), two or more vertebrae on the same subject showed Table 1. ‘‘Evaluation and Diagnostic Thresholds of the DXL Calscan’’. Mean (standard deviation) subject demographic variablesfor whole group, osteoporotic and non-osteoporotic subjects, respectively

Figure 1. The Demetech DXL Calscan and phantom. The British Journal of Radiology, April 2006

Whole group

Osteoporotic

Non-osteoporotic

n 140 Age (years) 62.7 (4.5) Height (cm) 159.4 (6.5) Weight (kg) 64.3 (10.7) BMI (kg m22) 25.3 (3.9)

70 63.2 (4.3) 159.2 (7.0) 62.6 (11.1) 24.7 (3.9)

70 62.2 (4.7) 159.5 (6.0) 66.0 (10.1) 26.0 (3.9)

BMI, body mass index.

337

J A Thorpe and S A Steel

degenerative changes. For these eight subjects the spine results were disregarded and diagnosis was made by total hip DXA alone. In one case this caused the subject to move from the osteoporotic to non-osteoporotic group. Recruitment continued until 70 osteoporotic and 70 nonosteoporotic subjects were enrolled.

Comparison of the results recorded at the 16 highest temperatures (range 23.8 ˚C to 26.9 ˚C) with the results at the 16 lowest (21.3 ˚C to 23.8 ˚C) did not change phantom BMD or precision significantly. For the 16 higher temperature scans, phantom BMD and precision was 0.347 g cm22 and 0.64% CV, respectively. For the 16 lower temperature scans, phantom BMD and precision was 0.348 g cm22 and 0.58% CV, respectively.

Calcaneal DXA assessment BMD of both heels was obtained using the DXL Calscan. To determine in vivo precision, 67 of the 140 subjects underwent a repeat acquisition of the right heel, with repositioning between each. The calcaneal regions of interest (ROI) were manually checked and, if deemed necessary, ROI position was corrected as per the user manual instructions. For the repeat Calscan acquisition, the second scan for each subject was analysed on a separate day to the first to reduce the possibility of operator bias during any ROI repositioning.

Results Operational utility The time from scan initiation to appearance of the results is 94 s, with an additional 4 min required to allow the X-ray tube to cool before another acquisition can be taken. The Calscan is able to image either heel from the same side of the device making it easier for the patient and minimizing floor space required where both heels are to be scanned. As for all equipment using ionizing radiation, the Calscan requires a standard radiation safety assessment but also an additional laser safety assessment. The footwell of the Calscan was of an open design, and had the advantage of allowing the operator to manually assist the positioning of the heel. The open design allowed easy access for the operator and was comfortable for the patient, but did require some attention to achieve the ideal positioning.

In vitro results and environmental effects Short term in vitro precision (coefficient of variation) was 0.76% CV (mean BMD 0.347¡0.0026 g cm22) with phantom repositioning, 0.75% (0.347¡0.0025 g cm22) without. Long term precision was 0.73%. The device was accurate, with no measurable difference between mean phantom BMD as measured on our machine compared with that of the central reference machine. Average phantom BMD and precision for the 32 environmental scans was 0.347 g cm22 and 0.62% CV. Average phantom BMD and precision for the 16 scans taken as soon as possible after a warm-up scan, i.e. after the enforced 4 min cooling down period between scans, was 0.347 g cm22 and 0.65% CV. For the 16 scans acquired after the tube had been allowed to cool for half an hour, average BMD was 0.348 g cm22 and 0.61% CV. During the 32 scans of the 16 environmental tests, the room temperature varied from 21.3 ˚C to 26.9 ˚C, the upper value being slightly outside the manufacturer’s recommended operating range of 15 ˚C to 25 ˚C. 338

Radiation and laser safety The effective radiation dose to the patient was ,0.1 mSv per scan and a controlled area of 0.5 m was defined around the device in order to comply with IRR 1999 [21]. At this distance, scatter dose to the operator would not exceed annual dose limits, even at maximum scan throughput. The laser assessment found the Calscan laser itself to be class 2 by UK/European/US standards and thus capable of causing eye damage, but the location of the laser within the footwell removed the possibility of accidental exposure and so the laser was deemed to be safe (class 1), provided the Calscan outer casing was in place. The permanent filtration and laser class were not marked on the casing as is required to comply with UK/ EU standards [22, 23]. A laser warning label was added to the Calscan and local rules were established that reflective objects should be kept clear of the footwell, as stated in the user manual. No other laser precautions were deemed necessary. No safety problems with the laser occurred during the project, but it was noticed that opaque black hosiery could produce spurious BMD results, although other hosiery did not.

In vivo results Subject demographics and bone density results are summarized in Tables 1 and 2, whilst coefficients of determination (adjusted R2) between key variables are shown in Table 3. Mean in vivo BMD of the right heel for all 67 subjects (19 osteoporotic, 48 non-osteoporotic) given repeat measurements was 0.357 g cm22 (range 0.186–0.518 g cm22, standard deviation 0.074 g cm22). Mean absolute difference between paired results was 0.0046 g cm22 (range 0–0.018 g cm22). Calscan precision for the 67 subjects as %CV (derived from root mean square) was 1.19%. Taking the osteoporotic and non-osteoporotic precision groups separately, mean BMD for the 19 osteoporotic subjects was 0.299 g cm22 (0.186–0.437 g cm22, SD 0.065). Mean absolute difference was 0.0042 g cm22 (0– 0.012 g cm22). Precision was 1.30%CV. For the 48 nonosteoporotic precision subjects mean BMD was 0.392 g cm22 (range 0.281–0.518 g cm22, SD 0.061). Mean absolute difference was 0.0048 g cm22 (0– 0.0018 g cm22). Precision was 1.09%CV.

Establishing triage thresholds In the revised NOS guidelines on peripheral DXA, Blake et al [20] recommend the use of peripheral devices in a triage role as an adjunct to axial DXA and suggest a The British Journal of Radiology, April 2006

Evaluation and diagnostic thresholds of the DXL Calscan Table 2. ‘‘Evaluation and Diagnostic Thresholds of the DXL Calscan’’. Mean (standard deviation) DXA bone density and T-score results for whole group, osteoporotic and non-osteoporotic subjects, respectively

Spine L2–L4 Total hip Non-dominant heel Dominant heel

BMD T-score BMD T-score BMD T-score BMD T-score

Whole group

Osteoporotic

Non-osteoporotic

0.948 (0.173) 22.10 (1.45) 0.832 (0.122) 21.39 (1.02) 0.356 (0.064) 21.96 (0.97) 0.356 (0.064) 21.96 (0.97)

0.822 (0.070) 23.15 (0.59) 0.774 (0.092) 21.88 (0.77) 0.328 (0.054) 22.37 (0.82) 0.328 (0.050) 22.38 (0.75)

1.076 (0.152) 21.03 (1.26) 0.892 (0.121) 20.89 (1.01) 0.383 (0.062) 21.53 (0.93) 0.383 (0.065) 21.54 (0.99)

Units: BMD (g cm22); T-score (St Dev). BMD, bone mineral density.

Table 3. ‘‘Evaluation and Diagnostic Thresholds of the DXL Calscan’’. DXA BMD Correlation (adjusted R2 value) Spine L2–L4

Spine L2–L4 1 Total femur Dominant heel Non-dominant heel

Total femur

Dominant heel

Nondominant heel

0.379 1

0.285 0.331 1

0.276 0.350 0.905 1

method for defining the two triage thresholds required. The upper of the two thresholds is set at a point above which only 10% of osteoporotic subjects would fall, whilst the lower threshold is a point below which only 10% of non-osteoporotic subjects would fall. Subjects who fall above the upper threshold would be assumed non-osteoporotic, whilst subjects who fall below the lower threshold assumed osteoporotic. Subjects falling between the two would be recommended for referral for axial DXA. Using the 140 subjects in this study, the upper and lower thresholds for Calscan for the non-dominant and dominant heels are shown in Figure 2. The proportion of subjects in the equivocal group, and therefore requiring axial DXA, is shown in Table 4. Based on these thresholds of 0.391 g cm22 and 0.306 g cm22 for the non-dominant and 0.395 g cm22 and 0.294 g cm22 for

the dominant heel, the referral rates for the Calscan in this group were 52.9% (non-dominant) and 58.6% (dominant), but with an error margin of ¡9% due to the small sample size. Of the seven non-osteoporotic subjects misclassified as osteoporotic by Calscan of either non-dominant or dominant heels, six had severe osteopenia (axial T-score ,22). A better estimate of the expected referral rate can be drawn from comparing our derived thresholds (and confidence intervals) to the mean and standard deviation of the Calscan reference data. The reference data are drawn from a population of 993 Swedish women between 15 years and 85 years of age [24] (381 between 50 years and 69 years), albeit a population without known risk factors. If we assume a hypothetical referral group with an even distribution of subject ages from 55 years to 70 years and the same spread and trend in BMD results as the Swedish reference group, the expected mean referral rate at the non-dominant heel would be 36.7%. Adjusting for the distribution of ages seen in our 140 subjects, the figure would be 36.8%.

Discussion The Calscan proved reliable, precise, accurate and easy to use. Calscan performance was stable within a normal range of room temperatures, and was not affected by recent movement of the device. There was no difference

Figure 2. (a) DXL Calscan upper and lower triage thresholds for the non-dominant heel. (b) DXL Calscan upper and lower triage thresholds for the dominant heel.


339

J A Thorpe and S A Steel Table 4. ‘‘Evaluation and Diagnostic Thresholds of the DXL Calscan’’. Referral by number of subjects

Thresholds:

Upper Lower Units: BMD (T-score). BMD in g cm22. T-score in standard deviations Above upper threshold: reassure Non-osteoporotic Osteoporotic Between thresholds: refer Non-osteoporotic Osteoporotic Below lower threshold: treat Non-osteoporotic Osteoporotic Units: number of subjects Referral rate

in performance when acquiring scans in quick succession, or with 30 min breaks between them, with the enforced 4 min period between scans appearing sufficient time for the X-ray tube to cool. Radiation dose to patient and scatter dose to operator were low and the device requires only a small controlled area. There is no lower limit of applicability and the Ionising Radiation Regulations still apply, requiring therefore the advice of a radiation protection advisor, device risk assessment, production of local rules, written procedures and appropriate training of staff. At 1.19% CV, precision in vivo for our 67 precision subjects as a whole was slightly superior to the 1.24% for the Calscan and 1.28% for the GE-Lunar Pixi reported by Hakulinen et al [8], who performed repeat scans on 38 (18 male, 20 female) subjects with a mean (SD) age of 59.7 years (¡9.4 years). Although at 1.30% CV the precision for our 19 osteoporotic precision subjects is poorer than the 1.09% for the 48 non-osteoporotic subjects, at 0.0042 g cm22 versus 0.0048 g cm22, the mean absolute error per repeat measurement was actually lower for the osteoporotic than for the nonosteoporotic group, and so the difference in precision can be explained by the difference in the mean BMD scores for the two groups. As with all DXA systems, attainment of good precision requires technical staff to be trained, experienced and to practice good technique. We found the coefficient of determination (R2) with spine and total hip DXA to be 0.28 and 0.35 at the nondominant heel. Correlation at spine was lower than the 0.59 reported by Martini et al [25], or the 0.61 reported by Hakulinen [8]. It is not clear if this is due to differences in the sample groups and the small size of the Martini and Hakulinen samples. Using T-scores, the upper and lower triage thresholds as defined by the NOS method for the Calscan were at 21.4 and 22.7, respectively, for the non-dominant heel. These T-scores are only applicable to post-menopausal white women aged 55–70 years who meet the normal criteria for axial bone densitometry examination. As with all T-scores, the exact threshold values depend on the reference range. Were this to be changed, then the Tscores would need to be recalculated from the underlying BMD scores of 0.391 g cm22 and 0.306 g cm22. In addition, the T-score thresholds of any peripheral devices employed in a triage role are likely to become more negative with advancing subject age [20], but the unreliability of spine DXA in subjects over 70 years of age makes the calculation of peripheral threshold values 340

Non-dominant heel

Dominant heel

0.391 (21.4) 0.306 (22.7)

0.395 (21.4) 0.294 (22.9)

30 8 33 41 7 21

31 7 32 50 7 13

52.9%

58.6%

problematic for such a group, without resorting to total hip BMD alone. There continues to be a growth in demand for bone densitometry services through increased awareness of health professionals and the public, rising healthcare costs of fragility fractures and the development and introduction of new bone protective treatments. The provision and availability of such services, however, remains patchy and inconsistent. In an area where demand on axial DXA is exceeding capacity, peripheral DXA could prove useful in a triage role to ensure best and most cost effective use of this resource. However, a comprehensive analysis of the resource implications of such an approach is required. Applying the triage thresholds to the population used for this evaluation would suggest that over 50% would require referral for axial DXA. As indicated, the study was not powered to provide an accurate assessment of referral proportion and the true figure is probably below 40%. Provided the cost per case for the heel DXA measurement is less than 60% that of a spine and hip measurement, there should be a net saving. Where the peripheral device is community or primary care based, there may be an increase overall in patients identified due to the more accessible nature of the service which would reduce the potential cost savings and also increase the burden on the prescribing budget. Where there is no access to axial DXA locally, peripheral DXA may play a role in identifying those at risk of fragility fracture provided it is used with care and in appropriate populations with clearly identified clinical risk factors. The Calscan device appears suitable for either role using the thresholds derived in this study. There is a high proportion (95%) of the more metabolically active trabecular bone in the calcaneus [15] which would suggest that this site is sensitive to mechanisms affecting bone metabolism. This, together with the advantage of being a weight bearing bone, should better reflect the changes occurring at the spine and hip than at other peripheral sites. The moderate correlation between the heel and axial sites observed in this study may be due to sample bias as the subjects were drawn from those attending for bone densitometry. The lack of agreement observed generally between sites is also partly due to the varying trabecular to cortical ratios with the spine being 50% trabecular and hip 40%. There is no published evidence to date that patients commenced on treatment on the basis of falling below the lower triage threshold by pDXA could be monitored The British Journal of Radiology, April 2006

Evaluation and diagnostic thresholds of the DXL Calscan

by pDXA. Ringe et al demonstrated promising results in heel BMD with ibandronate [26], but they employed a non-standard technique and do not compare the observed 15% increase at 2 years with the least significant change. It is also known that some bone protective treatments are only effective in reducing fractures in those defined osteoporotic by hip BMD [27]. There are no data yet on effectiveness of treatments in those targeted by the pDXA triage technique although use of the derived lower pDXA threshold provides 90% confidence that the patient would be found osteoporotic by spine and hip, with almost all the remainder severely osteopenic by spine or hip. Use of peripheral devices in a triage role as an adjunct to an established axial DXA service could bring substantial benefits to both patient and healthcare providers, and the Calscan is well suited for this purpose. However, it should be operated only by qualified personnel, used in selected populations and results interpreted in conjunction with clinical risk factors for fragility fracture.

Acknowledgments The authors wish to acknowledge Vincent Mann and Melanie Auty for their technical assistance in the environmental measurements for this paper.

References 1. Royal College of Physicians. Osteoporosis: Clinical guidelines for prevention and treatment. London, UK: RCP, 1999. 2. U.S. Department of Health and Human Services. Bone Health and Osteoporosis: A Report of the Surgeon General Chapter 8: Assessing the Risk of Bone Disease and Fracture. U.S. Department of Health and Human Services, Office of the Surgeon General, 2004. 3. National Osteoporosis Society. Position statement on the reporting of dual energy x-ray absorptiometry (DXA) bone mineral density scans. Bath, UK. NOS, August 2002. 4. Marshall D, Johnell O, Wedel H. Meta-analysis of how well measures of bone mineral density predict occurrence of osteoporotic fractures. Br Med J 1996;312:1254–9. 5. Stone KL, Seeley DG, Lui L-Y, et al. BMD at multiple sites and risk of fracture of multiple types: long-term results from the Study of Osteoporotic Fractures. J Bone Miner Res 2003;18:1947–54. 6. Eastell R. Treatment of postmenopausal osteoporosis. N Engl J Med 1998;338:736–46. 7. Formica C, Loro ML, Gilsanz V, Seeman E. Inhomogeneity in body fat distribution may result in inaccuracy in the measurement of vertebral bone mass. J Bone Miner Res 1995;10:1504–11. 8. Hakulinen MA, Saarakkala S, Toyras J, Kroger H, Jurvelin JS. Dual energy x-ray laser measurement of calcaneal bone mineral density. Phys Med Biol 2003;48:1741–52. 9. Hausler KD, Rich PA, Barry EB. Water bath and contact methods in ultrasonic evaluation of bone. Calcif Tissue Int 1997;61:26–9. 10. Swanpalmer J, Kullenberg R, Hansson T. The feasibility of triple-energy absorptiometry for the determination of bone mineral, Ca and P in vivo. Physiol Meas 1998;19:1–15.


11. Swanpalmer J, Kullenberg R, Hansson T. Measurement of bone mineral using multiple-energy x-ray absorptiometry. Phys Med Biol 1998;43:379–87. 12. Jonson R, Mansson LG, Rundgren A, Szucs J. Dual-photon absorptiometry for determination of bone mineral content in the calcaneus with correction for fat. Phys Med Biol 1990;35:961–9. 13. Swanpalmer J, Kullenberg R. A new measuring device for quantifying the amount of mineral in the heel bone. Ann N Y Acad Sci 2000;904:115–7. 14. Patel R, Blake GM, Fogelman I. Radiation dose to the patient and operator from a peripheral dual x-ray absorptiometry system. J Clin Densitom 1999;2:397–401. 15. WHO Technical Report Series 843. Assessment of fracture risk and its application to screening for postmenopausal osteoporosis. Geneva, Switzerland: World Health Organization, 1994.. 16. Grampp S, Genant HK, Mathur A, et al. Comparisons of non-invasive bone mineral measurements in assessing agerelated bone loss, fracture discrimination and diagnostic classification. J Bone Miner Res 1997;12:697–711. 17. Lu Y, Genant HK, Shepherd J, et al. Classification of osteoporosis based on bone mineral densities. J Bone Miner Res 2001;16:901–10. 18. Faulkner KG, Von Stetton E, Miller P. Discordance in patient classification using T-scores. J Clin Densitom 1999;2:343–50. 19. National Osteoporosis Society. Position statement on the use of peripheral x-ray absorptiometry in the management of osteoporosis. Bath, UK: NOS, November 2004. 20. Blake GM, Chinn DJ, Steel SA, Patel R, Panayiotou E, Thorpe J, et al. The Revised National Osteoporosis Society Position Statement on Peripheral x-ray Absorptiometry: a list of device specific thresholds for the clinical interpretation of pDXA examinations. Osteoporosis International; In press. 21. The Ionising Radiations Regulations 1999. London, UK, The Stationery Office Limited, ISBN 0 11 085614 7; 1999. 22. British Standards Institute (BSI). Medical electrical equipment. Particular requirements for safety. Specification for xray source assemblies and x-ray tube assemblies for medical diagnosis. British Standards Publishing BS EN 60601-228;1993. 23. European Committee for Electrotechnical Standardization (CENELEC). Safety of laser products – Part 1: Equipment classification, requirements and user’s guide. EN 60825-1; 1994. 24. Kullenberg R. Reference database for dual x-ray and laser Calscan bone densitometer. J Clin Densitom 2003;6:367–72. 25. Martini G, Valenti R, Gennari L, Salvadori S, Galli B, Nuti R. Dual x-ray and laser absorptiometry of the calcaneus: comparison with quantitative ultrasound and dualenergy x-ray absorptiometry. J Clin Densitom 2004;7: 349–54. 26. Ringe JD, Dorst A, Faber H, Ibach K, Preuss J. Threemonthly ibandronate bolus injection offers favourable tolerability and sustained efficacy advantage over two years in established corticosteroid-induced osteoporosis. Rheumatology 2003;42:743–9. 27. Cummings SR, Black DM, Thompson DE, Applegate WB, Barrett-Connor E, Musliner TA, et al. Effect of alendronate on risk of fracture in women with low bone density but without vertebral fractures: results from the Fracture Intervention Trial. JAMA 1998;280:2077–82.

341


SHORT COMMUNICATION

Time-dependent observer errors in pulmonary nodule detection 1

D MANNING,

PhD, FInstP,

1

S C BARKER-MILL,

PhD,

1

T DONOVAN,

MSc

and 2T CRAWFORD,

PhD

1

School of Medical Imaging Sciences, St Martin’s College, Lancaster LA1 3JD and 2Department of Psychology, Lancaster University, Lancaster, UK ABSTRACT. The work was carried out to investigate differences in visual search characteristics between groups of observers with different levels of experience in the task of pulmonary nodule detection in chest radiology and we report here on these differences in respect of time related decisions. Volunteer observers were divided into three groups depending on their level of expertise. There were eight radiologists, eight radiographers and eight novices. Their task was to detect pulmonary nodules in a test bank of 120 digitized posteroanterior (PA) chest radiographs. Five of the eight radiographers were tested twice: once before and once after a 6-month training programme in interpretation of the adult chest radiograph. During each test session the observers’ eye movements were tracked. Data on the observers’ decisions through Alternate Free Response Operating Characteristic (AFROC) methodology were correlated to their eye-movement and fixation patterns. True negative decisions from all observers were associated with shorter fixation times than false negative decisions. No correct negative decisions were made after fixations exceeding 3 s.

In previously reported studies we have investigated observer experience [1, 2] and the effects of lesion conspicuity [3] on performance in nodule detection from plain chest radiology. Comparisons of diagnostic performance through alternate free response receiver operating characteristic (AFROC) [4] showed that after training and extensive caseload experience radiographers’ detection rate improved to approach that of the experts. AFROC is a variation on receiver operating characteristic methodology that takes into account the location in the image of the observer decision. Through this technique the observer must indicate not only his decision on the disease status of the image (positive or negative for the presence of a pulmonary nodule) but also the location of the lesion. The observer groups were as follows:

N N N N

eight first-year student radiographers considered as novices to chest interpretation tasks; eight experienced clinical radiographers before training in chest interpretation; five of the eight radiographers after their training in chest interpretation; and eight radiologists.

Eye-tracking these observers gave insight into differences between the groups in terms of their visual search strategies and we concluded that, amongst other things, the experts were more economical in their patterns of search, carried out fewer fixations and spent less time on the task. After a training period that included a minimum of 500 cases but no specific instruction in search patterns, the radiographers developed spontaneously similar search strategies to those of the 342

Received 2 March 2005 Revised 24 August 2005 Accepted 30 September 2005 DOI: 10.1259/bjr/13453920 ’ 2006 The British Institute of Radiology

radiologists. But both radiologists and trained radiographers had a false negative error-rate in excess of 40% [2]. We accept that the requirements of AFROC are stringent (a false negative is defined by a missed lesion rather than an incorrect case decision); but this is still a significant miss-rate and in this paper we report our observations on the relationship between the experience of the observer, the duration of visual attention and the probability of a false negative error in the task.

Eye movement, visual attention and visual scan paths The field of view in humans extends over 180 ˚ but only the centre of the field provides sharp details. Consequently, we move our eyes (and heads) to bring interesting features into the centre. The pause over the point of interest is known as a foveal fixation and has duration of about 100–300 ms before we move our interest to a new location in a fast jump called a saccade. Fixations are characterized by multiple (clustering) nature when we dwell extensively on a location because the eyes do not remain stationary for long before losing sensitivity. Saccades are too fast for information to be gathered during their operation and so eye-tracking experiments assume that fixations represent the location of conscious attention. Viewers are not conscious of the pattern of their saccades during an observation of a scene, however, and it has been shown that scanpaths for an individual are modified by the task presented [5]. The path is the sequence in which image details are scrutinized but fixation point clusters illustrate the parts of the scene the observer find most interesting as The British Journal of Radiology, April 2006

Short communication: Time-dependent observer errors

opposed to the parts that receive no attention at all. It has been noted that the same details in an image can be accessed using quite different scanpaths but similar fixation clusters [6]. As a result, analysis of fixation cluster location and dwell time gives a better opportunity for performance comparisons than the scanpath sequence if the research question is focused on whether image features are being hit. However, scanpath data can give fascinating insights into how individuals prioritise locations of semantic interest although the analysis of these paths is difficult and sometimes inconclusive. Errors in radiology are known to be multifactorial but a sizable proportion are known to be perceptual in origin [7, 8]. This has led to interest in the possible sources of perceptual error and the visual strategies of radiologists in their task [9]. Early work on the visual dwell behaviour of radiologists during film reading was carried out by Kundel et al [10–12] who went on to later describe survival curve methods for analysing timerelated decision-making in those observers [13, 14]. Our aim here is to report on the timing of decisions, their accuracy and their relationship to experience by the survival analysis of decisions from eye-tracking data.

Aim The aim of this study was to report an observed relationship between the experience of the observer, the duration of visual attention and the probability of incorrect decisions in pulmonary nodule detection.

Materials and methods Observers Eight volunteers were recruited from a cohort of first year radiography students who acted as the novice group of film readers. Eight experienced radiographers volunteered from a post-graduate course on chest image interpretation at the commencement of their course. Five of these radiographers then offered themselves for repeat testing at the end of their 6 months training and 500 cases of experience. Eight radiologists with extensive and current experience in chest radiology volunteered to act as the most experienced group of observers. All observers gave their consent and all performance data were made anonymous. Recruitment of all the observers and the conduct of the experiments followed the ethical guidelines for experiments involving volunteer human subjects at St Martin’s College and Lancaster University. At the time when the data for these experiments were collected (in 2001) there was no requirement of COREC approval for NHS employees who were acting as volunteer participants in this type of research.

Detectable nodules Each observer viewed a bank of 120 digitized chest images of adults. The images contained 81 pulmonary nodules agreed as significant in pathological appearance from confirmed radiological reports. Nodules were The British Journal of Radiology, April 2006

roughly circular and ranged in diameter size from 5 mm to 20 mm with varying degrees of measured conspicuity [3]. Nine films contained more than one nodule and 30 nodules were located in these multinodule films. Normal films were included in the observer task and the complete test bank was divided into three sets of 40 images with prevalence-rates of 12%, 50% and 82%.

Observer performance measurement Alternate free response operating characteristic methodology (AFROC) was used [4]. This required observers to indicate a location to a decision for a lesion and to assign a score between 1 and 4 on their level of confidence in that decision. A zero score was allocated to all decisions of ‘‘no nodule present’’. In AFROC methodology false positive decisions are treated in the following way: the highest scoring false positive decision is the only one recorded per image which avoids the possibility of infinite values in summing false positive responses. Observer test sessions were never longer than 1 h to avoid the effects of fatigue on performance. There was a minimum 6-month interval between the before- and after-training observer tests on the radiographers to give an effective case-memory washout period.

Parameters All eye-tracking data comparing the performance of the observer groups were processed through the ASL (Applied Science Laboratories, Bedford, MA), software EYENALH. The measured parameter from the eye tracking for use in the present study was the accumulated dwell time at each decision point.

A fixation We defined a fixation (visual dwell) as a point of gaze remaining continuously within a 1 ˚ area for at least 100 ms. A 1 ˚ angle subtended from the eye to the image at a viewing distance of 40 cm approximates to a circle the size of a 2.5 cm diameter disc at arms length. Refixations were summed to give cumulative fixation times to clusters when there was overlap of the 1 ˚ areas up to a 5 ˚ area. These definitions are ones that are commonly used in work of this kind [9, 14]. The observers were allowed to search freely and no limit was imposed on the duration of inspection for any given image. A true negative decision was defined as a timed fixation of a lesion-free zone of the chest image that elicited a zero response from the observers on the AFROC scale. The dwell-time data for all fixations related to positive and negative decisions were analysed through the statistical package SPSSH to provide information on the percentage survival of decisions over time. These data allowed us to characterize the observers’ decisions in greater detail than true and false negative and true and false positive outcome, giving an opportunity to 343

D Manning, S C Barker-Mill, T Donovan and T Crawford

identify time-related features of decision outcomes. The time related information on decisions gave an opportunity to analyse time differences between correct and incorrect decisions whether positive or negative.

Results Time related decisions: survival analysis We investigated how the four possible decision outcomes of true and false positive (TP FP) and true and false negative (TN FN) related to the duration of gaze through a survival analysis of the fixation data. The results are presented in Tables 1–4 and Figures 1–4. Survival analysis is used in this context as a means of showing the proportions of decisions that are completed for each category (TP FP TN FN) at increasing accumulated time intervals of visual attention. It is analogous to the cell survival curves used in radiobiology to indicate the proportion of cells surviving increasing radiation doses [15]. So in Table 1 and Figure 1 for example, 50% of all true positive decisions made by the radiologists survived 2200 ms of accumulated visual attention on the tumour location but none of Table 1. Radiologists Survival duration (ms) for each decision group TN

FN

TP

FP

% surviving

0 200 220 450 680 750 790 820 900 1010 1800 2250

0 220 240 500 700 800 1000 1250 1600 2000 3000 4750

0 400 455 1000 1680 1800 2200 2500 3000 3500 4500 5000

0 220 250 750 1700 1800 2200 2700 3500 4000 4700 5500

100 95 90 80 70 60 50 40 30 20 10 0

TN, true negative; TP, true positive; FN, false negative; FP, false positive.

Table 2. Radiographers post training in chest interpretation Survival duration (ms) for each decision group TN

FN

TP

FP

% surviving

0 200 300 450 500 500 550 600 800 1000 1500 3000

0 220 330 500 500 550 700 1100 2000 2500 4000 5000

0 300 800 900 1200 1850 2500 3000 4400 5000 6000 6600

0 500 600 800 1000 2000 4000 5000 5500 6000 7000 8000

100 95 90 80 70 60 50 40 30 20 10 0


344

their true positive decisions took more than 5000 ms of visual scrutiny. In the tables we have highlighted the time values for the completion of all the observer decisions in their outcome categories to draw attention to the differences in duration between the true and false decisions in each case. For all observers the false negative decisions took them significantly longer than true negatives. For positive decisions, correct decisions were made faster than incorrect ones except in the case of the novices who dwelt longer on genuine lesions before deciding they were pathologies. The survival curves show that for all observers 50% of all their true negative decisions were made within 1000 ms of gaze duration and the false negative decisions have dwell-times somewhere between the positives (true and false) and the true negatives. This is consistent with the findings of others who have carried out this type of analysis [14] but our results indicate extension along the time axis with decreasing levels of experience.

Discussion The aim of this work was to analyse eye-tracking data to classify the false negative errors made when observer groups with different levels of experience are asked to detect pulmonary nodules. The different levels of experience in the groups give some insight into how, if at all, errors vary with expertise. The time related data provided measurable differences in the fixation times associated with observers making correct or incorrect decisions.

Survival analysis Observers’ decisions are made over time periods that can be related to the duration of visual fixations from eye-tracking data. Mean or median dwell times over trials and between readers can be calculated but dwell times are not normally distributed, making statistical comparisons difficult to interpret. The technique of survival analysis has been found useful in these circumstances [13, 14] and this operation on the data demonstrated some consistent findings for categories of decisions. Figures 1 and 2 show the family of decisions survival curves for radiologists and radiographers with chest interpretation training derived from the data in Tables 1 and 2. Virtually all the true negative decisions made by these, the most experienced observers, were made within 2 s of cumulative fixation time on an image feature. False negative decisions were characterized by longer dwell times but no negative decisions (either true or false) were made by radiologists after 4.75 s fixation. Figures 3 and 4 and Tables 3 and 4 show similar findings for the low experience groups but there is a tendency towards longer fixation times for incorrect negative decisions with decreasing levels of experience. In the case of the novices (Table 4) lesions were missed after a cumulative gaze duration of up to 8 s. The time difference in the proportions of true and false negative decisions made by these observers can be summarized by saying that in The British Journal of Radiology, April 2006

Short communication: Time-dependent observer errors Table 3. Radiographers pre training in chest interpretation Survival duration (ms) for each decision group TN

FN

TP

FP

% surviving

0 300 330 330 400 560 700 850 1000 1200 1600 2400

0 300 400 500 600 1200 1600 2000 2400 3000 4500 6500

0 350 800 1100 2000 3000 3300 3800 4200 5100 6500 8200

0 350 820 1200 2000 3100 3800 4800 5500 6500 8000 9000

100 95 90 80 70 60 50 40 30 20 10 0


Table 4. Novices Survival duration (ms) for each decision group

Figure 1. Time-related decisions for radiologists. Compared

TN

FN

TP

FP

% surviving

0 300 400 500 550 600 680 800 900 1500 2200 3000

0 300 400 500 600 1000 1500 1800 2200 3500 5000 8000

0 500 1000 1700 2000 2600 3800 4500 5500 6500 9000 12000

0 400 500 800 1200 1500 2000 3000 4500 5600 8000 10000

100 95 90 80 70 60 50 40 30 20 10 0

with less experienced observers the radiologists made their decisions faster. The separation between the positive and negative decision pair curves is closer than for all other groups.


this experiment, all negative decisions made after gaze duration of 3 s were incorrect. It seems that correct negative decisions (correct decisions that normal features are not nodules) tend to be made rapidly after fixation occurs and that this is a feature of expert performance. Conversely, incorrect negative decisions are characterized by extended gaze duration especially in novice readers. When areas of the images hold their attention for several seconds of accumulated fixation time, observers show a semantic interest in the appearances. This suggests that they are suspicious of the feature and they are operating on the information at a perceptual/cognitive level. These errors are not failures of detection but of recognition and decision and can be explained partly by the visual ambiguities of the image and partly by the level of experience of the reader. The finding may be important because of: (a) its potential for reducing false negative error if visual dwell information is fed back to the reader; and (b) its potential use in training schemes for radiologists and monitoring the effects of caseload experience. If these results are reproducible in other settings and our interpretation of their meaning is correct there are several ways that they may help in the improvement of observer performance. A simple expedient is to inform film readers The British Journal of Radiology, April 2006

Figure 2. Time-related decisions for radiographers posttraining. After experience and training over 6 months and 500 cases the radiographers have speeded their decision making and markedly reduced the separation of their true and false negative curves. Compare with Figure 3.

that their negative decisions are more likely to be incorrect when they are made after a period of indecision over a particular image feature. In short, we suggest informing observers that if a feature looks suspicious enough to warrant more than 2 s of their attention it is probably not innocent. More sophisticated aids linked to these findings 345

D Manning, S C Barker-Mill, T Donovan and T Crawford

discriminator in predicting whether a negative decision will be correct or incorrect. We consider this observation to be potentially important for feedback strategies for education purposes and to aid performance.

Acknowledgment We wish to thank the radiologists, radiographers and students, all of whom are associated with the School of Medical Imaging Sciences at St Martin’s College, Lancaster, UK, who kindly agreed to take part in this study. Eye-tracking equipment for this research was supplied through a grant from the Peter Barker-Mill Memorial Trust.

References

Figure 3. Time-related decisions for radiographers pre-training.

Figure 4. Time–related decisions for novices. Novices show a similar trend to the experienced observers although their positive decision curves are reversed and more widely separated. All their decisions extend further along the time axis.

might involve computer aided feedback to observers in real-time to give indication of the gaze duration for individual locations in the image.

Conclusion The data and their analysis from an eye-tracking experiment have given insights into the errors made by observers with different levels of experience. The most notable outcomes of this are that the duration of fixation on a feature in the chest image may be an effective 346

1. Manning DJ, Ethell S, Crawford T. An eye-tracking AFROC study of the influence of experience and training on chest xray interpretation. Medical imaging 2003; Image Perception Observer Performance and Technology Assessment. Editors Dev Chakraborty and Elizabeth Krupinski, Proc SPIE 2003;5034:257–266. 2. Manning DJ, Ethell SC, Crawford T, Donovan T. How do radiologists do it? The influence of experience and training on searching for chest nodules. Radiography Available online 23 March 2005. Item 49, www.sciencedirect.com/ science/journal/10788174. 3. Manning DJ, Ethell SC, Donovan T. Detection or decision errors? Missed lung cancer from posteroanterior chest radiograph. Br J Radiol 2004;77:231–5. 4. Chakraborty DP. Statistical power in observer performance studies: comparison of the receiver operating characteristic and free response method in tasks involving localization. Acad Radiol 2002;9:147–56. 5. Yarbus AL. Eye movements and vision. New York, NY: Plenum Press, 1967. 6. Kundel HL. Reader error, object recognition and visual search. In: Chakraborty D, Eckstein M, editors. Medical imaging 2004: image perception, observer performance and technology assessment. Proceedings of SPIE 2004;5372: 1–9. 7. Fitzgerald R. Error in radiology. Clin Radiol 2001;56:938–46. 8. Chief Medical Officer. Learning from failure: evidence and experience. An Organisation with a Memory. London: Stationery Office, 2000:1–7. 9. Gale AG. Human response to visual stimuli. In: Hendee WR, Wells PNT, editors. The perception of visual information, 2nd edition. New York, NY: Springer, 2000:134. 10. Kundel H, Nodine C, Carmody D. Visual scanning, pattern recognition and decision making in pulmonary nodule detection. Invest Radiol 1978;13:175–81. 11. Kundel H, Nodine C, Krupinski E. Searching for lung nodules: visual dwell indicates locations of false positiveand false negative decisions. Invest Radiol 1989;24:472–8. 12. Kundel H, Nodine C, Toto L. Searching for lung nodules: the guidance of visual scanning. Invest Radiol 1991;26: 777–81. 13. Anderson S, Auquier A, Hauk WW, Oakes D, Vandaele W, Weisberg HI. Statistical methods for comparative studies. New York, NY: John Wiley & Sons, 1980. 14. Kundel H. Visual search in medical images. In: Beutel J, Kundel H, Van Metter R, editors. Handbook of medical imaging Vol 1 Physics and psychophysics. Bellingham WA: SPIE, 2000:837–58. 15. Motulsky H. Intuitive biostatistics. Oxford: Oxford University Press Inc., 1995:Chapter 6.



SHORT COMMUNICATION

Study of scattered radiation for in-air calibration by a multipledistance method using ionization chambers and an HDR 192Ir brachytherapy source 1

N P PATEL,

MSc, DipRP,

2

B MAJUMDAR,

MSc, PhD, PDC

and 3V VIJAYAN,

MSc, DipRP, PhD

1

Department of Medical Physics, Acharya Harihar Regional Cancer Centre, Cuttack-753007, India, Department of Physics, Government College of Science, Raipur-492010, India and 3Health Physics Unit, Institute of Physics, Bhubaneswar-751005, India

2

ABSTRACT. The aim of this study is to estimate the room-scatter correction when measuring air kerma rate of an HDR 192Ir brachytherapy source by in-air calibration. The variation in scattered radiation due to the specially designed jig and from the room walls was also studied. Two therapy ion chambers of volume 0.1 cm3 and 0.6 cm3 were used in the present study. Air kerma was measured by placing the source at several distances between 10 cm and 20 cm from the chamber. The scatter radiation was determined by superimposing the theoretically derived model curve of known scatter (based on the inverse square law) over the plot of measured air kerma strength values. The scatter radiation was estimated in terms of percentage of the primary radiation at 10 cm measurement distance. The scatter estimated by the 0.6 cm3 chamber at two positions was 0.33% and 0.59%, respectively. Similarly the scatter estimated at two other positions by the 0.1 cm3 chamber was 0.58% and 1.11%. This variation in scatter with position as well as with the chamber was due to the varying scatter contribution from components of the measurement set-up. The scatter radiation becomes constant at a distance greater than 100 cm from the walls of the room. We conclude that a fixed chamber with changing source positions should be used in multiple-distance measurement of air kerma rate when using a measurement jig.

Calibration of brachytherapy sources is an end-user requirement prior to clinical use. The recommended quantity for specifying a brachytherapy source is the air kerma strength (AKS) [1]. Air kerma strength is defined as the product of air kerma rate at a calibration distance, d, in free space, K(d), measured along the transverse bisector of the source, and the square of the distance, d. A source calibration accuracy of ¡3% relative to existing AKS standards seems reasonable [2]. At present, high dose rate 192Ir sources are widely used for brachytherapy treatment due to their high specific activity and low effective photon energy. Calibration of an HDR 192Ir source is performed by in-air measurement or by a solid phantom technique using a thimble chamber. An alternative method is to use a welltype chamber. An IAEA technical document has discussed several factors including measurement distance, chamber size, positioning uncertainty, charge leakage and room scatter to be considered for in-air measurement [3]. The measurement distance should be chosen such that the combined uncertainty in source calibration due to other factors can be minimized. The optimum measurement distance estimated for a Farmer-type chamber with an HDR 192Ir source is 16 cm [4]. Experimental studies have estimated the room scatter using a large volume spherical The British Journal of Radiology, April 2006

Received 7 December 2004 Revised 28 September 2005 Accepted 3 October 2005 DOI: 10.1259/bjr/54848398 ’ 2006 The British Institute of Radiology

chamber (3.6 cm3) by multiple-distance measurement from 10 cm to 40 cm in a specially designed jig [5, 6]. The ESTRO guidelines recommend minimizing the positioning uncertainty and scatter contribution using a calibration jig with source positions at a distance of 10 cm from both sides of a centrally placed 0.6 cm3 Farmer-type chamber [7]. The Monte Carlo study found that the assumption of constant room scatter over the measurement distance is not valid and that in-air measurement needs correction for accurate estimation of AKS [8]. The calibration jig is designed to hold the ionization chamber and source applicator for in-air measurement. Although several factors that might affect the source calibration have been taken care of in designing the jig, the contribution of scattered radiation seems significant. The aim of the present study is to estimate the scattered radiation by a multiple-distance method using therapy level ionization chambers. The variation in scatter radiation inside the jig and from the walls of the measurement room is discussed.

Methods and materials We have used a slightly different method in the present study to estimate the scattered radiation from 347

N P Patel, B Majumdar and V Vijayan

that used by earlier experimental studies [5, 6]. In air measurement, the measured air kerma (Md) at distance d is the sum of the primary (Mp) and scatter (Ms) radiation, i.e. Md5Mp+Ms. The primary radiation (Mp) follows the inverse square law and the scatter radiation (Ms) is a constant. The product of Md and the square of the distance d is a variable quantity which is denoted by f(d) and expressed by Equation (1). f ðd Þ~Md ðd Þ2 ~ Mp zMs d 2

ð1Þ

The f(d) values for different source to chamber distances were normalized with respect to the f(d) value at minimum measurement distance and plotted against d. The scatter (Ms) shown in Equation (1) is unknown but can be determined by generating an equivalent curve from theoretical calculation. In the theoretical calculation, the air kerma from primary radiation Ep at minimum measurement distance was taken to be one and the inverse square law was used to calculate the air kerma at all other true measurement distances d95d+D, where D is the offset error in measuring the initial nominal distance d. The offset error D is due to the offset positioning of the chamber (cˇ) and source (sˇ), i.e. D5cˇ+sˇ. The variable quantities F(d) were determined using Equation (2): F ðd Þ~ Ep zEs d 2

ð2Þ

The scatter value Es was assumed constant over the range of measurement distance. The curve of normalized F(d) values vs distance d was generated by choosing the value of Es such that: h i Md ðd Þ2

plot vs: d

& Ep zEs d 2 plot vs: d

ð3Þ

then Ms5Es The condition as described in Equation (3) was accomplished when the theoretical curve F(d) was fitted to the measured normalized f(d) curve. It is evident from Equation (2) that for a constant value of Es, different curves of F(d) vs distance can be generated for primary radiations Ep with different offsets D. In the present study, nominal distances d for measurement of air kerma (Md) were chosen between 10 cm and 20 cm. Calculation was performed using Excel Workbook (Microsoft Corporation). A simple mathematical formula was derived to show the relationship between the scatter levels and ratio of air kerma for constant distance at two different positions inside the room. If r is the ratio of measured air kerma from unknown to known scatter position, air kerma from primary radiation (Mp) and the scatter at two positions are (Ms) and (Ms9), respectively, we can express the unknown scatter by Ms0 ~r Mp zMs {Mp

ð4Þ

The GammaMed Plus brachytherapy unit with an HDR 192Ir source supplied by MDS-Nordion Haan 348

Figure 1. Measurement set-up shows the jig with a 0.1 cm3 chamber and brachytherapy unit.

GmbH, Germany (at present Varian Medical Systems) was used in this experiment. Two different therapy ion chambers (volume 0.6 cm3 and 0.1 cm3) with UNIDOS dosimeter from PTW – Freiburg, Germany were used. The wall thickness and build up cap (60Co beam) of the 0.6 cm3 chamber were 0.054 g cm22 and 0.547 g cm22, respectively. The wall thickness of the 0.1 cm3 chamber without any build up cap was 0.12 g cm22. Figure 1 shows the specially designed measurement jig. The dimensions of the jig were about 30 cm 6 20 cm 6 27 cm and it was made up of non-scattering low Z materials of acrylic plates and wooden frames. Figure 2a shows the lateral view of the measurement set-up. The chamber holder was mechanically fixed at positions A and B in the jig whereas the applicator holder was shifted linearly along the track. A fine laser beam was projected over the jig to verify the sagittal, transverse and coronal cross-sections planes. Two parallel scale systems aligned with the help of laser beams were used to determine the measurement distance and the vertical position of the source applicator. A laser beam and magnifying glass were used to determine the offset position (cˇ) of the central axis of the chamber with the scale system. Figure 2b shows the positioning of the centres of the chamber and source with respect to the measuring scale. The source applicator (inner and outer diameter of 1.35 mm and 1.65 mm, respectively) can be placed at any scale pointer with an accuracy of about ¡0.01 cm. A plan view and the dimensions of the brachytherapy treatment room used in the present study are shown in Figure 3. The wall of the treatment room is made of concrete (density 2.35 g cm22). The figure indicates the location of measurement at the centre of the room and other locations used to study single and double wall scattering effects. The contribution to scattering from the walls at the centre of the room was assumed negligible. The locations of measurement for the wall effect study were at various distances from 12.5 cm to 125 cm from the single wall (square marks) and as well as from two walls (circle marks) as shown in the Figure 3. First, the jig was placed at a height of 120 cm from the floor at the centre of the room. The 0.6 cm3 chamber was fixed at position A in the jig. The source applicator was placed 1 cm from the chamber. Successive measurements were taken by vertical scanning of the source to The British Journal of Radiology, April 2006

Short communication: Scattered radiation for in-air calibration

Figure 2. (a) Lateral view of measurement set-up. The chambers 0.6 cm3 and 0.1 cm3 are placed at positions A and B,

respectively. The source position c is for the 0.6 cm3 and c’ for the 0.1 cm3 chamber that means the heights of the effective points of chambers differed by 1.4 cm. (b) The offsets in positioning of the centre of chamber (cˇ) and source (sˇ). The total offset in the setup distance is given by D52cˇ+sˇ.

determine the reference dwell position where the chamber shows maximum response. In the vertical scanning, the source was moved to dwell positions (stepping size of 0.2 cm) along the applicator. The reference dwell position where the chamber shows maximum response lies on the transverse axis that passes through the centre of the source as well as chamber. For measurement of air kerma, the source applicator was placed at measuring distance and the source was moved to the reference dwell position. Air kerma was measured for 100 s. A similar procedure was followed for measurements at other distances. The nominal distances (d) were 10 cm, 12 cm, 14 cm, 16 cm, 18 cm, and 20 cm. At each nominal distance, five source offset positions (sˇ5–0.4, –0.2, 0.0, +0.2, and +0.4 cm) were taken. Air kerma measurements at offsets of 10 cm distance were repeated as their readings were used for normalization. Hence 30 values of air kerma at different source to chamber distance were measured. Measured values were corrected for air attenuation and nonuniformity due to chamber size [9]. Transit time correction was not required as the dosemeter timer was used to collect the charge during an interval after the source stopped moving. The leakage current was negligible. A change in temperature and pressure was monitored and successive corrections were made. The offset of the chamber position (cˇ) was determined to calculate the total offset (D) as shown in Figure 2b. The scattered radiation Es was estimated as per the methodology described in the text using a computer. A similar procedure was followed to estimate the scatter for the 0.6 cm3 chamber at position B. In the case of the 0.1 cm3 chamber, a similar procedure (except the stepping distance was 0.1 cm and air kerma measurement period 200 s) was used to estimate the scatter at positions A and B. The jig was placed at different locations (circle and square marks) successively as shown in Figure 3 to study the scatter effects from the walls of the room. A similar procedure was followed to estimate the scattered radiation at every location by taking 18 air kerma measurements (only for three offset positions of the source) with the 0.6 cm3 chamber placed at position A in the jig. The scatter radiation vs distance from the single and double wall was determined. The British Journal of Radiology, April 2006

The relationship given in Equation (4) was verified experimentally. The estimated scattered radiation Ms at the centre of the room for the 0.6 cm3 chamber at position A in the jig was taken as the reference value. The estimated scattered radiation at other locations (taken as Ms9) was used to calculate the ratios of air kerma r (relative ionization) for the locations to the centre of the room. To verify the accuracy of calculated relative ionization (r), the air kerma for constant source to chamber distance was measured at the centre of the room and the other locations.

Results The maximum response of the 0.6 cm3 and 0.1 cm3 chambers was found at different source dwell positions 6.8 cm and 5.4 cm from the end of the source applicator, respectively. This means that the heights of the effective centres of the chambers from the base of the jig differed by 1.4 cm. Reproducibility (n55) of our measurement by repositioning of the source applicator was within 3%, 1%, and 0.1% at measurement distances of 1 cm, 5 cm, and 10 cm, respectively. The linearity of the dosemeter within the range of 10–200 s was 0.01% for the 0.6 cm3 chamber and 0.02% for the 0.1 cm3 chamber. Figure 4 shows the plot of normalized f(d) values from measured air kerma (dots) and theoretically derived F(d) (lines) vs nominal distance d for different offsets (D) for the 0.6 cm3 chamber at position B. It shows that the plot of measured f(d) values are best fitted by the theoretical curves of corresponding offsets (D) with varying scatter components Es of the primary at a measurement distance of 10 cm. The mean of these scattered components Es was 0.59¡0.06%. This means that the measured air kerma has a contribution from scatter (Es) of about 0.59% of the primary radiation at 10 cm distance and the scatter quantity is assumed to be constant over the range of measurement distance. The estimated scattered radiation in terms of percentage of primary radiation at 10 cm measurement distance for both the chambers at positions A and B is shown in Table 1. These results were obtained with the jig placed at the centre of the room. The uncertainty in the 349


Figure 3. Plan view of the brachytherapy treatment room. Not to scale.

positioning of chamber was about ¡0.02 cm, which resulted in uncertainty of ¡0.1% in the estimation of scattered radiation. The scatter measured by the 0.6 cm3 chamber at position A was 0.33% and when the chamber was moved inside the jig by 2 cm at position B, the measured scatter level increased to 0.59%. Similarly the scatter radiation measured by the 0.1 cm3 chamber at positions A and B was 0.58% and 1.11%, respectively. Even if the uncertainty in chamber position is taken into account, our results show a variation in the scattered radiation over the different measurement positions inside the jig. There is also a significant difference

Figure 4. The plot of normalized f(d) values from measured air kerma (dots) and theoretically derived F(d) (lines) vs nominal distance d for the 0.6 cm3 chamber at position B. The measured f(d) curves were fitted by theoretical curves F(d) of corresponding offsets (D) with varying scattered components (Es). The offsets (D) in nominal distance were +0.3 cm, +0.1 cm, 20.1 cm, 20.3 cm, and 20.5 cm, respectively. The percentage value in brackets is the scatter component. The mean of scattered components was 0.59¡0.06% at the centre of the room for the 0.6 cm3 chamber at position B in the jig. 350

between the results for the 0.1 cm3 and 0.6 cm3 chamber. It should, however, be noted that the centres of the chamber volume were located at different heights. An additional chamber holder, which was very close to the chamber volume in measurement with 0.1 cm3, could have increased the scattered contribution. Figure 5a shows the scattered radiation as a function of distance from a wall in single and double wall scattering, estimated by multiple-distance measurement using the 0.6 cm3 chamber at position A. It is observed that the scattered radiation becomes constant at about 0.33% at distances greater than about 100 cm from the walls for both single and double wall scattering. A steady increase in scattered radiation was observed below 50 cm from a single wall and 75 cm from the double wall. At 12.5 cm from the wall, scattered radiation was estimated at 2.0% and 2.7% of the primary for single and double wall, respectively. The deviation in measured relative ionization from calculated relative ionization as a function of the distance from a single and double wall for the 0.1 cm3 and 0.6 cm3 chamber is shown in Figure 5b. The calculated relative ionizations were determined from the estimated scattered radiation at different locations using the 0.6 cm3 chamber as shown in Figure 5a. Results show that calculated relative ionization is in good agreement with measured relative ionization with most of the deviations within ¡0.5% for both the chambers.

Table 1. Estimation of scattered radiation for the chambers Type of the chamber

3

0.6 cm chamber 0.1 cm3 chamber

Estimated value of scattered radiation for n55 (in % of primary radiation at d 510 cm) Position A

Position B

0.33 (¡0.04) 0.58 (¡0.05)

0.59 (¡0.06) 1.11 (¡0.04)


Short communication: Scattered radiation for in-air calibration

Figure 5. (a) Estimated scattered radiation as a function of distance from the wall in double and single wall scattering by

multiple-distance measurement using the 0.6 cm3 chamber at position A inside the jig. (b) Deviation in measured relative ionization from the calculated value as a function of distance from wall for both the chambers (DW, double wall; SW, single wall). The measurement distances were 12 cm, 16 cm, and 18 cm. The calculated relative ionizations were determined from estimated scattered radiations using the 0.6 cm3 chamber.

Discussion The chamber used in the present study was fixed and the source applicator was moved to different positions. The accuracy of positioning of the source applicator was about ¡0.01 cm. Displacement of the source inside the applicator was crucial in the measurement as it has a smaller diameter (0.9 mm) than the inner diameter of the applicator. Displacement of the source was not observed during reproducibility measurements at any reference position when performed by repositioning the applicator. However, when the applicator was moved to another position, the source was displaced laterally due to movement of the source guide tube. The maximum error in setup distance may increase by up to ¡0.032 cm as a result of the combined effects of source displacement and applicator positioning. In order to minimize the uncertainty in the estimation of the scattered radiation due to positioning error, we have considered five offset positions of the source. The mean of scattered values from five offsets was calculated. The standard deviations in the measured f values (after correction of scatter) were found to be well within 0.2% and 0.3% for the 0.6 cm3 and 0.1 cm3 chamber, respectively. Measurements were also compared with the threeequation solution technique described earlier by other authors [5, 6]. It was observed that the shape of the curve derived from Equation (3) contains a unique pair of scatter (Es) and offset (D) values. The solution technique was found to be very sensitive to measurement error. For example, a deviation of ¡0.2% in air kerma values from primary radiation (without any scatter contribution) at five distances gives the scatter value from –0.8% to 0.65% using the solution technique. However, in the present analysis, the estimated scattered radiation was found to be well within ¡0.1%. IAEA guidelines recommend that offset values determined from the solution of the three equations should not vary by more than ¡0.1 cm [3]. This limit can be retained only if the deviation in measured f values (after correction of scatter) is about ¡0.03%. Therefore, in the solution technique, separations between measurement distances must be very precise. In the present method, The British Journal of Radiology, April 2006

variation in separation is managed by curve fitting but it is essential to know the measurement distance for accurate estimation of scatter radiation. The air kerma at any point is the sum of primary and scatter radiation. If the multiple-distance method is used then constancy of scatter over the measurement distance is essential for its precise estimation. Our results show that in addition to the constant scatter contribution from the room, the components of the jig also contribute significantly to scatter, which varies from position to position inside the jig. A large difference in scatter values between positions A and B shows a high gradient scatter field over the inner to outer area of the jig. The curves in Figure 4 show that the measured f(d) values of each offset were in ‘‘best fit’’ with the theoretical curve F(d), if the scatter gradient over the central area of the jig is assumed to be low. This means the variation in scatter contribution to the chamber from multiple source positions is negligible. Therefore, a measurement set-up with a fixed chamber position should be chosen to minimize the variation in scatter inside the jig for the multiple-distance method. The optimum distance for a Farmer-type chamber to give minimum combined uncertainty effect from several factors was found at 16 cm [4]. IAEA recommends measurement at multiple distances between 10 cm and 40 cm [3]. In general, the calibration jig is required for the multiple-distance measurement, thus requiring estimation of scattered radiation for the local conditions. In case of the Farmer-type chamber, the measurement distances should be selected around the optimum distance of 16 cm and a separation of 2 cm between distances is sufficient to give satisfactory results for the multipledistance method. If a larger range of measurement distances is used, the variation in scatter from the jig could enhance the uncertainty in estimation of scatter radiation as well as air kerma rate.

Conclusion This study suggests that a fixed chamber with changing source positions should be used in multiple-distance 351


measurement of air kerma rate when using a measurement jig. The source should be placed at different positions over a short-range around the optimum measurement distance of 16 cm for the Farmer-type ionization chamber. Scattered radiation estimated from above method remains unchanged for the reproducible measurement set-up, which makes the procedure of frequent source calibration very simple.

References 1. American Association of Physicists in Medicine. Specification of brachytherapy source strength. AAPM Report 21. New York, NY: AAPM, 1987. 2. Nath R, Anderson LL, Meli JA, Olch AJ, Stitt JA, Williamson JF. Code of practice for brachytherapy physics: Report of the AAPM Radiation Therapy Committee Task Group No. 56. Med Phys 1997;24:1557–98. 3. IAEA-TECDOC-1274. Calibration of photon and beta ray sources used in brachytherapy - Guidelines on standardized

352

4.

5.

6. 7.

8.

9.

procedures at Secondary Standards Dosimetry Laboratories (SSDLs) and hospitals. Vienna: IAEA, 2002. DeWerd LA, Ezzell GA, Williamson JF. Calibration principles and techniques. In: Nag S, editor. High dose rate brachytherapy. New York: Futura Publishing Company Inc., 1994. Goetsch SJ, Attix FH, Pearson DW, Thomadsen BR. Calibration of 192Ir high-dose-rate afterloading systems. Med Phys 1991;18:462–7. Stump KE, DeWerd LA, Micka JA, Anderson DR. Calibration of new high dose rate 192Ir sources. Med Phys 2002;29:1483–8. ESTRO Booklet No. 8: A Practical guide to quality control of brachytherapy equipment. Mounierlaan 83/12 – 1200 Brussels (Belgium): ESTRO 2004. Palani Selvam T, Govindarajan KN, Nagarajan PS, Sethulakshmi P, Bhatt BC. Monte Carlo aided room scatter studies in the primary air kerma strength standardization of a remote afterloading 192Ir HDR source. Phys Med Biol 2001;46:2299–315. Kondo S, Randolph ML. Effect of finite size of ionization chambers on measurements of small photon sources. Radiat Res 1960;13:37–60.



SHORT COMMUNICATION

Implications of quality adjusted survival for clinical trials in radiation oncology B JONES,

MD, FRCR, MedFIPEM

Birmingham Cancer Centre, University Hospital Birmingham, Birmingham B15 2TH, UK ABSTRACT. Clinical trials in radiotherapy sometimes compare changes in radiation dose distribution using different radiation techniques. The use of quality adjusted survival can, in special circumstances, reduce the requirement of large patient numbers in order to show a significant difference in overall outcome. The provisos are that marginal improvements in survival or tumour control endpoints and a reduction in toxicity scores are present. The converse findings would also be amenable to this approach. Random sampling methods are used to construct a patient population where the first set of conditions is met. Further work is necessary to refine the absolute indications for this technique.

The results of modern clinical trials in radiotherapy frequently show marginal changes in survival, which necessitate very large numbers of patients to demonstrate even a trend of improvement in outcomes. For example, the four-armed trial of accelerated and hyperfractionated radiotherapy conducted in the USA by Fu and colleagues [1] involved 1113 entered patients, with statistical significance probability values ranging from 0.045 to 0.067 for local tumour control and disease free survival. Separate analysis of survival and side effect statistics is the norm, often with much less detail given in terms of reporting the latter. The costs of clinical trials – in terms of human and financial resources provided by the government and charitable organizations, as well as the restrictions imposed upon patients – are considerable in advanced countries such as the UK. Better ‘‘value for money’’ would be achieved if trial costs could be curtailed – this would at least allow more trials to be conducted. The concept of quality adjusted survival is not new [2] and could be more frequently applied in clinical trial analysis, although it is most frequently used in cost benefit studies by health economists. Clinical trials in oncology are usually concerned with the duration of local tumour control, survival times (both continuous variables) and the severity of side effects (normally graded as discrete variables). In the testing of new radiation therapy techniques, for example if dose distribution is changed while delivering the same total tumour dose, it might be that the same or similar local control and survival can be expected but that the quality of life may be the intended benefit [3]. Another example would be the delivery of a slightly higher radiation dose to the tumour while at the same time reducing the dose to normal tissues by means of charged particle beams [4], with a modest increase in tumour control and a reduction in side effects. There are potential ethical difficulties about the conduct of such trials where the dose distributions (obtained using the predictions of the laws of physics) The British Journal of Radiology, April 2006

Received 23 May 2005 Revised 27 July 2005 Accepted 7 October 2005 DOI: 10.1259/bjr/15389891 ’ 2006 The British Institute of Radiology

are judged to be so much better in the case of the new treatment [5]. However, past experience with ‘‘new’’ forms of radiotherapy required trials to demonstrate that theoretically predicted gains were not realised in practical situations [6, 7], although in these studies there was no change in dose distribution. There is also a dilemma when new treatments are phased in slowly owing to reduced treatment capacity: in such cases there might be opportunities to do randomized studies. It would be advantageous if subtle changes in tumour control, survival and simultaneous improvements in quality of life could be found from trials that incorporate smaller numbers of patients; results should be available sooner and the financial costs of trials would be considerably reduced. The present paper aims to demonstrate the potential advantage of quality-adjusted survival in a simulated clinical trial where the necessary conditions of marginal improvements of disease free survival and reduced complications occur.

Methods The disease free survival times of a cancer patient population are simulated by random sampling techniques for the two arms of a trial that compares conventional X-ray therapy (XRT) with charged particle beam (CPB) therapy given to slightly higher dose. The assumptions made are given in Table 1. This example is not meant to advance the cause of any particular form of therapy, but merely used to demonstrate the statistical principles. Mathematica and GraphPad Prism software are used to obtain the results. The quality-adjusted survival is calculated by multiplying the actual survival by a factor F defined as: F~

5{grade 5 353

B Jones

Discussion

Table 1. Assumptions made in modelled population

Mean survival in years with standard deviation in parentheses for a log-normal distribution Number of ‘‘cured’’ 10 year survivors not included in above distribution Grade 0 toxicity % Grade 1 toxicity % Grade 2 toxicity % Grade 3 toxicity % Grade 4 toxicity %

XRT

CPB

4 (2.7)

4.5 (3)

23

30

40 40 15 4 1

60 30 7 2 1

XRT, conventional X-ray therapy; CPB, charged particle beam therapy.

In this way, toxicity scores of 0, 1, 2, 3 and 4, respectively, have F51, 0.8, 0.6, 0.4, 0.2. The assumptions made are given in Table 1, including the cumulative toxicity scores. For the results shown in Figure 1, the survival log rank test provides p50.062 for a one tailed study. The side effect profile is separately analysed to give x2 (with four degrees of freedom)59.04, p50.061. When the qualityadjusted survival is assessed, then p,0.001, which is a substantial reduction. Essentially this is the same approximate result as would be obtained by multiplying the two probabilities, i.e. the probability of the null hypothesis being accepted for the uncorrected survival and for the side effect profile being the same in both arms of the trial (p50.061 6 0.001¡le;0.001). Rather than use the classical x2 test with n–1 degrees of freedom, where n is the number of categorical variables, the x2 test for trend can be used with only one degree of freedom, which inevitably reduces the p-value. By use of this test, the results given in Table 1 yield a p50.003, which is significant without combining with the disease free survival results. However, the necessary condition for the use of this test is that the outcome variable must be well ordered; this condition will not be satisfied in trials where there might be less grade 0 than grade 1 toxicity.

Figure 1. Kaplan-Meier plot of a simulated clinical trial comparing conventional X-ray therapy (XRT) with charged particle beam (CPB) therapy. There are 100 patients in each treatment arm and the assumptions made are given in Table 1. 354

At present there is some dissatisfaction with the constraints imposed by standard clinical trials in radiotherapy [8]. A wider spectrum of research methods, such as observational studies and mathematical modelling, is required to solve at least some of the problems confronted in a complicated discipline such as radiation oncology, where trials may not be appropriate, or considered unethical [9]. The relatively simple exercise presented in this paper demonstrates proof of principle that fusion of survival with quality of life end points may very substantially reduce the numbers of patients required to show that a new treatment confers significant overall benefits. The number of patients required to confirm marginal benefits in radiation oncology is discussed further by Bentzen [10]. To use the quality adjusted approach, certain safeguards would be required: e.g. that the ‘‘survival’’ trend is superior, with a suggested p-value of ,0.1; also that the side effect profile is directionally correct, i.e. that there would be reduced numbers of toxicity scores in each toxicity category .0 for the test treatment with statistical confirmation of a trend with p,0.1 Further sensitivity analysis along these lines should be performed by expert statisticians, with more complex calculations of the expected numbers required to achieve the required level of statistical confidence using the adjusted survival [10, 11]; recommendations of acceptable p-value ranges from the primary end point analyses might also be helpful. Quality adjusted survival must be regarded as either a primary or a secondary endpoint, to be used in special circumstances: this technique might even have a place where survival is unchanged, since further endpoints such as local tumour control or disease free survival may show significant differences when adjusted for quality of life. The quality survival concept has been applied elsewhere in oncology and some statisticians have already advocated this approach [11]. For example, in leukaemia [12], where four clinical states (viz. toxicity, treatment free of toxicity, no treatment nor symptoms, relapse) were defined. The average times spent in each state were weighted by utility coefficients that reflect relative value according to quality of life. This represents a much more complex process, which also acknowledges the reversible nature of some of the treatment-related side effects. However, in radiotherapy the serious side effects tend to be chronic and cumulative with time; consequently, such a sophisticated approach would seem to be unnecessary. The use of extant grading systems in oncology can be used in toxicity grade allocation [13] and in crude quality adjusted survival. Actuarial assessment of each side effect grade could be used, but again the numbers of patients and events may be so small that no significant differences would be found, as would be the case for grades 3–4 in Table 1. The present emphasis on obtaining improved survival in clinical trials ignores the quest for testing improved quality of life. The main funding agencies do not normally support or encourage: primary quality of life endpoint studies in which there is no expectation of enhanced tumour control; observational studies of more The British Journal of Radiology, April 2006

Short communication: Quality adjusted survival

complex or newer forms of therapy; large multicentre non-randomized studies that allow convergence of technique and good quality assurance as a prerequisite for multicentre trials. Testing of new radiotherapy techniques – such as charged particle beams – that are specifically designed to reduce side effects, yet either maintain or increase tumour control, might need to utilize quality based end points in trial analysis given that the conventional clinical trial end points may not be sufficiently sensitive where subtle changes in multiple outcomes occur. It is hoped that this paper will stimulate discussion and lead to better and more cost effective studies within the UK National Cancer Research Institute as well as in other countries.

References 1. Fu KK, Pajak TF, Trotti A, Jones CU, Spencer SA, Phillips TL, et al. A Radiation Therapy Oncology Group (RTOG) phase III randomized study to compare hyperfractionation and two variants of accelerated fractionation to standard fractionation radiotherapy for head and neck squamous cell carcinomas: first report of RTOG 9003. Int J Radiat Oncol Biol Phys 2000;48:7–16. 2. Nord E. Cost–value analysis in Health care: making sense out of QALY’s. Cambridge University Press, 1999. 3. Dearnaley DP, Khoo VS, Norman AR, Meyer L, Nahum A, Tait D, et al. Comparison of radiation side-effects of conformal and conventional radiotherapy in prostate cancer: a randomised trial. Lancet 1999;353:267–72.


4. Jones B, Rosenberg I. Particle therapy Cooperative Oncology Group (PTCOG40), Institute Curie 2004. Br J Radiol 2005;78:99–102. 5. Suit H, Goldberg S, Niemerko A, Trofimov A, Adams J, et al. Proton beams to replace photon beams in radical dose treatments. Acta Oncologica 2003;42:800–8. 6. Errington RD, Ashby D, Gore SM, et al. High energy neutron treatment for pelvic cancers: study stopped because of increased mortality. Br Med J 1991;302:1045–51. 7. Maor MH, Errington RD, Caplan RJ, et al. Fast neutron therapy in advanced head & neck cancer: a collaborative internal randomised trial. Int J Radiat Oncol Biol Phys 1995;32:599–604. 8. Munro AJ. The conventional wisdom and the activities of the middle range. Br J Radiol 2005;78:381–3. 9. Jones B, Dale RG, Carabe A. Conventional wisdom and activities of the middle range. Br J Radiol 2005;78:1119. 10. Bentzen SM. Towards evidence based radiation oncology: improving the design, analysis, and reporting of clinical outcome studies in radiotherapy. Radiother Oncol 1998;46:5–18. 11. Billingham LJ, Abrams KR. Simultaneous analysis of quality of life and survival data. Stat Methods Med Res 2002;11:25–48. 12. Levy V, Porcher R, Delabarre F, Leporrier M, Cazin B, Chevret S. French Cooperative CLL Group. Evaluating treatment strategies in chronic lymphocytic leukemia: use of quality-adjusted survival analysis. J Clin Epidemiol 2001;54:747–54. 13. Davidson SE, Burns MP, Routledge JA, Swindell R, Bentzen SM, West CM. Assessment of morbidity in carcinoma of the cervix: a comparison of the LENT SOMA scales and the Franco-Italian glossary. Radiother Oncol 2003;69:195–200.

355


Correspondence (The Editors do not hold themselves respondible for opinions expressed by correspondents)

Response to ‘‘Radiation dose measurement and optimization’’ The Editor—Sir, In his Letter to the Editor on ‘‘Radiation dose measurement and optimization’’ in the September 2005 issue of this journal [1], Dr Moores raised a number of rather paradoxical objections to the use of dose–area product (DAP) measurements for monitoring radiation doses to patients in diagnostic radiology. His arguments that DAP meters are unsuitable because: (i)

they do not measure absorbed dose to a specified medium, (ii) they combine two quantities (entrance surface dose and entrance field area) into one, where the contributions from each are indistinguishable, (iii) male patients have much higher DAP values on average than female patients, are based on a fundamental misconception of the role of diagnostic reference levels (DRLs) in the optimization of medical exposures. DRLs are essentially investigation levels that, if exceeded, trigger a detailed investigation into the causes of unusually high exposures. They are just the first step in the optimization process – a means for identifying those situations in most urgent need of investigation because the doses appear to be towards the top end of the distribution seen on a national scale. The choice of DAP as one of the quantities in which DRLs are expressed is based mostly on the practicability of making DAP measurements during radiographic and fluoroscopic examinations in the clinical environment and the widespread installation of DAP meters on X-ray imaging equipment throughout the health service. There are three essential requirements for the dose quantities to be used in DRLs; they should be unambiguously defined so that everyone clearly understands what is to be measured, they should be capable of simple, direct measurement (or calculation from indirect measurements) with readily available dosemeters of sufficient precision and accuracy, and they should provide an indication of the typical dose received by patients examined in a particular facility from a particular type of X-ray examination. Clear guidance on how DAP meters can be used to meet all these requirements is provided in the National Protocol for Patient Dose Measurements in Diagnostic Radiology [2] and in Guidance on the Establishment and Use of DRLs for Medical X-ray Examinations [3] published by IPEM. Other practical dose quantities will be needed for expressing DRLs in different imaging modalities such as CT, mammography and dental radiography. These have been clearly defined in many publications, including IPEM report 88 [3]. It is, therefore, rather puzzling to see Dr Moores describing this wide variety of dose quantities 356

as an intolerable situation, most likely due to ‘‘the lack of a clear understanding of the meaning and relevance of the term ‘‘optimization’’ in the field of diagnostic radiology’’. Dr Moores is correct in pointing out that the full optimization process (i.e. maximizing the ratio of benefit to harm) in diagnostic radiology, requires assessment of both the diagnostic efficacy of the examination and the resulting radiation risk. However, this is the purpose of the detailed investigation that is triggered by exceeding the DRL; it was never intended that the dose measurements made to determine whether the DRL is being exceeded should, by themselves, provide all the information necessary for complete optimization. One of the first investigations to carry out on exceeding a DRL might well be to see if the sample of patients on whom DAP measurements were made was biased towards excessively large patients, since the DRLs are set for average-sized patients. The fact that males tend to be generally larger than females and consequently will on average have higher DAP values, is no reason, per se, for not using DAP measurements to trigger such investigations. A further investigation might be necessary to determine whether the excessive DAP values were due to the use of large entrance surface doses or large field areas (or both). Optimization in diagnostic radiology does not mean simply maximizing image quality and minimizing patient dose, rather it requires radiologists to determine the level of image quality that is necessary to make the clinical diagnosis and then for the dose to be minimized without compromising this image quality. Radiologists, of course, subjectively assess the adequacy of their images every time they report on them, and it can be argued that no one else is in a better position to do so. What radiologists cannot do intuitively is to assess the patient dose, which is why practical patient dose monitoring techniques and DRLs are needed to trigger more detailed optimization investigations, if images that are locally considered to be ‘‘diagnostically acceptable’’ are being obtained with excessively high patient doses. Dr Moores’ plea for the development of ‘‘a scientific framework for the quantification of diagnostic outcome and therefore clinical benefit’’ will, no doubt, have an important role in these detailed optimization studies. But it is difficult to see how such a framework would eliminate the ‘‘general ignorance within the medical profession concerning the levels of radiation employed in diagnostic radiology’’, as he claims. Rather, the establishment and implementation of DRLs using practical dose quantities such as DAP, has probably done more to raise awareness of patient doses in the radiology professions than any other national initiative. The halving of the DRLs that has been achieved for many common types of X-ray examination in the UK over the past 20 years [4] with no apparent loss of diagnostic The British Journal of Radiology, April 2006

Correspondence

efficacy, provides ample evidence for the effectiveness of this simple first step in the optimization process. Yours etc., B F WALL Radiation Protection Division Health Protection Agency Chilton Didcot Oxon OX11 0RQ UK (Received 18 October 2005 and accepted 25 October 2005) DOI: 10.1259/bjr/34749787

References 1. Moores BM. Radiation dose measurement and optimization. Br J Radiol 2005;78:866–8. 2. Dosimetry Working Party of IPSM. National Protocol for Patient Dose Measurements in Diagnostic Radiology. Chilton, NRPB, 1992. 3. IPEM Report 88. Guidance on the Establishment and Use of DRLs for Medical X-ray Examinations, York, IPEM, 2004. 4. Hart D, Hillier MC, and Wall BF. Doses to Patients from Medical X-ray Examinations in the UK – 2000 Review. NRPB-W14, www.hpa.org.uk/radiation, 2002 [Accessed 25 November 2005].

Authors’ reply The Editor—Sir, I wish to thank B Wall for his comments on my letter and I would like to respond, first of all, to paragraphs 4 and 5 of his letter. Here, he appears to recommend an unproven optimization strategy for female patients, which is based upon the use of dose reference levels (DRLs) that are dependent upon the exposure of male patients. I believe that such an approach severely disadvantages female patients. It either ignores or neglects the well-known and significant differences in both entrance surface dose (ESD) and dose–area product (DAP) values for female patients, compared with those of males, as well as their own distinct optimization requirements. A comparison of male and female ESD and DAP values measured during a national survey of doses to patients in the UK [1] and the corresponding DRLs presented in the National Protocol for Patient Dose Measurements [2] indicates that National and European adult patient DRLs are governed by the levels of exposure to male patients. The reason for this lies, to a large extent, in the recommendation that the mean weight of an adult patient sample, in any ESD/DAP audit, should lie within 5 kg of 70 kg [2]. In the national patient dose survey [1], the average weight of female patients was roughly 60 kg whilst that of the male patients was 70 kg. Similar weight differences for female and male patients have also been noted in studies involving fluoroscopy [3] and radiography [4]. Also, the weights of female and male adult voxel phantoms, developed for computational radiation protection studies, have been designed around similar The British Journal of Radiology, April 2006

weight differences [5–7]. The decision to employ a singleweight dependent DRL value has never been scientifically justified as an effective basis for optimization of practices in respect of female patients. As I indicated in my letter, results of the 1985 national survey [1] demonstrate that the mean ESD for female patients undergoing a PA chest X-ray examination was 30% lower than that of the male patients. Such differences in ESD were also noted, for example, for the AP abdomen (27%), lat thoracic spine (33%), lat lumbar spine (25%) examinations. Other studies have also shown similar differences in ESD [4], for PA chest (70%), knee (100%) and AP skull (70%), examinations. The mean DAP value for female patients was significantly lower than that of male patients [1] for the PA chest (86%), thoracic spine (22%), barium meal (39%), IVU (43%), cholangiography and cholecystography (43%) examinations. Other studies have shown similar differences in DAP values for female patients [3]; barium meal (36%), barium swallow (45%) etc. In the case of the DAP values, such differences can arise from unknown variations in the ESD and field sizes employed. Under present circumstances, local DRLs derived for patients in a women’s hospital would not be relevant to females in a mixed-sex patient group of 70 kg average weight. Different audit standards, therefore, can exist for different female populations. Similarly, female DAP values for a PA X-ray examination of the chest may exceed their own population norm by up to 86% [1], due to unknown variations in either ESD or field size, before the need for an investigation is indicated. Such an indication may be dependent upon the patient mix, including weights, in the audit sample, as well as specifically radiation protection considerations. Why should easily demonstrable, optimized practices for female patients be hidden under patient sampling/weight considerations involving male patients? Besides exposure differences for female patients, other factors merit consideration in respect of optimization strategies for female patients:

N N N N N N N N

For many examinations, the female breast may, or could potentially, be irradiated, depending upon field size and projection, as well as technique. A large percentage of female patients have the potential to be pregnant, which may then involve the exposure of radiosensitive third parties. Computed effective doses per unit of entrance surface air kerma (ESAK) are significantly different, based upon voxel phantom studies [7]. Optimization of tube voltage and dose for digital chest radiography are different [8]. The relative frequencies of different types of examinations are different for female patients [9]. The distribution of examinations contributing to the genetically significant dose (GSD), are different [9]. The risks of fatal cancer attributable to fluoroscopy are different [10]. Organ doses during CT examinations are different [11].Parent not defined

Given such extensive differences, why are female patients not considered to be a separate exposed group 357

Correspondence

with its own DRL values, so that their own unique safety and imaging needs can demonstrably be given full and direct consideration in respect of radiation safety? The main reason for the widespread installation of DAP meters on X-ray imaging equipment throughout the NHS [paragraph 2 of B Wall’s comments] is due to the fact that funds were made available from central government (presumably on the advice of the NRPB), in the early 1990s, for the purchase and distribution of 200 DAP meters. Prior to this, extensive medical physics effort had been expended on developing the use of calibrated X-ray source data for patient dose assessments. This effort included the background science [12] and calibration methods [13]. Such an approach harmonises patient dose assessments in both radiotherapy and diagnostic radiology and is based upon over 50 years of scientific effort. Unfortunately this was ignored in the National Protocol [2]. The reduction in patient doses in diagnostic radiology over the past 20 years [paragraph 7 of B Wall’s comments] is due to the introduction of rare earth intensifying screens in radiography and sodium iodide phosphors in fluoroscopic units. These initiatives have been supported by the introduction of low attenuation materials in table-tops and Bucky stands. The UK national dose audits have faithfully recorded their well-documented effects. Our manufacturer colleagues are to be commended upon their contribution to patient dose reduction. Yours etc.,

2.

3.

4.

5.

6.

7.

8.

9.

B M MOORES Co-ordinator Radiological Unification Strategies [RADIUS] Group EC Radiation Protection Research Programme (Received 18 November 2005 and in final form 4 January 2006, accepted 11 January 2006) DOI: 10.1259/bjr/71189157

References 1. Shrimpton PC, Wall BF, Jones DG, Fisher ES, Hillier MC, Kendall GM. A national survey of doses to patients

358

10.

11.

12.

13.

undergoing a selection of routine x-ray examinations in English hospitals. NRPB-R200. Chilton, UK: National Radiological Protection Board, 1986. National protocol for patient dose measurements in diagnostic radiology. Chilton, Didcot: National Radiological Protection Board, 1992. Rowley KA, Hill SJ, Watkins R, Moores BM. An investigation into the levels of radiation exposure in diagnostic examinations involving fluoroscopy. Br J Radiol 1987;60: 167–73. Gallini RE, Belletti S, Berna V, Giugni U. Adult and child doses in standardised x-ray examinations. Radiat Prot Dosim 1992;43:41–7. Zankl M, Fill U, Petoussi-Henss N, Regulla D. Organ dose conversion coefficients for external photon irradiation of male and female voxel models. Phys Med Biol 2002;47:2367–85. Kramer R, Vieira JW, Khoury HJ, Lima FRA, Fuelle D. All about MAX: a male adult voxel phantom for Monte Carlo calculations in radiation protection dosimetry. Phys Med Biol 2003;48:1239–62. Kramer R, Khoury HJ, Vieira JW, Loureiro ECM, Lima VJM, Lima FRA, et al. All about FAX: a female adult voxel phantom for Monte Carlo calculation in radiation protection dosimetry. Phys Med Biol 2004;49:5203–16. Pascoal A, Patel R, Lawinski CP, Tabakov S. Optimization of tube voltage and dose for digital chest radiography – a study addressing patient size. In: Proceedings of UK Radiological Congress, 6–8 June, 2005. London, UK: British Institute of Radiology, 2005:2. Wall BF, Rae S, Darby SC, Kendall GM. The NRPB survey: methods and results. Dosimetry in diagnostic radiology, CRS 40. Chapter 6, 44–55. The Hospital Physicists Association 1984. Goodenough DJ. Lessons learned in radiology. Proceedings of an international conference held in Malaga, Spain, 26–30 March 2001, 145–155. International Atomic Energy Agency, 2001. Faulkner K, Moores BM. Radiation dose and somatic risk from computed tomography. Acta Radiologica 1987;28: 483–8. Harrison RM. Dose measurement in vivo and in vitro. CRS 40. Chapter 2, 8–17. The Hospital Physicist’s Association. 1984. TGR 32 Measurement of the performance characteristics of diagnostic X-ray systems used in medicine. Part 1: X-ray tubes and generators. The Hospital Physicist’s Association. 1984.



CASE OF THE MONTH

The angel in the marble 1

D J A BUTTERISS,

BSc, MBBS (Hons), FRCR,

2

M CLARKE,

FRCS, FRCOphth

and 1D BIRCHALL,

MB, BChir, FRCR

1

Department of Neuroradiology, Regional Neurosciences Centre, Newcastle General Hospital, Newcastle upon Tyne and 2Ophthalmology Department, Royal Victoria Infirmary, Newcastle upon Tyne, UK

Received 16 May 2005 Accepted 27 June 2005 DOI: 10.1259/bjr/33391462 ’ 2006 The British Institute of Radiology

A 56-year-old woman presented to the ophthalmology department complaining of subacute painless decrease in visual acuity in her right eye. She had a previous history of longstanding poor visual acuity in the left eye, which

had not been formally investigated. No other significant history was elicited. Fundoscopy revealed atrophy of the left optic disc and swelling of the right disc. No other significant abnormality was found. CT (Figure 1) and MRI (Figure 2) were performed on this patient. What are the abnormalities? What is the diagnosis in this case? What are the potential causes of optic atrophy in this condition?

Address correspondence to: Dr D Birchall, Department of Neuroradiology, Regional Neurosciences Centre, Newcastle General Hospital, Westgate Road, Newcastle upon Tyne NE4 6BE, UK.

(a)

(b)

(c)

Figure 1. (a) CT shows buckling of the right optic nerve and diffuse calvarial hyperostosis. (b) There is generalized hyperostosis of the skull base. (c) There is bilateral narrowing of the optic canals and superior orbital fissures.


359

D J A Butteriss, M Clarke and D Birchal

(a)

(b)

(c)

(d)

Figure 2. (a) Axial T2 image shows dilatation of the optic nerve sheath on the right, and optic nerve buckling. Diffuse hyperostosis is again evident. (b) Coronal short tau inversion recovery (STIR) imaging through the orbits confirms optic nerve sheath dilatation on the right, with increased signal within the nerve itself. The left optic nerve is atrophic. (c) More posterior the coronal STIR image shows compression of the optic nerves within the optic canals (arrows). (d) Sagittal T2 image demonstrates skull base and upper cervical hyperostosis, with distortion of the cervico-medullary junction and of the pituitary fossa.

Imaging findings CT demonstrates marked thickening and sclerosis of the calvarium and skull base. The optic canals are 360

stenosed bilaterally, particularly on the left. The right optic nerve is buckled. The intracranial contents are otherwise normal. In the absence of other causes of diffuse hyperostosis, these appearances are consistent The British Journal of Radiology, April 2006

Case of the month: The angel in the marble

with osteopetrosis. MRI shows left sided optic atrophy. On the right, there is dilatation of the optic nerve sheath and increased T2 signal within the optic nerve, appearances consistent with decompensating subacute optic nerve compression at the orbital apex. Sagittal imaging shows clival and skull base hyperostosis with distortion of the medulla and proximal cervical spinal cord, and compression of the pituitary gland.

Discussion Osteopetrosis, also known as Albers-Schonberg disease or marble bone disease, is an uncommon genetic disorder characterized by increased bone thickness and density, and is caused by impaired osteoclast maturation or function resulting in decreased bone resorbtion and modelling. Several forms have been identified, the most common being the infantile autosomal recessive type and the delayed autosomal dominant type. The recessive form is also known as malignant osteopetrosis and presents in early childhood with fractures and the sequelae of pancytopaenia due to marrow space obliteration. Neurological complications are common, and are secondary to raised intracranial pressure (ICP) due to calvarial expansion, and narrowing of the cranial foramina resulting in cranial nerve, spinal cord and vascular compression. The delayed form usually presents in early adulthood, although it demonstrates increasingly early onset in affected familial series. It may be asymptomatic and diagnosed incidentally on radiological investigations, but usually presents with easy fracturing or with neurological complications such as raised ICP or cranial nerve palsies. It is divided into two forms based on radiological and clinical criteria: type 1 demonstrates diffuse calvarial thickening and sclerosis with diffuse sclerosis of the axial skeleton; type 2 has less marked calvarial thickening, but the skull base is more severely affected than in type 1 and there is vertebral endplate sclerosis and bone-within-a-bone appearance [2, 3].


Overall, the most common neurological complication of osteopetrosis is visual loss and optic atrophy. There are three possible causes of visual loss [4]: (1) Optic nerve compression secondary to stenosis of the optic canals. (2) Papilloedema due to raised ICP. (3) Primary retinal degeneration. In this report, the likely cause of visual loss and optic atrophy is stenosis of the optic canals. The late presentation is an unusual feature of the case. Secondary optic atrophy is treatable, management dependent on the cause. Optic canal stenosis is treated by orbital deroofing, and papilloedema treated by reducing ICP, either by ventricular shunting, or by craniectomy alone or in conjunction with cranial vault expansion [5].

Acknowledgment Michelangelo, ‘‘I saw the angel in the marble and carved until I set it free.’’

References 1. Hall CM. International Nosology and Classification of Constitutional Disorders of Bone (2001). Am J Med Genet 2002;113:65–77. 2. Anderson PE, Bollerslev J. Heterogeneity of autosomal dominant osteopetrosis. Radiology 1987;164:233–5. 3. Cure JK, Key LL, Goltra DD, VanTassel P. Cranial MR imaging of osteopetrosis. AJNR Am J Neuroradiol 2000;21: 1110–5. 4. Stewart CG. Neurological aspects of osteopetrosis. Neuropathol Appl Neurobiol 2003;29:87–97. 5. Vanier V, Miller NR, Carson BS. Bilateral visual improvement after unilateral optic canal decompression and cranial vault expansion in a patient with osteopetrosis, narrowed optic canals, and increased intracranial pressure. J Neurol Neurosurg Psychiatry 2000;69:405–6.

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May 2006 Volume 79 Issue 941

May 2006, Volume 79, Issue 941 PET poised to alter the current paradigm for response assessment of non-Hodgkin’s lymphoma ●

Acute adverse reactions to magnetic resonance contrast media – gadolinium chelates ●

Observer variation in the evaluation of lumbar herniated discs and root compression: spiral CT compared with MRI ●

Initial experiences of using an automated volumetric measure of breast density: the standard mammogram form ●

Occupational radiation doses in interventional cardiology: a 15-year follow-up ●

A four-dimensional computer simulation model of the in vivo response to radiotherapy of glioblastoma multiforme: studies on the effect of clonogenic cell density ●

A quantitative study of IMRT delivery effects in commercial planning systems for the case of oesophagus and prostate tumours ●

●

Effect of ticlopidine in the prevention of radiation enteropathy

Radioprotective effects of hesperidin against gamma irradiation in mouse bone marrow cells ●

Characteristics and predictive factors of early-onset diarrhoea during pelvic irradiation ●

Optical density variations in CT films and their effect on image quality ●

Assessment of environmental disturbances to the static magnetic field in magnetic resonance installations ●

An investigation of search pattern extent in the threshold contrast detection task ●

Imaging pulmonary embolism in pregnancy: what is the most appropriate imaging protocol? ●

Primary subcutaneous sacrococcygeal ependymoma: a case report and review of the literature ●

●

Correspondence

●

A catastrophic complication


COMMENTARY

PET poised to alter the current paradigm for response assessment of non-Hodgkin’s lymphoma 1

A GUERMAZI,

MD

and 2M E JUWEID,

MD

1

Synarc Inc. and University of California, San Francisco, CA and 2University of Iowa Hospitals and Clinics, Iowa City, IA, USA

Received 14 September 2005 Revised 16 December 2005 Accepted 16 January 2006 DOI: 10.1259/bjr/31368041 ’ 2006 The British Institute of Radiology

Non-Hodgkin’s lymphoma (NHL) is a group of heterogeneous tumours originating in lymphoid tissue. The outcome of patients with lymphoma is, in general, better than non-lymphoid malignancies. Indeed, aggressive NHLs, which make about two-thirds of the NHL subtypes, are potentially curable (about 60% of patients are cured) with conventional and high-dose therapy. More than 70% of patients with newly diagnosed lymphoma respond well to chemotherapy alone or to chemoimmunotherapy with or without radiotherapy. In addition to the availability of new treatments, accurate staging and risk stratification with appropriate selection of treatments, as well as improved assessment of response to therapy, have resulted in a higher success rate in the treatment of lymphoma. The increasing number of treatment options available for patients with lymphoma currently being tested in clinical trials clearly requires the availability of standardized guidelines for assessment of response to therapy, which should ensure comparability of success rates among these trials. The International Workshop Criteria (IWC) for response assessment of NHL proved to be quite useful for standardized assessment of response to various lymphoma treatments tested in a significant number of clinical trials [1]. These 1999 criteria rely on clinical, biochemical, histopathological and imaging studies, especially CT. This first attempt at standardization of response criteria in NHL proved to be a significant step forward; yet with the advent of functional imaging with positron emission tomography (PET), particularly using fluorodeoxyglucose (FDG), the limitations of these criteria have become increasingly apparent. Six years after their publication in the Journal of Clinical Oncology, it was clear that the IWC were in need of Address correspondence to: Ali Guermazi. MD, Synarc Inc., 575 Market Street, 17th Floor, San Francisco, CA 94105, USA.

The British Journal of Radiology, May 2006

revision. Dr Bruce Cheson from Georgetown University in Washington DC, who organized the first International Workshop to Standardize Response Criteria for NHL in 1998, organized a meeting together with Dr Volker Diehl from the University of Cologne, Germany, called the International Harmonization Project (IHP), at which an invited international group of participants with clinical, histopathological and imaging expertise in NHL deliberated on the need for harmonization of clinical trial parameters and revision of the currently utilized response criteria in lymphoma. The meeting took place in June 2005 at the 9th International Conference on Malignant Lymphoma in Lugano, Switzerland. Six subcommittees were formed — response criteria, clinical features, endpoints, pathology/immunohistology, paediatrics and imaging. Each subcommittee presented the results of their deliberations. The imaging subcommittee, including the authors of this commentary among others, stressed the importance of the introduction of new technologies, such as PET and PET/ CT. Indeed, one of the most vexing clinical dilemmas in response assessment and management of NHL is the detection of active disease in a residual mass by conventional imaging (usually CT), following treatment despite resolution of clinical symptoms and normalization of laboratory tests (i.e. in patients with an otherwise clinical complete response). Such masses, usually seen in the mediastinum and abdomen/pelvis, may be seen in up to 40% of patients with aggressive NHL and an even higher percentage of patients with Hodgkin’s disease. Yet the vast majority of these masses represent necrosis and/or fibrosis and only a small percentage represents or contains residual active lymphoma. It is also noteworthy that the frequency of occurrence of these residual masses correlates with the size of tumour mass(es) at diagnosis, with a much more common occurrence of residual masses in patients with initially bulky compared with non-bulky lymphoma. 365

A Guermazi and M E Juweid

Obviously, characterization of these residual masses is important: residual lymphoma portends a poor prognosis and requires further salvage therapy, while an inactive residual mass usually confers an excellent prognosis without further treatment. Residual masses can be assessed with biopsy/fine needle aspiration or even surgery. However, these approaches, while generally accurate, are quite invasive and associated with some morbidity. Non-invasive and yet accurate approaches for assessment of these masses are clearly warranted. Conventional morphologic imaging (e.g. CT and conventional MRI) alone do not allow discrimination between active tumour and fibrosis and/or necrosis in residual masses. Most CT studies report poor specificity in the characterization of residual masses (40–50%). Conventional MRI has been shown to be slightly more reliable in this setting, but its performance is still unsatisfactory. Gallium scintigraphy (GS) has been shown to be substantially more accurate than CT and MRI for characterization of these masses, particularly in patients with aggressive lymphoma, with less reliable results in low-grade or follicular lymphoma that are substantially less gallium-avid. Importantly, GS is not well standardized and is also not a convenient imaging modality for both the patients and referring physicians because of the frequent requirement for imaging at multiple time points over several days to more than a week. The limited availability of the gallium-67 isotope, because of limited production and distribution, poses another significant problem. All of these issues in addition to the only moderate accuracy (about 60–70%) resulted in a relatively limited utilization of GS in both the USA and Europe. The advent of PET promises to solve many of the problems associated with the use of GS, being much more convenient to use with less than 2 h required for a PET study to be completed. More importantly, PET clearly has higher accuracy than GS and morphologic imaging for assessment of most subtypes of lymphoma, including follicular histology. PET alone or coupled with CT (PET/ CT) has demonstrated a high accuracy (i.e. . 80%) for assessment of response to therapy in patients with aggressive NHL, particularly with respect to characterization of residual masses, and it is now generally recognized that PET is the most accurate non-invasive method for differentiating necrosis and/or scar from viable tumour in residual masses [2–6]. Largely because of the recognized high accuracy of PET in characterizing the nature of residual masses following treatment of lymphoma, the IHP workshop has unequivocally agreed that PET or PET/ CT should be recommended in assessment of response or restaging of aggressive NHL and Hodgkin’s lymphoma (HL) at the completion of therapy. This decision is supported by several persuasive studies by various groups of investigators who evaluated the predictive value of PET following first-line therapy of patients with aggressive NHL and/or HL compared with CT or conventional staging methods [3–5]. More recently, Juweid et al [6] went one step further and compared a proposed response classification based on integration of FDG-PET into the IWC (IWC+PET) with the IWC alone in patients with aggressive NHL. This retrospective study included 54 patients with aggressive NHL (87% diffuse large cell) who underwent PET and CT 1–16 weeks 366

(median 4 weeks; 83% within , 3–11 weeks) after 4–8 cycles of chemotherapy (84% received 6–8 cycles) mostly with cyclophosphamide, doxorubicin, vincristine and prednisone (CHOP) with or without rituximab. These patients were assessed for complete response (CR), unconfirmed complete response (CRu), partial response (PR), stable disease (SD) and progressive disease (PD) by the IWC and by specifically defined IWC+PET-based response designations, derived by integrating the PET findings into the previously defined IWC designations. Progression-free survival (PFS) was used as the primary outcome measure to compare the IWC-assigned and IWC+PET-assigned response designations. The investigators found that there was a strikingly higher proportion of patients with CR by IWC+PET compared with IWC alone (35/54 versus 17/54, respectively), with no patients designated as CRu by IWC+PET compared with 7 with IWC alone, because such patients were either designated as CR if PET was negative (fibrosis/necrosis) or PR if PET was positive (residual tumour). This latter finding is significant since, in fact, several studies have shown a substantial rate of CRu designations by the predominantly CT-based IWC classification due to the findings of residual masses by CT. Thus, using the IWC+PET classification, the vast majority of these CRus would be re-classified as CR or, less frequently, as PR. The most critical finding of the study by Juweid et al, however, was that based on PFS as an outcome measure, the IWC+PET-based response classification was more accurate than IWC-based classification in predicting the ‘‘expected’’ outcome of patients in the various response categories. For example, whereas twothirds of the patients designated as PR by IWC+PET progressed within a median of only 8.5 months, almost 60% of patients with a PR designation by IWC were progression-free at a median of nearly 3 years post-therapy. Interestingly, in the 35 patients with CR by IWC+PET, PFS was identical to that in the only 17 patients with CR by IWC (31.5+ months), although slightly more than half of the IWC+PET CRs were actually thought to be in PR or even only having SD. This indicates that the ‘‘quality’’ of a CR by IWC+PET, in terms of conferring a much more favourable prognosis compared with PR or SD, was very similar if not equal to the quality of a CR by IWC. The potential implications for the management of patients with aggressive NHL are apparent: whereas patients with PR or SD by IWC+PET should be evaluated for persistent disease and, if confirmed, considered for additional therapy, patients with CR by IWC+PET may be safely observed. In contrast, at least the PR designation by IWC, assigned to slightly more than one-third of the patients in this retrospective study, provides incorrect characterization of the actual response to treatment in an unacceptably high fraction (,60%) of patients with such designation, casting serious doubts about its usefulness. Obviously, the study by Juweid et al [6] had its limitations: the retrospective nature of the evaluation performed; the relatively small number of patients evaluated; and the somewhat variable timing of PET following treatment and variable number and type of chemotherapy cycles administered. Yet this ‘‘proof-of-principle’’ study, along with similar evidence of superiority of PET compared with CT, were convincing enough that the IHP adopted IWC+PET-based definitions of CR, PR, SD and PD similar to those presented in the study by Juweid et al [6], The British Journal of Radiology, May 2006

Commentary: Response assessment of NHL by PET

with a proposal to completely eliminate the CRu designation. These new definitions were recently presented by Dr Bruce Cheson at the 2005 Annual meeting of the American Society of Hematology in Atlanta, Georgia [7]. In addition to the proposed elimination of the CRu designation, a noteworthy change compared with the criteria proposed by Juweid et al [6] is that the criteria will now take into consideration that some subtypes of NHL, such as extranodal marginal zone lymphoma (MZL) of mucosaassociated lymphoid tissue (MALT) are frequently not FDG-avid, hence assessment will still be based on conventional imaging. The criteria will now also be applicable to patients with typically FDG-avid indolent NHLs, such as follicular NHL, and also to patients with HL. The details of these revised criteria will be reported in an upcoming manuscript, to be submitted for publication shortly. A more imaging-focused report of the imaging committee of the IHP will follow subsequently. Despite the clearly favourable contribution of PET or PET/CT on response assessment of NHL, it should be noted that, like any other modality, PET or PET/CT are not perfect, as shown by several studies including the study of Juweid et al [6]. In the latter study, one-third of the patients (4/12) with a PR designation by IWC+PET based on positive PET findings suggesting residual disease remained progression-free for more than 3 years. False-positive PET findings at the site of residual masses may occur due to thymic rebound in the vicinity of a residual mass in the superior mediastinum, sometimes mistaken as uptake in the mass, or post-therapy inflammatory changes, the latter substantially more frequent following radiation than after chemotherapy or chemoimmunotherapy [5, 8]. False-positive findings outside the site of residual masses also may be caused by rebound thymic hyperplasia, or infectious or inflammatory processes including sarcoidosis. With PET/CT there is a significant improvement in the diagnostic accuracy of PET, principally because of the more accurate anatomic localization of the PET findings provided by the almost-simultaneously acquired CT. The generally precise registration of the PET and CT images largely eliminates the problem of false-positive physiologic FDG uptake, such as that in muscle or brown fat, which is highly variable between patients [8]. In fact, a recent study on the staging and re-staging of 73 lymphoma patients showed that PET/CT was superior to PET alone with reported accuracies of 93% and 83%, respectively [9]. To further improve the diagnostic accuracy of PET, it will also be critical to address the issue of standardization of PET technique (e.g. timing after last treatment, timing of scanning after the injection) in future studies. Finally, it is important to note that PET and PET/CT are likely to be useful not only for response assessment after completion of therapy, but also for initial staging and early response assessment after only a few cycles (e.g. 1–4) of a 6– 8-cycle chemotherapy or chemoimmunotherapy regimen. In fact, several studies have shown that an interim assessment with PET after 2–4 cycles of chemotherapy is at least as accurate an assessment of disease response in terms of prediction of progression free survival as an end of treatment assessment [8, 10–12].


In summary, PET or PET/CT is poised to alter the response assessment paradigm in patients with lymphoma which is currently largely based on conventional imaging and has, in fact, already resulted in a significant and important modification of the currently widely used and accepted IWC for response assessment of lymphoma. The full report of this revision is eagerly awaited.

References 1. Cheson BD, Horning SJ, Coiffier B, Shipp MA, Fisher RI, Connors JM, et al. Report of an international workshop to standardize response criteria for non-Hodgkin’s lymphomas. NCI Sponsored International Working Group. Joˆ Clin Oncol 1999;17:1244–53. 2. Reske SN. PET and restaging of malignant lymphoma including residual masses and relapse. Euroˆ Joˆ Nucl Med Mol Imaging 2003;Suppl. 1:S89–96. 3. Jerusalem G, Beguin Y, Fassotte MF, et al. Whole-body positron emission tomography using 18F-fluorodeoxyglucose for posttreatment evaluation in Hodgkin’s disease and non-Hodgkin’s lymphoma has higher diagnostic and prognostic value than classical computed tomography scan imaging. Blood 1999;94:429–43. 4. Spaepen K, Stroobants S, Dupont P, et al. Prognostic value of positron emission tomography (PET) with fluorine 18 Fluorodeoxyglucose ([18F]FDG) after first-line chemotherapy in non-Hodgkin’s lymphoma: is [18F]FDG-PET a valid alternative to conventional diagnostic methods? Joˆ Clin Oncol 2001;19:414–9. 5. Naumann R, Vaic A, Beuthien-Baumann B, et al. Prognostic value of positron emission tomography in the evaluation of post-treatment residual mass in patients with Hodgkin’s disease and non-Hodgkin’s lymphoma. Broˆ Joˆ Haematol 2001;115:793–800. 6. Juweid ME, Wiseman GA, Vose JM, Ritchie JM, Menda Y, Wooldridge JE, et al. Response assessment of aggressive non-Hodgkin’s lymphoma by integrated international workshop criteria and fluorine-18-fluorodeoxyglucose positron emission tomography. Joˆ Clin Oncol 2005;23:4652–61. 7. Cheson BD, Pfistner B, Juweid ME, Specht L, Rosen ST, Gascoyne R, et al. Revised response criteria for malignant lymphomas from the members of the International Harmonization Project (IHP) of the Competence Network Malignant Lymphoma. Blood 2005;106:108, (Abstract). 8. Juweid ME, Cheson BD. Role of PET in lymphoma. J Clin Oncol 2005;23:4577–80. 9. Allen-Auerbach M, Quon A, Weber WA, Obrzut S, Crawford T, Silverman DH, et al. Comparison between 2-deoxy-2-[18F]fluoro-D-glucose positron emission tomography and positron emission tomography/computed tomography hardware fusion for staging of patients with lymphoma. Mol Imaging Biol 2004;6:411–6. 10. Spaepen K, Stroobants S, Dupont P, et al. Early staging positron emission tomography (PET) with fluorine 18 fluorodeoxyglucose ([18F]FDG) predicts outcome in patients with aggressive non-Hodgkin’s lymphoma. Blood 2001;98:726a, (Abstract). 11. Jerusalem G, Beguin Y, Fassotte MF, et al. Persistent tumor 18F-FDG uptake after a few cycles of polychemotherapy is predictive of treatment failure in non-Hodgkin’s lymphoma. Haematologica 2000;85:613–8. 12. Fields PA, Mikhaeel G, Hutchings M, van der Walt J, Nunan T, Schely SA. The prognostic value of interim positron emission tomography scans combined with immunohistochemical data in diffuse large-B-cell lymphoma. Hematologica 2005;90:1711–3.

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Acute adverse reactions to magnetic resonance contrast media – gadolinium chelates 1

A LI, MBChB, FRCR, 2C S WONG, MBChB, 2M K WONG, 2 M C AU YEUNG, MBBS, FRCR, FHKCR

MBBS, FRCR, FHKCR,

2

C M LEE,

MBBS, FRCR, FHKCR

and

1

Department of Radiology and Organ Imaging, United Christian Hospital, 130 Hip Wo Street, Kwun Tong, Hong Kong, SAR and 2Department of Radiology, Princess Margaret Hospital, Hong Kong, SAR ABSTRACT. The objective of this study was to evaluate the clinical safety of intravenous gadolinium-based contrast media used in patients who underwent MRI at a single institution. Acute adverse reactions to intravenous gadolinium-based contrast media used for MRI at the Princess Margaret Hospital, Hong Kong, SAR, from January 1999 to November 2004 were recorded in an incidence log book. The medical records of patients’ demographics were retrospectively reviewed and the nature, frequency and severity of the adverse reactions were investigated and documented. The incidence of acute adverse reactions to intravenous gadolinium-based contrast media was 0.48% (45 patients with 46 adverse reactions). The severity of these adverse reactions were 96% mild, 2% moderate (one patient developed shortness of breath that required oxygen supplementation and intravenous steroidal management) and 2% severe (one patient developed an anaphylactoid reaction, but successfully recovered through timely resuscitation). No patients were recorded as having contrast extravasation and none died as a result of any adverse reaction. Among the 45 patients who developed adverse reactions, three patients (6.7%) had prior adverse reactions to iodinated contrast media, three (6.7%) had prior reactions to a different gadolinium-based contrast agent, one (2%) had asthma and nine (20%) had a history of drug/food allergy. Overall, 41% of the adverse reactions were not documented in the final MRI report or the clinical medical records. Gadolinium-based contrast media are safe and well tolerated by the vast majority of patients. In our study, the adverse reaction rate (0.48%) and the incidence of severe anaphylactoid reaction (0.01%) concur with those reported in the literature. Although most of the symptoms are mild and transient, these adverse reactions must be accurately documented and managed.

MRI is often considered a safer alternative to CT for patients who cannot tolerate iodinated contrast media. The intravenous gadolinium-based contrast media is widely used in MRI for a number of reasons: it improves sensitivity of lesion detection, and provides better diagnostic specificity and more accurate depiction of the extent of the disease involvement. Although the total adverse reaction rates are in the range of 0.17% to 2.40%, many of the problems that lead to concerns about the use of iodinated contrast agents also exist in the use of gadolinium-based contrast media [1–3]. The purpose of this report is to quantify the frequency and manifestations of adverse reactions associated with the use of gadoliniumbased contrast media in our regional hospital.

Materials and methods From January 1999 to November 2004, 18 142 MRI examinations were performed at the Princess Margaret Hospital, Hong Kong, SAR (including those performed on patients referred from Yan Chi Hospital and Caritas Medical Center, Hong Kong, SAR). Gadolinium-based contrast media was used in 9528 (52.8%) examinations. All adverse reactions that occurred in patients within 1 h 368

Received 9 May 2005 Revised 24 August 2005 Accepted 17 October 2005 DOI: 10.1259/bjr/88469693 ’ 2006 The British Institute of Radiology

of intravenous administration of the gadolinium-based contrast media were documented in an incidence log book by the nursing staff or the responsible radiologist. This information was either provided voluntarily by the patients or observed by the MRI radiographers, nurses or radiologists. The following details were recorded in the incidence log book: demographic data of the patients (name, medical record number, sex and age); medical history (e.g. asthma, or previous allergic reactions such as allergies to iodinated or gadolinium-based contrast media); severity and nature of adverse reactions; treatments given and actions taken; the name and signature of the reporter and responsible radiologists. This information, the medical records and the MRI radiology reports relevant to the patients were reviewed. An adverse reaction was defined to be any unfavourable or unintended alteration in the clinical status that was temporally associated with the use of contrast media, even if it was not considered to be related to administration of the drug. The sensations of hot or cold, transient sensations of taste or mild pain at the injection site were excluded as adverse reactions. The severity of the reaction was subdivided into mild, moderate and severe. Mild reactions were defined as symptoms that were transient, or that required minimal or no therapy. The British Journal of Radiology, May 2006

Adverse reactions to MR contrast media

Moderate reactions were more severe in nature or occurred for a longer duration than mild reactions; therapy was required but the patient was not considered critical. Patients who had severe reactions required resuscitation and their condition was considered to be possibly life-threatening.

Results A total of 45 (0.48%) of the 9528 patients were recorded as having adverse reactions to the gadolinium-based media. One patient developed repeated mild contrast reactions in two separate MRI studies (counted as one patient in our study). The mean age of the study population was 42 years (median 42 years, range 12–79 years). Five patients (11%) were in the paediatric age group (16 years or younger). 58% of the patients were female (n526). Table 1 displays the adverse reactions that occurred in patients with a history of potentially predisposing conditions. Among the 45 patients who developed adverse reactions to gadolinium-based contrast media, 8 patients (18%) had previous exposure to an iodinated contrast media (one patient had a prior adverse reaction to the iodinated contrast media); 34 patients (75%) had no previous exposure; status was unknown in 3 patients (7%). Nine of the 45 patients (20%) who developed adverse reactions had prior exposure to gadoliniumbased contrast media (three patients had prior adverse reactions to the gadolinium-based contrast media); 33 patients (73%) had no prior exposure; the contrast history was unknown in three patients (7%). Nine of the 45 patients (20%) had a history of allergies (eczema, to drugs, alcohol, etc.) and one patient (2%) had a history of asthma. Table 2 displays the type, number, and severity of adverse reactions that occurred after the use of intravenous gadolinium-based contrast media. Gadoterate meglumine (Dotarem; Guerbet, France) was used in three patients, Gadodiamide (Omniscan; Nycomed, New York, NY) in 23 patients and Gadopentetate dimeglumine (Magnevist; Berlex, Schering AG, Germany) in 19 patients. The dose was administered according to the manufacturers’ instructions. Two of the patients who developed adverse reactions underwent MR angiography, which involved the use of power injectors to administer the contrast agent at a rate of 2 ml per second. In all other cases, the contrast media was manually administered by injection. Table 1. The medical history of patients who had adverse

Table 2. The type and severity of the adverse reactions Adverse reactions

Severity (number of patients)

Urticaria Rash Nausea/vomiting Vasovagal attack Dizziness Confusion Dyspnoea Chest discomfort Palpitation Anaphylactoid shock Total

Mild (4) Mild (11) Mild (18) Mild (2) Mild (3) Mild (1) Mild (1), Moderate (1) Mild (1) Mild (2) Severe (1) 45

None of the patients in our study with a history of adverse reactions to gadolinium-based (three patients) or iodinated contrast media (one patient) had pre-medication to prevent recurrence of such adverse reactions. Among these 45 patients, the majority of adverse reactions to gadolinium-based contrast media were mild and transient (96%) (Table 2). The most frequent mild reaction was nausea and vomiting (40%) and 33% of patients were affected by urticaria and a rash. All patients with nausea, vomiting and urticaria responded to conservative management, and the seven patients who developed a rash were treated with a stat dose of oral diphenhydramine. One patient developed dyspnoea that was treated with oxygen supplementation, intravenous steroids and fluids, and further inpatient management. The patient was subsequently discharged uneventfully. This patient did not have a history of respiratory disorders such as asthma or chronic obstructive airway disease—bronchospasm was the presumed cause of the dyspnoea. Among the 9528 patients who underwent contrast MRI, one patient (0.01%) developed an anaphylactoid reaction. This 77-year-old patient was referred for outpatient MRI for follow up of a known malignant melanoma of the nostril that had intracranial extension. The patient had undergone a contrast enhanced MRI study using gadodiamide without complications, but developed anaphylactoid shock shortly after injection of gadopentate dimeglumine. The patient did not have a history of asthma, allergies, or previous sensitivities to drugs including iodinated contrast media. The patient was successfully resuscitated (by securing the airways, use of oxygen, and intravenous diphenhydramine, steroids, adrenaline and fluids) and was hospitalized for intensive care and later discharged uneventfully. No contrast extravasation was recorded for any patient and none died as a result of an adverse reaction.

reactions Parameters

Asthma History of allergy MRI contrast agent Previous exposure Adverse reaction Iodinated contrast agent Previous exposure Adverse reaction

Discussion

Patient history Yes

No

Unknown

1 (2%) 9 (20%)

37 (82%) 35 (78%)

7 (16%) 1 (2%)

9 (20%) 3

33 (73%)

3 (7%)

8 (18%) 1

34 (75%)

3 (7%)


Gadolinium-based contrast agents constitute the largest group of MRI contrast media and are considered to be very safe and well tolerated. Gadopentetate dimeglumine, the first agent approved by the Food and Drug Administration (FDA) for clinical use in the USA (1988), has a total adverse reaction rate in the range of 0.17– 2.40% [1–3]. Nelson et al [1] reported the adverse reaction rate in a patient population of 15 496 given gadopentetate dimeglumine was 2.4%. The overall adverse reaction 369

A Li, C S Wong, M K Wong et al

rate was 1.2% if patients voluntarily reported their symptoms instead of being directly questioned. The safety of the gadolinium chelates is largely based on their stability in vivo. The chelates are designed to bind gadolinium ions tightly, preventing the possible release of free gadolinium ions, which are extremely toxic to the body. The differences between the gadolinium chelates in terms of chemical stability and transmetallation gives rise to some theoretical concerns. Transmetallation refers to the substitution of the chelates by copper and zinc ions in the body leading to the release of free gadolinium ions. Both gadopentetate dimeglumine and gadodiamide can inhibit the zinc-dependent angiotensin-converting enzyme (ACE) in the body and may also give rise to transmetallation [4–7]. To our knowledge, no harmful effects in humans caused by free gadolinium ion deposition resulting from the clinical use of these agents have been reported. All commercially available gadolinium chelates have comparable adverse reaction profiles including mild reactions such as nausea and severe anaphylactoid reactions [4]. For reference, the incidence of urticaria in clinical trials for gadopentetate dimeglumine is 0.3% in 1068 patient [8], for gadodiamide it is 0.7% in 439 patients [9], and for gadoterate meglumine it is 0.4% in 518 patients [10]. These figures are not statistically differentiable. Numerous cases of severe anaphylactoid reactions associated with the use of gadolinium-based contrast media have been reported [11–13]. For example, a fatal reaction to gadopentetate dimeglumine was reported in 1995. This patient did not have a history of asthma or previous allergic reactions to iodinated contrast media [14]. The incidence rate of severe anaphylactoid reaction in our series was 0.01%. This concurs with a large series by Murphy et al [3], in which two severe reactions to gadopentetate dimeglumine out of 21 000 patients were reported. For reference, in another study by Caro et al, the risk of a life-threatening event with iodinated radiographic contrast media was 0.031% for low-osmolarity contrast media, and 0.157% for conventional ionic contrast media [15]. Our results again confirm that gadolinium-based contrast media is safer than iodinated contrast media. In 1994, Witte et al reported a case of lifethreatening anaphylactoid reaction that occurred after the administration of intravenous gadoteridol (Prohance; Bracco Diagnostics, Princeton, NJ) in a patient who had previously received gadopentetate dimeglumine [16]. To our knowledge, our study is the first reported case in the English language literature of a patient who developed severe anaphylactoid reaction after intravenous administration of gadodiamide without having exhibited any adverse reaction with the prior use of gadopentetate dimeglumine. Although the incidence of life-threatening reactions to gadolinium-based contrast media is low, anaphylactoid reactions are among the worst adverse reactions that might occur with the use of MRI contrast agents. Therefore, it is prudent to provide sufficient training for personnel and to have available adequate resuscitation facilities in the MRI suite if these agents are to be used. Nelson et al reported that the likelihood of adverse reactions to gadolinium-based contrast media is about eight times higher in patients who have had previous reactions to gadolinium-based contrast media. The 370

severity of a second adverse reaction tends to be more pronounced than the first [1]. In our series, all three patients who had prior adverse reactions to gadoliniumbased contrast media developed adverse reactions after repeated use of such an agent. These three patients were not pre-medicated to reduce risk of contrast reactions. In an attempt to reduce the risk of a reaction to the contrast media, a different gadolinium-based media (gadodiamide) was administered instead of the previously used gadopentate dimeglumine. However, adverse reactions still developed despite the use of a different gadoliniumbased contrast media, but these reactions were mild and transient (rash and nausea). The frequency of adverse reactions among patients who have a history of adverse reactions to gadolinium-based contrast media cannot be determined from our data. However, our observations may support the findings of the study by Nelson et al. The role of pre-medication with corticosteroids as a preventive measure in patients with prior reactions to gadolinium-based contrast media has not been well established. In a study by Murphy et al [3], two patients were pre-medicated before the gadolinium contrast agent was administered. One patient had a single oral dose of 50 mg diphenhydramine 1 h before the procedure, and the second patient had an oral dose of 50 mg diphenhydramine 1 h before the procedure and 20 mg prednisolone every 6 h (starting 13 h before the examination). Both patients had the same adverse reaction as with the previous gadolinium exposure. The second patient returned for a further MRI study and was given the same pre-medication, but a different contrast agent from the previous one. He did not develop any complications. Murphy et al [3] recommends the use of 12–24 h pre-medication with corticosteroids and subsequent use of a different or lower osmolar MRI contrast agent as a preventative measure among this group of patients. The evidence of benefits of such pre-medication has not been well established. Until more data are available, gadolinium-based MRI contrast media should be used with caution in patients with a history of adverse reactions, in particular severe reactions. The frequency of adverse reactions to gadoliniumbased contrast agents is about 2.3–3.7 times higher in patients with a history of reactions to iodinated contrast material [1, 2]. In our study, three of the 45 patients who experienced adverse reactions to gadolinium-based contrast media had also had previous reactions to iodinated contrast media, but the rash and nausea symptoms were mild and transient. Again, administration of gadoliniumbased contrast media among these types of patients should be made with caution, because the role of steroid premedication has not been well established. Patients with asthma and various allergies are also at greater risk of adverse reactions. Nelson et al [1] reported that the frequency of adverse reactions is 1.5 and 1.9 times higher, respectively, compared with patients who have none of these allergies. Among the 45 patients in our study who had adverse reactions (Table 1), one patient had asthma and nine patients had a prior allergy (2 urticaria, 2 seafood, 1 eczema, 3 alcohol and 1 aspirin related). In this study, 19 out of 46 cases of contrast adverse reactions (41%) were not documented in the final MRI report or the clinical medical records. The adverse reactions in these cases were all mild and transient. The British Journal of Radiology, May 2006

Adverse reactions to MR contrast media

Proper documentation of the adverse reaction is not only essential for medico-legal purposes, but also more importantly for the interest and well being of the patient. If precautions could be taken before the repeated administration of gadolinium-based contrast media, then the awareness and management of adverse reactions would be more efficient. Therefore, the accurate recording of the reactions associated with gadolinium-based contrast media should be as thorough as they are for reactions associated with other drugs. Several limitations were identified in the design of this study. The design was a retrospective review of the incidence log book, which was originally intended for documenting all acute adverse reactions to the gadolinium chelates for medical records and for departmental audit. Although clear instructions were given to staff in the department, all acute adverse reactions may not have been documented, in particular those that were trivial in nature. Voluntary reporting by patients was also likely to lead to under-reporting. Additional minor reactions might have been reported if patients were questioned directly or by way of a questionnaire. Only acute adverse reactions were recorded, occurring within 1 h between the gadolinium-based contrast media injection and the patient leaving the department after observation. In the study by Nelson et al, 44.9% of adverse reactions occurred more than 1 h after the gadolinium-based contrast media injection. Those reactions were all mild and transient, with no severe or life-threatening events recorded [1]. In such circumstances, the incidence of acute adverse reactions in our patient population still concurs with those reported in the literature. In conclusion, gadolinium-based contrast media is safe and well tolerated by the vast majority of patients. Although most adverse reactions are mild and transient, life-threatening anaphylactoid reactions do occur. Personnel should be well trained and resuscitation facilities should be readily available in the MRI examination room. Accurate documentation and cautious management of all adverse reactions related to gadolinium-based contrast media should be made in the same manner as for other drug reactions.

References 1. Nelson KL, Gifford LM, Lauber-Huber C, Gross CA, Lasser TA. Clinical safety of gadopentetate dimeglumine. Radiology 1995;2:439–43.


2. Niendorf H, Dinger J, Haustein J, Cornelius I, Alhassan A, Clauss W. Tolerance data of Gd-DTPA: a review. Eur J Radiol 1991;13:15–20. 3. Murphy KJ, Brunberg JA, Cohan RH. Adverse reactions to gadolinium contrast media: a review of 36 cases. AJR Am J Roentgenol 1996;167:847–9. 4. Runge VM. Safety of approved MR contrast media for intravenous injection. J Magn Reson Imaging 2000;12:205–13. 5. Wedeking P, Kumar K, Tweedle MF. Dissociation of gadolinium chelates in mice: relationship to chemical characteristics. Magn Reson Imaging 1992;10:641–8. 6. Tweedle MF, Wedeking P, Kumar K. Biodistribution of radiolabeled, formulated gadopentetate, gadoteriodol, gadoterate, and gadodiamide in mice and rats. Invest Radiol 1995;30:372–80. 7. Corot C, Idee JM, Hentsch AM, et al. Structure-activity relationship of macrocyclic and linear gadolinium chelates; investigation of transmetallation effect on the zinc-dependent metallopeptidase angiotensin-converting enzyme. J Magn Reson Imaging 1998;8:695–702. 8. Goldstein HA, Kashanian FK, Blumetti RF, et al. Safety assessment of gadopentetate dimeglumine in U.S. clinical trials. Radiology 1990;174:17–23. 9. Sze G, Brant-Zawadzki M, Haughton VM, Maravilla KR, et al. Multicenter study of gadodiamide injection as a contrast agent in MR imaging of the brain and spine. Radiology 1991;181:693–9. 10. Oudkerk M, Sijens PE, Van Beek EJ, Kuijpers TJ. Safety and efficacy of dotarem (Gd-DOTA) versus Magnevist (GdDTPA) in magnetic resonance imaging of the central nervous system. Invest Radiol 1995;30:75–8. 11. Tishler S, Hoffman JC Jr. Anaphylactoid reactions to i.v. gadopentetate dimeglumine. Am J Neuroradiol 1990;11:1167–9. 12. Tardy B, Guy C, Barral G, Page Y, Ollagnier M, Bertrand JC. Anaphylactic shock induced by intravenous gadopentetate dimeglumine. Lancet 1992;339:494. 13. Sahni H, Jayakumar PN, Lalla RK, Vasudev MK, Kovoor JME. Anaphylactoid reaction following intravenous gadolinium dimeglumine - a case report. Ind J Radiol Imag 2002;12:1:81–2. 14. Jordan RM, Mintz RD. Fatal reaction to gadopentetate dimeglumine. AJR Am J Roentgenol 1995;164:743–4. 15. Caro JJ, Trindade E, McGregor M. The cost-effectiveness of replacing high-osmolarity with low-osmolarity contrast media. AJR Am J Roentgenol 1992;159:869–74. 16. Witte RJ, Anzai LL. Life-threatening anaphylactoid reaction after intravenous gadoteridol administration in a patient who had previously received gadopentetate dimeglumine. Am J Neuroradiol 1994;15:523–4.

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Observer variation in the evaluation of lumbar herniated discs and root compression: spiral CT compared with MRI 1

J C VAN RIJN, MD, 2N KLEMETSO, MD, 1J B REITSMA, MD, PhD, 1P M BOSSUYT, PhD, 2F J HULSMANS, MD, 4 2 3 2 PhD, W C PEUL, MD, PhD, G J DEN HEETEN, MD, PhD, J STAM, MD, PhD and C B L M MAJOIE, MD, PhD Departments of 1Clinical Epidemiology and Biostatistics, 2Radiology and 3Neurology, Academic Medical Center, Meibergdreef 9, 1105 AZ, Amsterdam and the 4Department of Neurosurgery, Leiden University Medical Center, Albinusdreef 2, 2333 ZA, Leiden, The Netherlands

ABSTRACT. Spiral CT is considered the best alternative for MRI in the evaluation of herniated discs. The purpose of this study was to compare radiological evaluation of spiral CT with MRI in patients suspected of herniated discs. 57 patients with lumbosacral radicular syndrome underwent spiral CT and 1.5 T MRI. Two neuroradiologists independently evaluated 171 intervertebral discs for herniation or ‘‘bulge’’ and 456 nerve roots for root compression, once after CT and once after MRI. We compared interobserver agreement using the kappa statistic and we performed a paired comparison between CT and MRI. For detection of herniated or bulging discs, we observed no significant difference in interobserver agreement (CT kappa 0.66 vs MRI kappa 0.71; p50.40). For root compression, we observed significantly better interobserver agreement at MRI evaluation (CT kappa 0.59 vs MRI kappa 0.78; p50.01). In 30 of 171 lumbar discs (18%) and in 54 of 456 nerve roots (12%), the observers disagreed on whether CT results were similar to MRI. In the cases without disagreement, CT differed from MRI in 6 discs (3.5%) and in 3 nerve roots (0.7%). For radiological evaluation of lumbar herniated discs, we found no evidence that spiral CT is inferior to MRI. For evaluating lumbar nerve root compression, spiral CT is less reliable than MRI.

In a series of comparative radiological studies on the evaluation of lumbar herniated discs, MRI was found to be ‘‘slightly better’’ than [1–4] or equal to [5–8] CT. We found one study suggesting CT to be more accurate than MRI [9], but overall, MRI is preferred and CT seems to have comparable capacity to diagnose herniated discs. Therefore, current guidelines designate MRI as the firstchoice investigation and suggest CT as the alternative in the evaluation of the lumbar back if MRI is contraindicated or unavailable [11–13]. Besides herniated discs, the direct evaluation of nerve roots by MRI has been considered an important asset to facilitate decision making in patients with back pain [14– 17]. No evidence, however, is available on the capabilities of spiral CT to evaluate lumbar nerve root compression. The purpose of our study was to compare spiral CT with MRI in a series of patients suspected of lumbar herniated discs. For each technique we assessed interobserver agreement in detecting herniated discs, bulging discs and nerve root compression. We also performed a paired, direct comparison between spiral CT and MRI.

Methods This study was conducted between June 1999 and June 2000, as part of a larger project on the diagnostic process 372

Received 16 August 2005 Revised 23 November 2005 Accepted 9 December 2005 DOI: 10.1259/bjr/26216335 ’ 2006 The British Institute of Radiology

of patients with lumbosacral radicular pain at the University of Amsterdam, The Netherlands. The Institutional Review Board approved the study protocol. Patients were recruited from the neurology outpatient department. Those eligible were patients referred by their general practitioner with lumbosacral radicular syndrome (LRS) with suspected disc herniation at the levels L3–L4 to L5–S1 in whom conservative treatment had been unsuccessful. LRS was defined according to the national general practitioners’ guideline and the consensus statement on diagnosis and treatment of LRS defined by the Dutch Neurology Society [18]. The hallmark of this definition is continuous monoradicular or multiradicular pain below the knee with a primary suspicion of disc herniation. Excluded were patients younger than 18 years or older than 70 years, pregnant women, patients with a previous history of lumbosacral herniation or lumbosacral surgery, as well as patients with contraindication for MRI. Eligible patients received written and oral information about the study. After the neurologist had confirmed the diagnosis of LRS, consenting patients were subjected to spiral CT and MRI within 1 week. No specific treatment was given within this period. The British Journal of Radiology, May 2006

Spiral CT compared with MRI in the evaluation of suspected lumbar herniated discs

Imaging techniques Spiral CT Lumbar CT examinations were performed on a 2-slice CT-Twin scanner (Philips Medical Systems, Best, The Netherlands). Helical CT-scans were made with 120 kV, 265 mAs (effective dose: 6.4 mSv), table feed of 1.4 mm s-1, 2 mm61 mm collimation (effective slice width, 1.1 mm) and 0.5 mm increment from the level L3 to the bottom of S1. The gantry angle was aligned through the disc space of L4–L5. In addition to the original axial images with the bone window (level 400 HU/window 1600 HU), reformatted 4 mm axial sections of all scanned interspaces were made parallel to the interspaces using the soft tissue window (level 50 HU/window 180 HU). Also reformatted ‘‘0 mm’’ (i.e. slice width in the order of 1 mm) sagittal sections were made using the bone window, and 4 mm sagittal sections using the soft tissue window. The quality of the CT-images used in this study, obtained with a 2-slice CT-scanner, is comparable with the quality of more modern, multislice CT-scanners. The effective slice width in this study was 1.1 mm, which is only slightly more than the effective slice width of multislice CT-scanners. The effective slice width is usually 25–30% higher than the collimation width and thus in the order of 0.9–1.0 mm for, e.g. a collimation of 16 mm60.75 mm [19, 20].

MRI Lumbar MR examinations were performed with a 1.5 T Signa LX Scanner (General Electric Medical Systems, Milwaukee, WI) using a dedicated lumbar spine surface coil. The protocol included sagittal spin-echo T1 weighted (repetition time (TR) 500 ms, echo time (TE) 14 ms) and proton density/T2 weighted (TR 3500 ms, TE 20–120 ms) fast spin echo images with 4 mm slice thickness, 0.5 mm intersection gap, 2006512 matrix and 29 cm629 cm field of view. In addition, axial spin-echo T1 weighted (TR 520 ms, TE 12 ms) and fast spin-echo T2 weighted (TR 4500 ms, TE 120 ms) images were obtained from the increment of L3 to the bottom of S1 with 4 mm slice thickness, 0.5 mm intersection gap, 2006256 matrix and 15 cm615 cm field of view. Axial images were obtained without angulation.

Image evaluation Two experienced neuroradiologists (CB and FJ) independently evaluated all CT and MR images with knowledge of the side and level of symptoms. Per observer, the images were evaluated in two independent sessions. There was at least 6 weeks between these sessions. The images of either CT or MRI were presented per patient. It was randomly allocated whether CT or MRI was evaluated in the first session. Three lumbar discs were examined per patient at levels L3–L4 to L5–S1. Each disc was scored for the The British Journal of Radiology, May 2006

presence of a herniation. No distinction between protrusion and extrusion was made; both were considered ‘‘herniated disc’’. If no herniation was detected, readers evaluated the presence of a bulging disc. The definition of a bulging disc was according to the description by Jensen [21]: ‘‘circumferential symmetric extension of the disc beyond the interspace’’. Nerve roots L3 through S1 were evaluated per side (eight nerve roots per patient) within the scanned area. A nerve root was defined as the part of the nerve between the central canal and the ganglion. As the scan protocol for both techniques was from the increment of vertebra L3 through the bottom of S1, the radiologists were not able to evaluate a minimal proximal part of root L3 and a minimal distal part of root S1. A five-point scale was used per nerve root, anchored as definitely no root compression, possibly no root compression, indeterminate, possibly root compression, and definitely compression. For the analyses, these responses were dichotomized as ‘‘root compression’’ (possibly or definitely) or ‘‘no root compression’’ (all other categories). MRI examples of the different categories have been presented in an earlier report on the same series of patients [22].

Data analysis Per technique, we calculated interobserver agreement for the evaluation of herniated discs and bulging discs by constructing 363 tables. Data were matched for lumbar level (L3–L4 to L5–S1). Presence of nerve root compression was analysed using a 262 table. These data were matched for each side (left and right) and for each nerve root (L3 to S1). To acknowledge possible correlation between intervertebral discs and nerve roots within one patient, bootstrap sampling was used to calculate the unweighted interobserver kappa statistics and standard-errors for each contingency table [23]. A two-sided z-test was then used to compare the kappa statistics of CT with MRI. To differentiate between CT versus MRI differences due to disagreement and ‘‘true’’ differences (no disagreement amongst observers), we performed a paired comparison of the results of CT and MRI. We limited the analysis to the detection of herniated discs. Therefore, the readings for each technique were recoded to ‘‘herniated disc’’ or ‘‘no herniated disc’’. The latter category comprised bulging discs and discs reported as having no abnormality. The results were coded per technique as follows: (1) both observers agreed on the presence of a Hernia Nuclei Pulposi (HNP), (2) observers disagreed or (3) both observers agreed on the absence of a HNP. Subsequently, a 363 table including the results of both observers was constructed to compare both techniques on a per patient basis. The McNemar test for paired data was used to detect differences between modalities. All calculations were performed with SAS 9.0 and SPSS 11.0. Values lower than p.0.05 were considered to indicate statistical significance. 373

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Results

root, both readers reported root compression at MRI but not at CT evaluation (case is presented in Figure 3).

64 consecutive eligible patients with LRS were identified. Three patients did not undergo MRI because of claustrophobia. Four more patients had incomplete data of spiral CT and MRI, leaving 57 patients with 171 lumbar discs and 456 lumbar roots for analysis. Table 1 shows the interobserver agreement. There was no significant difference between CT and MRI with respect to interobserver agreement of herniated disc and bulging disc evaluation, but for root compression the kappa statistic was significantly higher for MRI. Figure 1 shows an example of a case in which CT and MRI agree on a HNP with root compression without any interobserver disagreement. Table 2 shows the results of the paired comparison of CT versus MRI in detecting herniated discs and root compression. Of 171 discs, observers disagreed on herniation at CT evaluation in 20 discs (12%) compared with 12 discs (7%) at MRI. The difference was not significant (McNemar, p50.22). In 141 (83%) cases there was no disagreement at CT evaluation as well as no disagreement at MRI evaluation. In three of these cases both readers reported a HNP at CT, but not at MRI evaluation, and three times both readers reported a HNP at MRI but not at CT evaluation. Observers disagreed on the presence of root compression when evaluating CT in 40 of 456 nerve roots (8.8%) compared with 22 roots (4.8%) when evaluating MRI. The difference was significant (McNemar p50.03). An example of a case in which observers disagreed on root compression at CT but not on MRI is presented in Figure 2. In 402 roots (88%) there was no disagreement at CT evaluation as well as no disagreement at MRI evaluation. Of these cases, twice both readers reported root compression at CT but not at MRI evaluation. In one

Discussion In this study we found no evidence that spiral CT is inferior to MRI in the evaluation of herniated lumbar discs and bulging discs, but we observed better interobserver agreement at MRI with respect to the evaluation of nerve root compression. The paired comparison revealed that in the vast majority of the observed differences between spiral CT and MRI there was interobserver disagreement involved (either at CT or at MRI or at both evaluations). In these cases, we cannot discriminate between differences due to reader variability or due to ‘‘real’’ difference between modalities. The observed differences between CT and MRI without any disagreement indicate possible ‘‘real’’ differences. However, after re-evaluation of these cases we can conclude that all HNPs and all cases of root compression were detectable on the CT images as well as on the MR images. It remains impossible to determine which one of these imaging techniques corresponded best with ‘‘the truth’’ because there is no proper reference standard available for this study. However, we observed no particular pattern in the distribution of discordant readings: the differences between spiral CT and MRI were divided equally amongst both techniques. This supports the assumption that spiral CT and MRI are comparable in herniated disc evaluation. Observer disagreement is a general obstacle in the radiological evaluation of lumbar intervertebral discs, both for spiral CT as well as for MRI. In an earlier published study in the same series of patients, we

Table 1. Spiral CT and MRI interobserver kappa-statistics with 95% confidence intervals Lumbar disc evaluation Spiral CT

MRI Observer 2

Observer 1

HNP

Observer 2 Bulge

HNP 49 10 Bulge 3 24 No HNP 1 9 Total 53 43 Kappa 0.66 (0.56–0.75) Difference50.05; p50.40

No HNP

Total

Observer 1

HNP

6 9 60 75

65 36 70 171

HNP 52 Bulge 7 No HNP 1 Total 60 Kappa 0.71 (0.63–0.80)

Bulge

No HNP

Total

3 20 15 38

1 4 68 73

56 31 84 171

Root compression Spiral CT

MRI Observer 2

Observer 1

RC

Observer 2 No RC

RC 35 16 No RC 24 381 Total 59 397 Kappa 0.59 (0.47–0.70) Difference50.19; p50.01

Total

Observer 1

RC

51 405 456

RC 45 No RC 12 Total 57 Kappa 0.78 (0.69–0.87)

No RC

Total

10 389 399

55 401 456

HNP, Hernia Nuclei Pulposi; RC, root compression.

374



Figure 1. 22-year-old male with left-sided lumbosacral radicular syndrome of L5 and paresis of the anterior tibial muscle and with hypaesthesia of L5 and S1. Images show a left-sided herniated disc at level L4–L5 with root compression of L5 reported at CT as well as at MRI by both observers (arrows). (a) Axial CT L4–L5. (b) Axial T2 weighted MRI L4–L5.

investigated the possible causes of interobserver disagreement in MRI evaluation [22]. We identified a lack of consensus on nomenclature of bulging discs as the main cause of disagreement. Other causes were additional pathology (e.g. spondylolisthesis or collapsed vertebral bodies) and small sized lesions, which were believed to be missed more easily. The disagreement on nerve root evaluation can for a large part be attributed to the lack of a proper standardized and reproducible method to categorize root compression. We chose to use a five point scale to evaluate root compression. Compared with a dichotomous test result, the advantage of this method was that it

resembled clinical practice because the observers were given the opportunity to express their uncertainty. A limitation of this method was that we were not able to retrospectively retrieve the decisive items to get to a radiological diagnosis. Future studies should focus on reducing observer disagreement through the development of standardized radiological nomenclature for lumbar back abnormalities. Development of universal selection criteria to identify specific target groups of patients with back pain and more specific imaging strategies are needed to improve efficiency in decision making. This is also supported by the recent MR-based study by

Table 2. Paired comparison of spiral CT versus MRI Herniated disc detection MRI CT

HNP

Disagree

No HNP

Total

HNP Disagree No HNP Total

43 6 3 52

4 2 6 12

3 12 92 107

50 20 101 171

Root compression MRI CT

RC

Disagree

No RC

Total

RC Disagree No RC Total

32 12 1 45

1 8 13 22*

2 20 367 389

35 40* 381 456

HNP, both observers agreed on the presence of a HNP; No HNP, both observers agreed on the absence of a HNP. *Significantly more disagreement at CT evaluation (McNemar p50.03). HNP, Hernia Nuclei Pulposi; RC, root compression.


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Figure 2. 43-year-old male with left-sided lumbosacral radicular syndrome of S1. At both CT and MRI there is disagreement on root compression of L5 on the opposite side of signs and symptoms. (a) Axial CT L4–L5. Observers disagree on whether root L5 on the right (asymptomatic side), is compressed (arrow). (b) Axial T2 weighted MRI L4–L5. Observers agree on compression of L5 on the right (asymptomatic side) (arrow).

Cihangiroglu in 2004 [24]. To investigate the additional value of myelography (either conventional or by CT or MR imaging) might also be a valuable goal for future studies. Especially in the evaluation of nerve root impingement, it is suggested that myelography can play an important role [25].

In our study, the observers were not blinded for side and level of symptoms. This was done to mimic practice. As part of a larger project, we also obtained blinded evaluations of the same images. We observed no significant differences between blinded and not blinded evaluations [22]. After re-evaluation of the discordant

Figure 3. 58-year-old male with left-sided lumbosacral radicular syndrome of L3. (a) Axial CT L3–L4. Both observers reported ‘‘definitely no compression’’ of L3 on the left (arrow). (b) Axial T2 weighted MRI L3–L4. Both observers reported ‘‘definitely compression’’ of L3 on the left (arrow). 376



results in the present study, we have no reason to believe that the evaluations were influenced by knowledge of suspected side and level of the LRS. Our study was limited to the evaluation of herniated discs, bulging discs and root compression. MRI has better qualities to depict all surrounding soft tissue in the lumbar area. This property of MRI is one of the major reasons for neurosurgeons to use MRI to determine whether a patient is a candidate for surgical intervention. We did not investigate to what extent spiral CT might be useful within surgical triage. In summary, we recommend MRI as the investigation of choice in patients suspected of herniated discs. Nevertheless, a significant number of patients are not able to undergo MRI due to claustrophobia or other contraindications [26, 27]. In these patients, spiral CT seems to be an excellent alternative. However, for the evaluation of root compression, spiral CT appears to be not as reliable as MRI.

References 1. Goscinski I, Ulatowski S, Urbanik A. Comparison of the clinical usefulness of magnetic resonance (MR), computer tomography (CT) and radiculography (R) in diagnostic lumbar discopathy. Przegl Lek 2001;58:885–8. 2. Janssen ME, Bertrand SL, Joe C, et al. Lumbar herniated disk disease: comparison of MRI, myelography, and postmyelographic CT scan with surgical findings. Orthopedics 1994;17:121–7. 3. Forristall RM, Marsh HO, Pay NT. Magnetic resonance imaging and contrast CT of the lumbar spine. Comparison of diagnostic methods and correlation with surgical findings. Spine 1988;13:1049–54. 4. Pevsner PH, Ondra S, Radcliff W, et al. Magnetic resonance imaging of the lumbar spine. A comparison with computed tomography and myelography. Acta Radiol Suppl 1986;369:706–7. 5. Albeck MJ, Hilden J, Kjear L, et al. A controlled comparison of myelography, computed tomography and magnetic resonance imaging in clinically suspected lumbar disc herniation. Spine 1995;20:443–8. 6. Tullberg T, Grane P, Rydberg J, et al. Comparison of contrast-enhanced computed tomography and gadoliniumenhanced magnetic resonance imaging one year after lumbar discectomy. Spine 1994;19:183–8. 7. Thornbury JR, Fryback DG, Turski, PA, et al. Disk-caused nerve compression in patients with acute low-back pain: diagnosis with MR, CT myelography and plain CT. Radiology 1993;186:731–8. 8. Jackson RP, Cain JE, Jacobs RR, et al. The neuroradiographic diagnosis of lumbar herniated nucleus pulposus: II. A comparison of computed tomogrpahy (CT), myelography, CT-myelography, and magnetic resonance imaging. Spine 1989;14:1362–7.


9. Dellerud R, Johansen JG, Johnsen UL, et al. Differentiation between contained and noncontained lumbar disk hernias by CT and MR imaging. Acta Radiol 1995;36:491–6. 10. Eysel P, Rompe JD, Schaub T, et al. Value of imaging techniques of lumbar intervertebral disk prolapse. Z Orthop Ihre Grenzgeb 1994;132:371–6. 11. Patel N. Surgical disorders of the thoracic and lumbar spine: a guide for neurologists. J Neurol Neurosurg Psychiatry 2002;73(Suppl. I):i42–8. 12. Milette PC. Classification, diagnostic imaging, and imaging characterization of a lumbar herniated disk. Radiol Clin N Am 2000;38:1267–92. 13. Herzog RJ. The radiologic assessment for a lumbar disc herniation. Spine 1996;21(24S):19S–38S. 14. Vroomen PC, De Krom MC, Wilmink JT, et al. Diagnostic value of history and physical examination in patients suspected of lumbosacral nerve root compression. J Neurol Neurosurg Psychiatry 2002;72:630–4. 15. Pfirrmann CW, Dora C, Schmid MR, et al. MR Image-based grading of lumbar nerve root compromise due to disk herniation: reliability study with surgical correlation. Radiology 2004;230:583–8. 16. Nygaard OP, Jacobsen EA, Solberg T, et al. Nerve root signs on postoperative lumbar MR imaging. A prospective cohort study with contract enhanced MRI in symptomatic and asymptomatic patients one year after microdiscectomy. Acta Neurochir (Wien) 1999;141:619–23. 17. Jarvik JJ, Hollingworth W, Haegerty P, et al. The longitudinal assessment of imaging and disability of the back (LAIDBack) study. Spine 2001;26:1158–65. 18. Stam J. Consensus on diagnosis and treatment of lumbosacral root entrapment syndromes. Ned Tijdschr Geneeskd 1996;140:52. 19. Prokop M. General principles of MDCT. Eur J Radiol 2003;45:S4–S10. 20. Prokop M. Multislice CT: technical principles and future trends. Eur Radiol 2003;13 (Suppl. 5):M3–13. 21. Jensen MC, Brant-Zawadski MN, Obuchowski N, et al. Magnetic resonance imaging of the lumbar spine in people without back pain. N Engl J Med 1994;331:69–73. 22. Van Rijn JC, Klemetso N, Reitsma JB, et al. Observer variation in MRI evaluation of patients suspected of lumbar disk herniation. AJR Am J Roentgenol 2005;184:299–303. 23. Efron B, Tibshirani RJ. An introduction to the bootstrap. Chapman & Hall, 1993. 24. Cihangiroglu M, Yildirim H, Bozgeyik Z, et al. Observer variability based on the strength of MR scanners in the assessment of lumbar degenerative disc disease. Eur J Radiol 2004;51:202–8. 25. Bartynski WS, Lin L. Lumbar root compression in the lateral recess: MR imaging, conventional myelography, and CT myelography comparison with surgical confirmation. AJNR Am J Neuroradiol 2003;24:348–60. 26. Melendez JC, McCrank E. Anxiety-related reactions associated with magnetic resonance imaging examinations. JAMA 1993;270:745–7. 27. Kanal E, Shellock FG, Talagala L. Safety considerations in MR imaging. Radiology 1990;176:593–606.

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Initial experiences of using an automated volumetric measure of breast density: the standard mammogram form 1

M JEFFREYS,

PhD,

2

R WARREN,

FRCR,

3

R HIGHNAM,

PhD

and 4G DAVEY SMITH,

DSc

1

Centre for Public Health Research, Massey University, Private Box 756, Wellington, New Zealand, Addenbrooke’s Hospital, Cambridge, 3Mirada Solutions Ltd, Oxford and 4Department of Social Medicine, University of Bristol, UK

2

ABSTRACT. Limitations of area based measures of breast density have led several research groups to develop volumetric measures of breast density, for use in predicting risk and in epidemiological research. In this paper, we describe our initial experiences using an automated algorithm (standard mammogram form, SMF) to estimate the volume of the breast that is dense from digitized film mammograms. We performed analyses on 3816 mammograms of 626 women, who were part of the Glasgow Alumni Cohort and had mammograms taken within the Scottish Breast Screening Programme between 1989 and 2002. Absolute volume of dense breast tissue (SMF volume) and the percentage of the volume of the breast that is dense (SMF%) were calculated. The median (interquartile range) of SMF volume was 66 cm3 (48 to 98), and of SMF% was 23.4% (18.6 to 29.7). SMF%, but not SMF volume, was positively related to a six category classification (SCC) of visually assigned area-based breast density (increase in ln(SMF%) per category increase in SCC: 0.04% (95%CI: 0.03–0.05). The SMF algorithm produced lower SMF volume for craniocaudal (CC) compared with mediolateral oblique (MLO) views, but CC/MLO differences for SMF% were small. The mean right/left difference for ln(SMF volume) was 20.027 cm3 (95% confidence interval (CI) 20.044 to 20.009) and of ln(SMF%) was 0.005% (95% CI 20.008% to 0.019%). We present these initial data as a background for future analytical work using SMF.

The magnitude of the relationship between breast density and breast cancer [1] has led to recognition that breast density may have potential use as a biomarker for breast cancer risk [2]. Breast density is conventionally estimated by using the area of the mammogram that appears to be ‘‘glandular’’ and dividing by the total breast area. This is performed either visually, using classifications such as the Wolfe system [3], or using a computer-based thresholding technique [1]. Each of these classifications results in a measure of breast density which is a strong determinant of breast cancer risk. Several recognized breast cancer risk factors are positively related to area measures of breast density, including height [4, 5], parity [6] and age at first birth [7]. There are two notable exceptions to this, namely age and body weight, both of which are positively related to breast cancer risk, but inversely related to breast density [5, 8]. An inverse association between body mass index (BMI) and percent breast density is inevitable, since inherent in the definition of these measures is the assumption that fatty tissue is non-dense, and women with a high BMI have higher amounts of fat in the breasts. In addition to this discrepancy, there are other concerns associated with the visual methods of assessing breast density. These include: (i) the subjectivity of visual measures; (ii) variations in visual density with breast compression and X-ray exposure; (iii) consideration of area measures of the breast, despite its three-dimensional structure. 378

Received 6 April 2005 Revised 16 July 2005 Accepted 1 September 2005 DOI: 10.1259/bjr/24769358 ’ 2006 The British Institute of Radiology

Due to these concerns regarding area-based measures, recent research has been directed towards volume-based measurements, which try to model the volume of glandular tissue. We describe here initial results from the Standard Mammogram Form (SMF2) technology, version 2.2, a fully automated objective measurement tool to estimate the volume of glandular tissue in the breast from a mammogram [9, 10]. The SMF algorithm explicitly considers breast compression, exposure and tube voltage, and computes two volumetric measures of breast density, (i) the absolute volume (cm3) of the breast that is dense (SMF volume) and (ii) the percentage of the volume of the breast which is dense (SMF%). SMF is different from other volumetric research methods in that it incorporates a full physics model rather than using step-wedges in each image [11, 12].

Materials and methods The data presented were obtained from women who are part of the Glasgow Alumni Cohort, which has been described in detail elsewhere [13]. The included women were students who were registered at the University of Glasgow during the period 1948–1968 and who attended an annual medical examination at the Student Health Service. Surviving cohort members were contacted by postal questionnaire in 2001. Those women who replied to the questionnaire and were living in Scotland were asked to give consent for access to The British Journal of Radiology, May 2006

Volumetric breast density using SMF

screening mammograms taken under the Scottish Breast Screening Programme (SBSP) between 1989 and 2002. Women were informed that their mammograms would be digitized. All screening mammograms for each woman were retrieved from the eight breast screening centres and were digitized on site with a Canon FS300 digitizer scanner at a resolution of 100 mm with 8-bit precision by a single radiographer. Data on exposure and tube voltage were collected. Both the postal questionnaire survey and the acquisition of digital mammograms received ethical approval from the Multi-centre Research Ethics Committee (Scotland). For the visual assignment of density categories, scanned images were displayed at 300 mm resolution on a flat-panel display system. At this resolution, the images were about the same size as a mammogram film. All images were displayed to appear as if viewed on a light-box. No other adjustment or image post-processing was applied during the reading period. We have previously reported on the similarity in density measures obtained when these assessments are made from the digitized image compared with the original film [14]. Visual density measures were made by one radiologist experienced in density assessment (RW) using a sixpoint categorical scale of the percentage of the breast area that appeared dense. The categories were: 0%, 1– 10%, 11–24%, 25–49%, 50–74% and >75% and are referred to in this paper as the six category classification (SCC). These scales were chosen to make our work comparable with that of other researchers [15]. All mammograms for each woman were presented consecutively to the radiologist. The volume of the mammogram that appeared dense was estimated using SMF, v2.2. This is a computer algorithm which models the image formation process to decompose the breast into fatty and non-fatty tissue. It achieves this through estimating the thickness of dense tissue in each column of tissue between each pixel on the projected image and the X-ray source. The image is then standardized to remove the dependence on the parameters which were used to form the image. The input parameters required are side (left or right), view (craniocaudal (CC) or mediolateral oblique (MLO), current time product (mAs), tube voltage (kVp) and filter and target materials. The algorithm is fully automated, only requiring user intervention if there has been a data entry error, for example if a right sided mammogram was entered as left sided, since this causes the breast segmentation algorithm to fail. Detailed explanations of the physics behind the model have been published previously [9, 10].

Statistical analyses Descriptive analyses were performed on all mammograms of all women. Mann-Whitney tests were used to test the differences in density measures obtained from left and right mammograms, and from different views (CC and MLO). Because of the log-normal distribution of the data, both SMF and SMF% were log transformed prior to analysis, and the natural log of these measures was used in the regression models. The new variables are The British Journal of Radiology, May 2006

referred to as ln(SMF volume) and ln(SMF%). Random effects linear regression models were used to estimate associations between age and SMF density measures, taking into account the clustered nature of the data (several mammograms per woman). Further analysis was based on paired mammograms (i.e. left vs right, CC vs MLO) taken on the same day. For these, the first visit per woman was used, since it is at this visit that both views have been routinely performed since 1994, before two-view mammography became routine at all visits by 2003. Because of the continuous nature of the SMF data (both volume and percentage), Bland-Altman plots [16] of the natural log of the SMF measures were used instead of kappa statistics to assess the agreement between paired data. These plot the mean difference between pairs (expected to be zero) against the average value (ln(SMF volume) or ln(SMF%)) of that pair.

Results There were 3566 women in the original Glasgow Alumni Cohort, of whom 2169 (61%) were sent a postal questionnaire in 2001. These were the women who could be traced through the National Health Service Central Register and were still alive. The response rate was 59% (n51285). Of the respondents, 935 women (73%) were still living in Scotland. 277 of these women (30%) had never had a screening mammogram, and two women refused access to their films. The SMF algorithm was run on all 3968 mammograms belonging to 649 of the remaining 656 women (films of seven women were omitted inadvertently). The programme failed on one image and produced a result classified as ‘‘not excellent’’ for 29 (0.7%) further images. 23 (3.5%) further women (122 images) were excluded as they reported having had breast cancer in the 2001 questionnaire. Analyses are based on the remaining 626 women with 3816 mammograms. The median age at first breast screening was 53.6 years (range 40.0–71.5 years). Eight women were over 65 years at the time of their first mammogram. The median (interquartile range (IQR)) of the absolute volume of SMF was 66 cm3 (48–98 cm3), and of SMF% was 23.4% (18.6– 29.7%), see Figure 1. There was a non-linear relationship between the total volume of the breast and the volume of the breast which was dense, see Figure 2. In women with small overall breast volume, the absolute volume of SMF dense tissue was low, but the percentage of the volume of the breast which was dense was variable. In women with larger breast volume, the absolute volume of SMF dense tissue was variable, whereas the percentage of the volume of the breast which was dense tended to be smaller. Assessment of the association between the two SMF measures and the SCC system showed no relationship between absolute SMF volume and SCC, but a strong positive association between SMF% and SCC (Table 1). Each category increase in SCC was associated with a 0.04% (95%CI: 0.03–0.05) higher ln(SMF%). There were small but significant associations between each of the two SMF measures with age at mammography. The regression coefficient per year older, based on the 379

M Jeffreys, R Warren, R Highnam and G Davey Smith

Figure 1. Distribution of the volume of dense breast tissue in 3816 mammograms.

ln(absolute SMF volume) was 20.008 cm3 (95%CI: 20.010 to 20.005). For ln(SMF%), the regression coefficient was 20.021% (20.023 to 20.020). Comparisons of median SMF values stratified by mammogram view (CC vs MLO) indicated that the SMF algorithm consistently produced lower SMF volume for CC compared with MLO views (Table 2). The CC/MLO differences for SMF% were smaller. Comparing left and right sides, there was a small difference in the median absolute SMF volume, being slightly larger for left than right breasts, but no difference for SMF%. Correlations between the paired mammograms taken on the same day were high. For mammograms taken on the same day, the left–right correlation was high, r50.92 (p,0.001) for ln(SMF volume) and r50.85 (p,0.001) for ln(SMF%). These correlations were equally high for CC and MLO views. Bland-Altman plots (Figure 3) showed that although left/right agreement was good for the majority of women, there were some women whose values lie well outside the mean reference range (¡2 standard deviations of the mean difference between the left/right measures). The mean difference (right minus left) of the ln(SMF volume) was 20.027 cm3 (95% confidence interval (CI) 20.044 to 20.009) and of ln(SMF%) was 0.005% (95% confidence interval 20.008% to 0.019%). 380

Figure 2. Relationship between total breast volume and standard mammogram form (SMF) volume/SMF% in 3816 mammograms.

Discussion This study describes our initial experiences of using a new method to assess mammographic breast density, in which we estimate the proportion of the volume of the breast that is dense and the absolute volume of dense tissue. This new method is fully automated, and produced a useable density value for over 99% of images. We have shown that the percentage measure

Table 1. Association between area percent density and standard mammogram form (SMF) density in mammograms of 649 women SCC

0% 1–10% 11–24% 24–49% 50–74% >75%

n

359 542 649 822 753 625

Absolute SMF volume (cm3)

SMF%

median

IQR

median IQR

69 67 64 69 71 58

53 49 46 51 50 40

to to to to to to

103 95 85 99 107 93

16 17.6 19.6 23.7 29.2 34.4

14.1 to 18.1 15.7 to 20.4 17.7 to 22.1 21.1 to 26.8 25.4 to 33.1 29.1 to 40.4

SCC (Six category classification) is based on a visual assessment of the percentage of the area that appears dense. IQR (interquartile range).


Volumetric breast density using SMF Table 2. Association between view, side and standard mammogram form (SMF) density in mammograms of 649 women View

CC MLO p-value Side Left Right p-value

Absolute SMF volume (cm3)

SMF %

Median

IQR

Median IQR

625 3191

58 67 ,0.001

42–94 49–98

25.4% 20.3–32 23.0% 18.3–29 0.098

1921 1895

67 65 ,0.001

49–98 47–97

23.3% 23.4% 0.69

n

18.7–29.5 18.6–29.8

IQR, interquartile range.

is closely associated with frequently used visual assessments of density. We had no gold standard against which to compare the absolute volume of density, but the good agreement between left and right SMF estimates provides an assurance regarding the internal consistency of the method. Unlike the visual and semi-automated systems to assess breast density, which have been related to breast

Figure 3. Bland-Altman plots showing agreement between left/right pairs of mammograms, 626 women. Note: BlandAltman plots of the difference against the mean of left and right measures, with horizontal lines showing the mean difference of 0 and limits of agreement (¡2 standard deviations). Both standard mammogram form (SMF) and SMF% have been log transformed for these analyses.


cancer in several studies [1], SMF has not yet been validated directly against the incidence of breast cancer. Combining a biological understanding of breast density with knowledge of the physics of mammogram acquisition, we believe that this is likely to be as strong a predictor of breast cancer as the more conventional methods of mammographic density assessment. Since the estimation of density using this method is entirely objective, we expect that the non-differential misclassification during visual estimation, which attenuates associations between density and breast cancer, will not occur. Associations between SMF and breast cancer may therefore be of greater magnitude than the four to six fold risks reported between extreme categories of visually estimated or computer-assisted methods [1]. A limitation of all currently proposed volumetric systems, including SMF, is that they fail to acknowledge that the non-fatty component consists of groups of fibrous, glandular and other tissues together. It has been suggested that the (possibly unconscious) acknowledgment of these structures by radiologists when assigning the visual percent density may explain the power of these assessments in determining breast cancer risk. Comparison of visual area-based density measures have shown that correlations between these and MR images are high [17], and of similar magnitude to the associations which we report. Although correlation coefficients are not recommended as measures of assessing agreement [16, 18], we present these results as a comparison with those reported in previous studies. For example, in a small sample of pre-menopausal women, Pearson correlation coefficients were between 0.86 and 0.96 [19], very similar to the results we present. Recent data suggest that the association between breast density and breast cancer risk is similar in pre- and post-menopausal women [20]. We found that the SMF algorithm appears to estimate higher SMF% values in CC compared with MLO mammograms, but lower SMF volume. This may be because the MLO, but not the CC, view includes the axillary tail, which is primarily composed of fat. Furthermore, the MLO image captures all the tissue at the very back of the breast, so has a higher SMF volume, but since that tissue is often mostly fatty, the SMF% for CC is higher. If SMF can be proven in validated studies to be a useful marker of breast cancer risk, its potential is enormous. Unlike some methods of measuring volumetric breast density [11, 12], SMF does not require step wedges to be included during the mammogram acquisition. We have shown that it can be used with historical mammograms, which is a strong feature for use in epidemiological studies. Although SMF does currently require calibration data such as mAs and kVp to be known, work is well underway to remove this requirement (RH, personal communication, 2005). Development that would enable volume measurements of this kind without the need for calibration data would enable wider use for epidemiology in a multicentre setting. Despite continuing reported associations between risk factors, breast density and breast cancer [2, 20], there is a need for such work to be refined. First, we propose that the relative amounts of dense and non-dense tissues should be considered as two separate outcomes, with 381

M Jeffreys, R Warren, R Highnam and G Davey Smith

both routinely reported. This has been previously suggested [5], but not adhered to in the majority of published studies. This will allow better modelling of associations between risk factors and density, without the potential confounding influence of measures of body fat on the results. We have found that BMI is positively related to SMF volume [21], as would be expected if SMF volume is thought of as a proxy marker for breast cancer risk. Second, investigations into the biological mechanisms underlying associations between breast density and breast cancer risk are needed. Some work in this area has begun. For example, insulin-like growth factor (IGF) and its main binding protein IGFBP-3 have both been related to breast density [22, 23]. In summary, our findings suggest that the novel technique of estimating the volume of dense breast tissue, which involves modelling mammographic breast density using a fully automated system, has the potential to be a marker of breast cancer risk available for use in large epidemiological studies. Work is underway to investigate whether SMF can predict breast cancer risk.

7.

8.

9. 10.

11.

12.

13.

Acknowledgments We are grateful to the women who participated in this study for allowing access to their mammograms, to Pat Forrest who performed the digitization and for the assistance of all the Scottish Breast Screening Centres. This work was undertaken while Dr Jeffreys (neé Okasha) was employed at the University of Bristol. Dr Highnam is employed by Mirada Solutions, where he developed the SMF algorithm. We are grateful for the financial support provided by Breast Cancer Campaign, Breast Cancer Research Trust and World Cancer Research Fund International. The Centre for Public Health Research is supported by a programme grant from the Health Research Council of New Zealand.

References 1. Boyd N, Lockwood G, Byng J, Tritchler D, Yaffe M. Mammographic densities and breast cancer risk. Cancer Epidemiol Biomarkers Prev 1998;7:1133–44. 2. Heine JJ, Malhotra P. Mammographic tissue, breast cancer risk, serial image analysis, and digital mammography. Part 1. Tissue and related risk factors. Acad Radiol 2002;9:298–316. 3. Wolfe JN. Breast patterns as an index of risk for developing breast cancer. AJR Am J Roentgenol 1976;126:1130–7. 4. Gram I, Funkhouser E, Tabar L. Anthropometric indices in relation to mammographic patterns among peri-menopausal women. Int J Cancer 1997;73:323–6. 5. Boyd NF, Lockwood GA, Byng JW, Little LE, Yaffe MJ, Tritchler DL. The relationship of anthropometric measures to radiological features of the breast in premenopausal women. Br J Cancer 1998;78:1233–8. 6. Vachon CM, Kuni CC, Anderson K, Anderson VE, Sellers TA. Association of mammographically defined percent

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Occupational radiation doses in interventional cardiology: a 15-year follow-up 1,2

E VANÕ, PhD, MD

PhD,

1

L GONZALEZ,

PhD,

1,2

J M FERNANDEZ,

BSc,

3

F ALFONSO,

PhD, MD

and 3C MACAYA,

1

Department of Radiology, Complutense University Medical School 28040 Madrid, Spain, 2San Carlos University Hospital, Medical Physics Service and 3Cardiovascular Institute, 28040 Madrid, Spain ABSTRACT. This report describes occupational radiation doses of interventional cardiologists over 15 years and assesses action undertaken to optimize radiation protection. Personal dosimetry records of nine staff cardiologists and eight interventional cardiology fellows were recorded using personal dosemeters worn over and under their lead aprons. The hospital in which this study was conducted currently performs 5000 cardiology procedures per year. The hospital has improved its facilities since 1989, when it had two old-fashioned theatres, to include four rooms with more advanced and safer equipment. Intensive radiation protection training was also implemented since 1989. Initially, some individual dose values in the range of 100– 300 mSv month21, which risked exceeding some regulatory dose limits, were measured over the lead apron. Several doses in the range of 5–11 mSv month21 were recorded under the apron (mean510.2 mSv year21). During the last 5 years of the study, after the implementation of the radiation protection actions and a programme of patient-dose optimization, the mean dose under the apron was reduced to 1.2 mSv year21. Current mean occupational doses recorded under the lead apron are 14% of those recorded during 1989–1992 and those recorded over the apron are 14-fold less than those recorded during 1989–1992. The regulatory dose limits and the threshold for lens injuries might have been exceeded if radiation protection facilities had not been used systematically. The most effective actions involved in reducing the radiation risk were training in radiation protection, a programme of patient-dose reduction and the systematic use of radiation protection facilities, specifically ceiling-suspended protective screens.

Radiation exposure is a significant concern for interventional cardiologists (ICs) because workloads and the complexity of procedures have increased over the past few years without a corresponding increase in the number of specialists [1]. Although reduced scatter radiation in catheterization laboratories compared with that in old X-ray system laboratories, improved radiological protection facilities, and better, more inclusive radiation protection training for ICs have substantially reduced the risk of radiation exposure, the complexity and number of procedures have increased. Therefore, interventional cardiology is recognized as a high-radiation-risk practice [1–3], and evaluation and follow-up of occupational doses should be considered an important part of quality assurance (QA) programmes. Address correspondence to: Prof. Luciano Gonzalez. This study was partially funded by the European Commission 5th Framework Programme, Contract DIMOND FIGM-CT-2000-00061, the Spanish Department for Science and Technology (project BFI2003-09434) and the Spanish Nuclear Safety Council. Validation of some results with TLD chips was carried out with experimental equipment partially funded with EC FEDER resources.


Received 11 May 2005 Revised 1 August 2005 Accepted 1 September 2005 DOI: 10.1259/bjr/26829723 ’ 2006 The British Institute of Radiology

Several aspects of radiation safety in the practice of cardiology have been addressed by the American College of Cardiology in a consensus document [4]. The UNSCEAR 2000 report [5] states that fluoroscopic procedures are by far the largest source of occupational exposure in medicine. Cardiac catheterization, in particular, can represent a major source of exposure. A study performed in the UK [6] indicated that ICs receive a mean annual dose of 0.4 mSv, twice that received by radiologists and many times that received by nurses and technicians. There are substantial differences in occupational doses between cardiac laboratories [7–10]. This is caused by differences in X-ray systems (old film-based systems versus digital units) and their particular settings, levels of training in radiation protection, frequency of use of radiation protection facilities and personal dosemeters, and workloads of specialists. Renaud et al [11] described a 5-year follow-up of the radiation doses received by the in-room personnel of three cardiac catheterization laboratories and concluded that some workers may have exceeded the occupational limit for the lens of the eye. Lens injuries have been 383

E Vanõ, L Gonzalez, J M Fernandez et al

reported for several interventional radiology suites in which radiation protection conditions were not appropriate to the level of risk [12]. This report describes occupational radiation doses from interventional cardiology in a university hospital over a period of 15 years and the actions that were taken to optimize radiation protection. Data were gathered from a dosemeter worn on the trunk of the body under the apron and a dosemeter worn outside the apron, as recommended by the International Commission on Radiological Protection (ICRP) [1].

Methods and materials Follow-up of IC’s personal dosimetry records was performed in a university hospital currently performing more than 5000 procedures per year in four catheterization laboratories with nine staff cardiologists and eight fellows. In 1989, this interventional cardiology service used two old-fashioned X-ray units. In 1994, a Philips Optimus M-200 Poly C X-ray unit (Philips, Best, The Netherlands), installed in 1988, was upgraded and an old CGR unit was exchanged for a Philips Integris HM-3000. In 2000, two new Philips Integris H-5000 units were installed. All systems now have protective screens suspended from the ceiling. This radiation protection tool, which had previously not been installed in one of the rooms, was not used regularly by all specialists until they were made aware of its importance. In addition, lead aprons, thyroid protectors and lead glasses were also available and are used routinely at present (with a few exceptions). Two personal dosemeters with thermoluminescent dosimetry chips, as recommended by the radiation protection service of the hospital, were used for occupational dosimetry: one was worn on the trunk of the body under the apron and the other was worn outside the apron at the level of the collar or the left shoulder. A dosemeter under the apron provides an estimate of the dose to the organs of the shielded region. A dosemeter worn outside the apron supplies an estimate of the dose to the organs of the head and neck, including the thyroid and lenses of the eyes (if unshielded), but greatly overestimates the doses to organs of the trunk. Results obtained from both dosemeters were used to estimate the occupational effective dose as recommended by the NCRP [13] and ICRP [1]. Dosemeters were read monthly by a public dosimetry service accredited and audited by the National Regulatory Authority. Before 1992, training in radiation protection for ICs was scant, if performed at all. Subsequently, a radiation protection training programme was initiated in accordance with national regulations [14]. Of the staff cardiologists working in the centre, 90% attended the courses and were accredited in radiation protection, as required by the National Regulatory Authority. Some new cardiologists, especially fellows, did not attend the courses. New regulations in force since 1999 [15] require a second level of radiation protection training for interventionalists, which includes training in radiation protection of patients and QA, as recommended by the ICRP [1]. Training in radiation protection of patients is also required by European Directive 43/97/EURATOM [16]. 384

An interactive CD-ROM, co-sponsored by the European Commission [17], is used to provide radiation protection training for residents and fellows who commence work in interventional suites during the intervals between radiation-protection training courses. A copy of this CD-ROM is given to all new doctors on commencement of duty at the hospital’s interventional cardiology service. In addition, refresher sessions on radiation protection are presented periodically. Detailed analysis of personal dosimetry records of IC personnel is conducted every month. This is followed by individual interviews with persons exposed to monthly doses greater than 1.0 mSv under the apron (1/20 of the annual effective dose limit) or greater than 7.5 mSv over the apron (1/20 of the annual lens dose limit). In addition, a progressive audit programme was implemented to detect high patient doses, facilitate clinical follow-up in cases of likely skin radiation injury and to implement corrective action when necessary. Since 1999, a national standard [15] stipulates that patient doses in interventional procedures must be estimated and recorded. Because this patient-dose audit has reduced patient doses, occupational doses have also been reduced [2]. Personal dosimetry services typically provide monthly estimates of Hp(10) (the dose equivalent in soft tissue at 10 mm depth), which is usually compared with the annual limit of effective dose and with the eye lens limit [18], and Hp(0.07) (the dose equivalent in soft tissue at 0.07 mm depth) [18]. Usually, no significant differences between values are found in cardiac catheterization suites. The values reported in this paper are for estimates of Hp(10) obtained from personal dosimetry readings. The effective dose, E, can be estimated [13] from the dosemeter values for Hw (under the apron at the waist, although this position is not critical) and Hn (above the apron at the neck) from the equation: E~0:5Hw z0:025Hn NCRP report 122 [13] contains specific recommendations for calculating the effective dose when protective aprons are worn during diagnostic and interventional medical procedures involving fluoroscopy. In addition to the above formula, it states that the effective dose can be estimated as Hn/21 if only one dosemeter is worn on the neck outside the apron.

Results The data from occupational dosimetry were allocated to one of three periods for purposes of analysis.

First period (1989–1992): investigation of high dose values and implementation of a customized radiation protection programme Table 1 shows the findings from this period. Most values were in the range of 100–300 mSv month21, but in one case a dose of 1600 mSv month21 was recorded by the left shoulder dosemeter outside the lead apron. Values in the range of 5–11 mSv month21 were recorded The British Journal of Radiology, May 2006

Occupational radiation doses in IC Table 1. Individual monthly high values of personal dose equivalent Hp(10) and total Hp(10) values under apron during the year (except for cases indicated in the footnotes). Capital letters and numbers in the staff column are an internal code allowing traceability of the reported data Staff member

Year

Max. mSv/month

Total Hp(10) (under apron) (mSv)

I1 Senior cardiologist B1 Senior cardiologist F1 Senior cardiologist G1 Senior cardiologist R1 Senior cardiologist G1 Senior cardiologist G1 Senior cardiologist A1 Senior cardiologist F2 Senior cardiologist R1 Fellow cardiologist B1 Fellow cardiologist C1 Resident

1989 1989 1989 1990 1990 1991 1992 1992 1992 1992 1992 1992

51 8 4.6 62 65 346 180 155 54 1640 185 179

7.4 27.8a 12.3 5.2 9.2 27 4.2 7.1 23.7 47b 3.1c 11

(over apron) (under apron) (under apron) (over apron) (over apron) (over apron) (over apron) (over apron) (over apron) (over apron) (over apron) (over apron)

a

Values over apron not available. Incorrect use of the dosemeter cannot be excluded. Value under apron during the month receiving 1640 mSv over the apron. 47 mSv are 2.9% of the dose over the apron. Abnormal irradiation of the over-apron dosemeter was not demonstrated. c Only some months. b

under the apron. An initial complete evaluation of the radiation protection conditions of the catheterization laboratories was done, after which follow-up of abnormal values was investigated and corrective actions proposed. Consequently, the occupational medical service of the hospital advised some staff to abstain from catheterization duties for several months. The National Regulatory Authority was informed of these actions. Lens injuries would have occurred in those situations if the corrective actions had not been put into practice immediately.

Second period (1993–1998): consolidation of the radiation protection programme Training courses in radiation protection and seminars with ICs (including fellows) were commenced, new Xray systems with radiation protection facilities were installed, and a formal programme of quality control (QC) and strategies to reduce patient and staff doses were launched. Maximum monthly dose values (over the apron) ranged from 7 mSv to 10 mSv, with the exception of a new fellow, for whom high readings of up to 28 mSv month21 were recorded on two occasions. The highest yearly Hp(10) values under the apron were between 2 mSv and 3 mSv.

Third period (1999–2004): implementation of occupational radiation protection in the QA programme During this period, the frequency of the X-ray system QC programme increased from once yearly to two or three times per year. The old CGR X-ray system was removed in 1999. Full characterization was done by measuring patient entrance dose, image quality and scatter radiation levels for all fluoroscopy and cine modes. Closer contact with the maintenance engineers was established to customize the operation modes to fulfil the image quality requirements of the cardiologists while keeping doses as low as possible. Since this The British Journal of Radiology, May 2006

initiative, patient dose values were measured, recorded in a database and analysed periodically. Since 2000, the MARTIR training CD-ROM [17] has been distributed to new personnel joining the interventional cardiology service, and radiation protection refresher seminars are held two or three times per year. Individual real-time occupational dosimetry has also been implemented for some procedures. Electronic dosemeters (Unfors EED-30; www.unfors.se) measure the dose accumulated by the specialist throughout a procedure and the maximum dose rate, which provides information about the correct use of the protective screen. Maximum values recorded by dosemeters placed over the apron were lower during the third period than during the second period and ranged between 3 mSv month21 and 4 mSv month21. The maximum dose under the apron was generally 2 mSv month21, but some abnormally high values were recorded for specialists doing electrophysiology cardiac procedures (in service since 2000). A maximum over-the-apron dose of 26 mSv month21 was recorded for one specialist. The workloads during the three periods were similar: five to six procedures per day and room, shared between one to three cardiologists. Some of the fellows stayed at the hospital for short periods and often performed many procedures per day to improve their skills. Typical workloads were two to four procedures per day for staff and three to six procedures per day for fellows. Table 2 shows monthly doses before, during and after radiation protection training. Mean and median doses decreased significantly after the training courses. Unpaired t-test analysis revealed statistically significant differences between means for doses before and after the training periods. In two-tailed tests, p-values were less than 0.05 (p50.01 for 1996 vs 1991; p50.02 for 1995 vs 1992). Table 3 presents the annual dose values for the three periods. Only personal dosimetry records comprising all the monthly dose values were used. Data in which background dose values were recorded by the over-apron dosemeter of specialists who had a substantial workload were excluded from the analysis, as this indicated that the personal dosemeter had not been used. Between 20% and 30% of the cardiologists 385

E Vanõ, L Gonzalez, J M Fernandez et al Table 2. Relevant dose values (in mSv/month) under the lead apron, before, during and after the training courses on RP for numbers of IC specialists indicated Year

Sample

Range

1991 1992 1993 (training) 1994 (training) 1995 1996

8 11 7 12 10 13

1.9–26.5 0.9–24.2 1.0–4.4 0.6–13.0 0.7–4.1 0.4–5.8

neglected to send their personal dosemeters to the dosimetry service for processing every month. Differences between under-apron doses during 1989– 1992 and the other two periods were statistically significant (p,0.01), and a more significant difference was noted for values over the apron (p,0.004). Table 4 presents estimates of the transmitted fraction of energy across different lead aprons with thickness equivalents in the range of 0.25–0.5 mm lead. The IPEM software application [19] for spectra from 70 kVp to 90 kVp was used for calculations. The real spectra of scattered radiation in the catheterization rooms are difficult to determine. However, the Xray beam used for interventional cardiology in our laboratories (with the Philips Integris systems) typically ranges between 80 kVp and 110 kVp. Thus, the energy degradation in the scattering process would yield dosetransmitted fractions of between 3.3% and 8.3% for 0.25 mm lead aprons, between 1.5% and 4.9% for 0.35 mm lead aprons, and between 0.5% and 2.4% for 0.5 mm lead. Thus, a dose under the apron of between 0.5% and 8.3% of the values measured over the apron was considered compatible with the personal protection used and was regarded as a good indicator of proper use of personal dosemeters. The same criterion has been used to reject unreliable data, and the values from Tables 2 and 3 are fully compatible. In summary, the radiation protection programme during the 15-year period reduced the effective dose to cardiologists by one order of magnitude, avoiding cases of high individual doses. The real mean effective dose for cardiologists in our centre during the last 4 years of our study was 1.2 mSv year21, which is compatible with results recently reported by Delichas et al [20] (1.2– 2.7 mSv/procedure, a dose of 0.7–1.5 mSv year21 for a workload of 50 procedures per month).

Discussion Several questions arise from the results presented in this paper. First, it should be determined if the high doses

Mean¡SD

9.0¡9.3 7.4¡8.5 1.9¡1.0 3.0¡3.3 1.8¡1.2 1.5¡1.6

Median

5.1 3.7 1.6 1.6 1.3 0.9

reported during the years 1989–1992 are real dose values or incidental readings caused by inappropriate use of the dosemeters. In fact, the bulk of the results in Table 1 should correspond to real dose values received by the cardiologists during a period in which there was no culture of safety: ceiling-suspended screens were absent or unused, the X-ray systems were used in relatively high-dose fluoroscopy modes and film cine acquisition was done at 25 frames s21. The high dose values shown in Table 1 cannot be considered a consequence of the incorrect use of the dosemeters. All abnormal doses were reported to the doctors wearing the dosemeters and investigated with them, and no reason was found to suggest that incidental dosemeter irradiation occurred. For the 1640 mSv measured at the left shoulder of a visiting cardiologist in 1 month, it was not possible to prove any abnormal dosemeter irradiation. The dose measured by the dosemeter worn under the apron was 42 mSv in that month, 2.8% of the dose over the apron. This figure is compatible with the transmitted fraction across the lead apron (Table 4). Moreover, experimental measurements in one of the cardiology rooms used by a fellow simulating clinical conditions produced doses in good agreement with the dosemeter readings, taking into account the presumed work rate, fluoroscopy time and frame rate per procedure, and the mean scatter dose rate for a non-pulsed fluoroscopy mode. Distance is an important factor that could increase (or decrease) the scatter dose rate. A distance of 65 cm between the cardiologist and the isocentre has been supposed, but a variation of 15 cm nearer to the patient could increase the occupational dose by 70%. In addition, considering that the protective screens—typically equivalent to a shielding of 0.5–1.0 mm lead—can reduce the scatter dose by a factor of 100 if properly used, differences in the reported occupational doses in the scientific literature of two orders of magnitude measured over the lead apron are not surprising. In fact, Pratt and Shaw showed that the relationships between the cardiologist’s eye dose and factors such as the dose efficiency of the X-ray equipment, scattered-dose rates, examination

Table 3. Mean values (and standard deviation) in mSv/year of occupational doses of cardiologists during the periods referred to. The percentage of dose under apron in relation to the dose over apron is indicated in the Hp(10) ‘‘under apron’’ column between brackets Number of reliable data

Period

15 24 11

1989–1992 1993–1998 1999–2004

386

over apron

Hp(10) under apron

Effective dose (NCRP, using two dosemeters)

Effective dose (NCRP, using over-apron dosemeter)

259¡249 31¡15 18¡7

10.2¡8.6 (3.9%) 1.7¡1.1 (5.5%) 1.4¡0.4 (7.7%)

11.6 1.6 1.2

12.3 1.5 0.86


Occupational radiation doses in IC Table 4. Protection of different lead aprons for X-ray beams filtered with 3 mm Al and generated at the kVp indicated kVp

Protective apron mm lead equivalent

Fraction of energy transmitted (%)

90 90 90 80 80 80 70 70 70

0.25 0.35 0.50 0.25 0.35 0.50 0.25 0.35 0.50

8.3 4.9 2.4 5.7 3.0 1.3 3.3 1.5 0.5

protocols and workload are complex and vary from centre to centre [21]. Data considered reliable are scarce in Tables 1 and 2 because between 1989 and 1996, a significant number of cardiologists did not use the personal dosemeters during all procedures or overlooked the established procedure of sending the dosemeters to the medical physics service monthly. Compliance with the radiation badge policies is one of the main problems in many interventional cardiology services. Reported occupational dose values are often surprisingly low and the reason is not a high level of radiation protection, but a lack of use of personal dosemeters. McCormick et al [10] reported that after a mandatory radiation protection training programme, compliance with the radiation badge policy was only 36% in 1999, reaching 67% in 2000 and 77% in 2001 for physicians and nurse clinicians. Therefore, confidence in the mean dose values determined by the regional dosimetry services, and sometimes by the regulatory bodies, to assign occupational doses is open to discussion, as stated by UNSCEAR [5]. Another important point for improving the occupational dosemetry data is reporting dose results from dosemeters over and under the lead apron [1, 13] and combining them to calculate a more realistic effective dose. The over-apron dosemeter provides very useful information on the risk of lens injuries in interventional suites. The data in Table 3 show that differences up to 15% with the conventional under-apron dose approach can be found when considering the proposed NCRP formula [13]. Finally, one may wonder if occupational dose values as high as those reported for the period 1989–1993 could be reached with new X-ray systems and radiation protection facilities. Fortunately, the probability is low. Modern interventional cardiology X-ray systems have significantly decreased the radiation level for the patient and, as a consequence, the scatter radiation level. In addition, radiation protection facilities [8, 22], especially ceiling-suspended screens, are in common use, access to advice of medical physics experts is more frequent and dose parameters are included in QA programmes, among other positive factors. However, the workload of cardiologists and the complexity of procedures are increasing, and thus vigilance should be maintained. In the new X-ray laboratories, high levels of scattered dose are still measured and sometimes a surprising lack of training in radiation protection is the cause of avoidable and unjustified occupational over-irradiations. The British Journal of Radiology, May 2006

Conclusions Occupational doses measured on specialists who are routinely using their personal dosemeters show that the radiation protection level has significantly improved in the last decade. A reduction in the effective dose by a factor of 10 has been achieved. The most successful action to reduce occupational doses has been training in radiation protection. The use of ceiling-suspended protective screens in a systematic way by the cardiologists and the programme of patient dose reduction were important complementary actions. New X-ray equipment also contributed to further dose reductions, but its relative impact cannot be distinguished from the training effect in this study because of their interdependence. Another significant conclusion is that mean values of the occupational doses in catheterization laboratories could provide an incorrect estimate of the real radiological risk if some specialists are not using their personal dosemeters on a regular basis.

Acknowledgments The authors thank Mercedes Lago for her help in gathering dosimetry data.

References 1. ICRP Publication 85. Avoidance of radiation injuries from medical interventional procedures. Ann ICRP. Oxford, UK: Pergamon, Elsevier Science Ltd, 2000;30(2). 2. Vano E. Radiation exposure to cardiologists: how it could be reduced. Heart 2003;89:1123–4. 3. Finkelstein MM. Is brain cancer an occupational disease of cardiologists? Can J Cardiol 1998;14:1385–8. 4. American College of Cardiology. Radiation safety in the practice of cardiology. ACC expert consensus document. J Am Coll Cardiol 1998;31:892–913. 5. United Nations Scientific Committee on the Effects of Atomic Radiation. Sources and effects of ionizing radiation. UNSCEAR 2000 report to the General Assembly, with scientific annexes. Annex E. Occupational radiation exposures. (Available at www.unscear.org). New York: United Nations, 2000. 6. Hughes JS, O’Riordan MC. Radiation exposure of the UK population—1993 review. NRPB-R263 (1993). Quoted in UNSCEAR 2000 report. 7. Tsapaki V, Kottou S, Vano E, Komppa T, Padovani R, Dowling A, et al. Occupational dose constraints in interventional cardiology procedures: the DIMOND approach. Phys Med Biol 2004;49:997–1005. 8. Kuon E, Schmitt M, Dahm JB. Significant reduction of radiation exposure to operator and staff during cardiac interventions by analysis of radiation leakage and improved lead shielding. Am J Cardiol 2002;89:44–9. 9. Vano E, Gonzalez L, Guibelalde E, Fernandez JM, Ten JI. Radiation exposure to medical staff in interventional and cardiac radiology. Br J Radiol 1998;71:954–60. 10. McCormick VA, Schultz CC, Hollingsworth-Schuler V, Campbell JM, O’Neill WW, Ramos R. Reducing radiation dose in the cardiac catheterization laboratory by design alterations and staff education. Am J Cardiol 2002;90:903–5. 11. Renaud L. A 5-year follow-up of the radiation exposure to in-room personnel during cardiac catheterization. Health Physics 1992;62:10–5. 12. Vano E, Gonzalez L, Beneytez F, Moreno F. Lens injuries induced by occupational exposure in non-optimized

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interventional radiology laboratories. Br J Radiol 1998;71:728–33. NCRP report 122. Use of personal monitors to estimate effective dose equivalent and effective dose to workers for external exposure to low-LET radiation. Bethesda, MD: NRCP, 1995. Royal Decree 1891/1991, from the Industry and Energy Department, on installation and utilisation of X-ray devices for medical diagnostic purposes. In: Spanish State Official Bulletin January 3rd 1992: 138–48 [In Spanish]. Royal Decree 1976/1999, from the Health and Consumer Affairs Department, establishing quality criteria in radiodiagnostics. In: Spanish State Official Bulletin of January 29th 1999: 45891–900 [In Spanish]. Council Directive 97/43 Euratom, on health protection of individuals against the dangers of ionizing radiation in relation to medical exposure, and repealing Directive 84/ 466 Euratom. Official Journal of the European Communities No L 180, 9th July 1997:22–27. MARTIR (Multimedia and Audiovisual Radiation Protection Training in Interventional Radiology). CD-ROM. Radiation Protection 119, European Commission. Directorate

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General Environment, Nuclear Safety and Civil Protection. Luxembourg, 2002 (available for free from the Publication Department of the European Commission, Office for Official Publications of the European Communities, Luxembourg, e-mail: [email protected]). ICRU report 51. Quantities and units in radiation protection dosimetry. Bethesda, MD: International Commission on Radiation Units and Measurements, 1993. IPEM report 78. The Institute of Physics and Engineering in Medicine. Catalogue of Diagnostic X-ray Spectra and other data. Cranley K, Gilmore BJ, Fogarty GWA, Desponds L. Software prepared by Reilly AJ and Sutton D. 1997. Delichas M, Psarrakos K, Molyvda-Athanassopoulou E, Giannoglou G, Sioundas A, Hatziioannou K, et al. Radiation exposure to cardiologists performing interventional cardiology procedures. Eur J Radiol 2003;48:268–73. Pratt TA, Shaw AJ. Factors affecting the radiation dose to the lens of the eye during cardiac catheterization procedures. Br J Radiol 1993;66:346–50. Kuon E, Birkel J, Schmitt M, Dahm JB. Radiation exposure benefit of a lead cap in invasive cardiology. Heart 2003;89:1205–10.



A four-dimensional computer simulation model of the in vivo response to radiotherapy of glioblastoma multiforme: studies on the effect of clonogenic cell density 1

G S STAMATAKOS,

PhD,

1

V P ANTIPAS,

PhD,

1

N K UZUNOGLU,

PhD

and 2R G DALE,

PhD

1

In Silico Oncology Group, Microwave and Fibre Optics Laboratory, School of Electrical and Computer Engineering, National Technical University of Athens, 9 Iroon Polytechniou St., GR 157 80 Greece and 2Department of Radiation Physics and Radiobiology, Hammersmith Hospitals NHS Trust and Faculty of Medicine, Imperial College, Charing Cross Hospital, Fulham Palace Road, London W6 8RF, UK ABSTRACT. Tumours behave as complex, self-organizing, opportunistic dynamic systems. In an attempt to better understand and describe the highly complicated tumour behaviour, a novel four-dimensional simulation model of in vivo tumour growth and response to radiotherapy has been developed. This paper presents the latest improvements to the model as well as a parametric validation of it. Improvements include an advanced algorithm leading to conformal tumour shrinkage, a quantitative consideration of the influence of oxygenation on radiosensitivity and a more realistic, imaging based description of the neovasculature distribution. The tumours selected for the validation of the model are a wild type and a mutated p53 gene glioblastomas multiforme. According to the model predictions, a whole tumour with larger cell cycle duration tends to repopulate more slowly. A lower oxygen enhancement ratio value leads to a more radiosensitive whole tumour. Higher clonogenic cell density (CCD) produces a higher number of proliferating tumour cells and, therefore, a more difficult tumour to treat. Simulation predictions agree at least semi-quantitatively with clinical experience, and particularly with the outcome of the Radiation Therapy Oncology Group (RTOG) Study 83-02. It is stressed that the model allows a quantitative study of the interrelationship between the competing influences in a complex, dynamic tumour environment. Therefore, the model can already be useful as an educational tool with which to study, understand and demonstrate the role of various parameters in tumour growth and response to irradiation. A long term quantitative clinical adaptation and validation of the model aiming at its integration into the treatment planning procedure is in progress.

The rapid growth and resilience of tumours make it difficult to believe that they behave as random, disorganized and diffuse cell masses and suggests instead that they are emerging, opportunistic systems [1, 2]. If this hypothesis holds true, the growing tumour and not only the single cell [3] must be investigated and treated as a self-organizing complex dynamic system. This cannot be done with currently available simple in vitro/ in vivo models or common mathematical approaches. Therefore, there is a need for novel computational models to simulate the complexity of solid tumour growth and invasion, combining a range of disciplines including medical, biological, engineering and statistical physics research [1, 2]. Recent research efforts have focused on modelling of tumour response to various therapeutic modalities but, unfortunately, current models of dynamic processes need substantial improvement due to the complexity of the problem and the paucity of large series’ of clinical data. This paragraph provides a brief outline to several of the concepts and earlier research efforts. Duechting et al The British Journal of Radiology, May 2006

Received 7 June 2004 Revised 8 August 2005 Accepted 30 August 2005 DOI: 10.1259/bjr/30604050 ’ 2006 The British Institute of Radiology

[4, 5] introduced a simulation model which concerns only the in vitro case or the early avascular stages of small in vivo tumours and is based on a consideration of the distinct phases of the cell cycle. Kocher et al [6, 7] developed a simulation model of the development of a tumour spheroid and its response to radiosurgery. However the detailed geometry of the clinical tumour as depicted by imaging data has not been considered in this model. Instead, an equivalent spherical tumour is considered in place of the generally arbitrarily shaped actual tumour. Additionally, detailed cell cycle phase biology (phases G1, S, G2, M) has not been taken into account, with grouping of the cells into only proliferating and dormant classes being considered instead. None of the above mentioned models have been applied to large clinical tumours of varied geometrical shapes, and none of them simulates conformal shrinkage for an arbitrarily shaped clinical tumour undergoing treatment. In the tumour growth models presented by Kansal et al [1, 2], a discretising grid is used in which each geometrical cell is able to contain a large number of biological cells, but the 389

G S Stamatakos, V P Antipas, N K Uzunoglu and R G Dale

grid has not been used to discretise clinical tumours of arbitrary shape. Neither is the response of the tumour to irradiation addressed in this model. Swanson et al [8, 10] and Mandonnet et al [9] have developed clinically significant spatiotemporally models of tumour growth and invasion concerning glioblastoma multiforme (GBM). Nevertheless, although growth and invasion constitute fundamental phenomena related to GBM treatment optimization, the investigators have not focused on the radiobiological mechanisms underlying radiotherapy. Byrne et al and Alarcon et al [11–14] have developed mathematical models of avascular tumour growth and angiogenesis evolution pertinent mainly to the initial stages of tumour development. Although valuable insight can be gained using such models, extension to clinical voluminous tumours is not an a priori manageable task. An effort to overcome these shortcomings has previously been made by our group through the development of a four-dimensional patient-specific in vivo simulation model [15–17]. All parameters used in the model have already been defined and can be determined (in principle) experimentally or clinically. Therefore, use of new mathematically dictated parameters of ambiguous physical meaning is avoided. Furthermore, the authors believe that the development of an experimental and clinical biology based model provides substantial insight into the interdependence of the mechanisms involved, even if some parameters cannot currently be accurately quantified for individual patients. This work presents the latest advances and improvements of a four-dimensional, patient individualized, in vivo simulation model of tumour response to radiotherapy and discusses attempts at parametric and clinical validation. To this end, the outcome of a pertinent clinical study is exploited. It is also noted that an eventual combination of the model with the approaches of the previously mentioned research groups might be worth a careful investigation. In such a case, a more scientifically complete and clinically meaningful picture of the various aspects of tumour modelling might emerge. It is suggested that the work presented might also be considered a step towards shaping an emerging analytical-computational discipline of ‘‘In Silico Radiation Oncology’’.

Tumour cell distribution In the following, a brief outline of the model’s construction is given. 1. A discretizing cubic mesh is superimposed upon a three-dimensional reconstruction of the tumour, including its necrotic region and the surrounding anatomical features, based on the imaging data. 2. Within each geometrical cell of the mesh, a number of equivalence classes (compartments) of the contained biological cells are defined based on their distribution over the various phases within or out of the cell cycle. Sufficient registers are used in order to store the current state of each equivalence class (e.g. time spent in phase G1, etc.). 3. The mesh is scanned every hour. 390

4. The basic biological ‘‘laws’’ (metabolic activity, cell cycling, mechanical restrictions, cell survival probability following irradiation with dose D, etc.) are applied on each geometrical cell at each scanning. 5. A spatial and functional restructuring of the tumour may take place during each scanning as e.g. new cells may be produced (leading to differential tumour growth) or as existing cells may disappear (leading to differential tumour shrinkage). 6. The simulation predictions can be two- or threedimensionally visualized at any simulated instant of interest. The basic philosophy of our approach can also be found (http://www.in-silico-oncology.iccs.ntua.gr/) at [17].

Recent improvements Vessel network, neoangiogenesis and oxygen supply based on imaging data The observation that angiogenesis occurs around tumours was made nearly 100 years ago [18–21]. The hypothesis that tumours produce a diffusible ‘‘angiogenic’’ substance was put forward in 1968 [22, 23]. Mammalian cells require oxygen and nutrients for their survival, and functional cells must therefore be located within a distance of 100–200 mm from the nearest capillary blood vessels, which is the diffusion limit for oxygen. Vessels in an embryo are assembled from endothelial precursors and this primitive network subsequently expands by sprouting (angiogenesis) or intussusception, in which interstitial tissue columns are inserted into the lumen of pre-existing vessels and partition the vessel lumen [24]. Tumour vessels develop by sprouting or intussusception from pre-existing vessels. Circulating endothelial precursors, shed from the vessel wall or mobilized from the bone marrow, can also contribute to tumour angiogenesis [25, 26]. Tumour cells can also grow around an existing vessel to form a perivascular cuff. Without blood vessels, tumours cannot grow beyond a critical size or metastasise to another organ [27]. In contrast to normal vessels, tumour vasculature has the following distinct characteristics: 1. Tumour vessel ultrastructure is abnormal [27] 2. Tumour vessel ultrastructure is highly disorganized [27] 3. Tumour vessels are tortuous and dilated, with uneven diameter, excessive branching and shunts [27] 4. Tumour blood flow is chaotic and variable [27, 28], and leads to hypoxic and acidic regions in tumours [27, 29] 5. Vessel walls have numerous ‘‘openings’’, widened interendothelial junctions, and a discontinuous or absent basement membrane [27] 6. Tumour vessels are ‘‘leaky’’ and have tremendous heterogeneity in leakiness over space and time [27, 30, 31] Definition of the imaging based tumour layers dictates the number and the metabolic state of the individual The British Journal of Radiology, May 2006

Simulation of in vivo glioblastoma multiforme response to radiotherapy

biological cells included within each layer. During the simulation process, and in the case of tumour growth, the normal tissue capillaries are shifted away and tumour capillaries are generated in their place [27, 32]. Consequently, the new tumour cells are assumed to be sufficiently oxygenated and able to divide. A ‘‘proliferating layer’’ is assumed to exist between the external surface of the gross tumour volume and a hypothetical surface (HYP) enclosing its necrotic kernel and lying 1.5 mm further out. The tumour volume contained between HYP and the surface of the necrotic region has been assumed to contain large numbers of dormant G0 cells; therefore this is called the ‘‘G0 cell layer’’. This layer contains a substantial number of dormant cells around the necrotic area of the tumour as it appears on the imaging data. We have assumed that the clonogenic cell density (CCD) in the ‘‘proliferating layer’’ is two times the CCD in the G0 cell layer. CCD in the necrotic or dead cell layer of the tumour has also been assumed to be one fifth of the CCD in the proliferating cell layer.

Consideration of oxic and hypoxic cells The resistance of cells that are hypoxic at the time of therapy will influence the efficacy of treatment with radiation, chemotherapy and combined modality regimens. Tumour cell response to ionizing radiation is strongly dependent upon oxygen, any given dose killing substantially fewer hypoxic than oxic cells. The radiation dose that allows a particular level of survival tends to increase by the same factor at all levels of survival when oxygen is removed. This allows calculation of oxygen enhancement ratio (OER) for the same level of biological effect. For most cells, OER for X-rays is around 3.0 [6, 7, 33]. Some researchers [33–35] report that OER reduces for radiation doses to 3.0 Gy or less. In practical terms, within a tumour microenvironment the oxic cells are those which proliferate, whereas the hypoxic cells are dormant or G0 cells. Extensive work has been done to measure hypoxia in human brain tumours (especially for gliomas) [36–41]. Different LQ a and b parameter values for the oxic (G1, S, G2, M) and hypoxic (G0) cells are considered. The interrelation between the hypoxic and oxic LQ parameters is given by the following expressions [7]: aoxic ahypoxic ~ OER

ð1Þ

boxic OER2

ð2Þ

bhypoxic ~

Nygren and Ahnstrom [42] suggest that OER can range from 2.0 to 3.0, and Palcic et al [34], Stuschke et al [43] and Speke and Hill [44] reported a value for OER of 2.3, 2.7, 2.75, respectively. It should be noted that hypoxic cells in the clinical setting can become oxic again when either new microvasculature vessels have emerged in their vicinity, or when the space between them and the nearest blood vessels has been cleared of other cells already killed by irradiation. The British Journal of Radiology, May 2006

An advanced algorithm leading to conformal tumour shrinkage The tumour shrinkage process usually tends to behave as a conformal contraction [45]. The biological rationale for the cells to be ‘‘pulled’’ towards the centre-of-mass of the tumour is that the surrounding normal tissues exert a rather uniform pressure upon the tumour in such a way that the brain tends to recover its (physiological) normal shape (homeostasis). Nevertheless, deviations of this rule due to local inhomogeneities are to be expected. In order to satisfactorily simulate this process in conjunction with the rest of the simulation strategy, the ‘‘centre-of-mass algorithm’’ (CMA) is introduced. Its primary mission is to ‘‘pull’’ individual tumour cells towards the centre-of-mass of the entire tumour. It is emphasised that the term ‘‘centre-of-mass’’ is not used in a strict context as it refers to a tumour of uniform mass density. The model does not consider the eventual appearance and behaviour of new tumour foci out of the space occupied by the primary tumour due to infiltration of adjacent tissues. Any geometrical cell which is in two-dimensional side contact with at least one other geometrical cell of the main tumour mass is considered to belong to the main cohesive tumour mass. Let us suppose that the distance from the centre of an isolated geometrical cell C2 to the nearest geometrical cell occupied by the main tumour mass is larger than one geometrical cell. The content of the isolated geometrical cell tends to move towards the centre of the main cohesive tumour mass. Only shifts along the x+ or x- or y- or y+ or z- or z+ directions are allowed. During the application of CMA, the distance from the centre of C2 to the centre of the main tumour along each of the six possible directions is calculated. Subsequently, the least distance is selected. In cases where more than one distance has the same least value, the direction is selected randomly. After completion of the next discretization mesh scan, C2 will be closer to the tumour centre-of-mass and after a number of scans, C2 will be connected with the nearest cell of the main tumour mass.

A more realistic approach to tumour cell distribution A random number generator is used to produce a uniform cell distribution over the individual time units constituting each phase within or out of the cell cycle (e.g. S, G0), this being a more realistic approach. This is because cells within the same cell cycle phase are not considered to be synchronized during the initialization process.

Parameter value selection: the paradigm of in vivo glioblastoma multiforme The selected paradigm is a recently irradiated glioblastoma multiforme tumour. An oncology specialist delineates the gadolinium enhanced T1 weighted MRI imaging-based apparent boundary of the tumour as well as the boundary of its necrotic region. As a first 391


approach, the neovasculature field is assumed to coincide with the area of the tumour where pronounced metabolism is apparent on the imaging data. A cube defining the anatomical region of interest is superimposed on the three-dimensionally reconstructed tumour and surrounding anatomical features. The cube includes a volume discretizing mesh. The dimensions of each geometrical cell of the mesh, considered to be able to accommodate 106 biological cells [46] (NBC5106), are 1 mm61 mm61 mm. A number of researchers have focused on the measurement of the duration of cell cycle (Tc), especially for gliomas. Hoshino et al [47] have reported a mean Tc of 57 h. Crafts et al [48] suggested that Tc can range from 2 days to 3 days. Hoshino and Wilson [49] have mentioned a Tc of 75.6 h. Pertuiset et al [50] have found an average value of 1–2.5 days. The Tc values considered in this paper have been 24 h, 48 h and 72 h. Salmon et al [51] have suggested that proliferating tumour cells would spend their time in the various cell cycle phases as follows: fraction of time spent in G1: TG150.4 TC, S: TS50.39 TC, G2: TG250.19 TC and M: TM50.02 TC. After irradiation, most often the reproductively dead cells will continue to cycle for (usually) 1–3 divisions before their ultimate (biological) death. In the model developed by our group, reproductively dead cells are assumed to undergo two mitoses before their biological death [17]. Reproductively dead cells and their offspring which are still cycling are considered proliferating until their ultimate biological death because they are detectable through imaging modalities. This point emphasises the conceptual differences that may arise between an engineering and a medical physics/clinical approach to the same biological phenomenon. Exploratory simulation runs have shown that if still-cycling reproductively (but not yet biologically) dead cells are added to the unaffected proliferating cells, an increase in the number of the latter by up to a factor of 10 (1 log) is to be expected during a typical radiotherapy course. A standard fractionation scheme (2 Gy once a day, 5 days per week, 60 Gy in total) has been simulated. The LQ model parameters of the tumour have been assumed to be aoxic50.17 Gy21, boxic50.02 Gy22, and ahypoxic5 (0.17/OER) Gy21, bhypoxic5(0.02/OER2) Gy22 for a GBM with known mt p53 gene [52] and aoxic50.6 Gy21, boxic50.06 Gy22, ahypoxic5(0.6/OER) Gy21, bhypoxic5 (0.06/OER2) Gy22 for a GBM with known wild type (wt) p53 gene [7, 53]. Both the aoxic and boxic parameters are assumed to remain constant throughout the cell cycle. For visualization purposes, cells are ‘‘painted as dead’’ during the time interval between a lethal cell hit and necrosis or apoptosis. For the specific type of poorly differentiated tumour under consideration, and for simplification reasons, all non-clonogenic cells have been considered to be necrotic and sterile cells have not been taken into account. The contribution of the living non-clonogenic cells will be considered in a future version of the model. A typical tumour volume of 20 cm3 contains 4–5% clonogenic tumour cells [54], i.e. the typical CCD is around 107 cm23 [55]. Most calculations of the biological effect of radiation on tumours assume that the CCD is uniform. But in practice, tumours will almost certainly have a non-uniform CCD [56]. This factor has not been 392

modelled except by Brahme and Agren [57] and Webb and Nahum [56]. At present, as there is paucity of experimental data in vivo, a range of reasonable values has been assumed. Webb and Nahum [56] have reported that the variation of CCD across tumours is a very important factor, especially for brain tumours. Hence the CCD has been assumed to be 16104 cells mm23, 26104 cells mm23, 36104 cells mm23, 46104 cells mm23. Within each geometrical cell of the discretizing mesh, the initial distribution of the clonogenic cells throughout the cell cycle, the G0 and the necrotic phases depends on the layer of the tumour to which the geometrical cell under consideration belongs [17]. The cell loss factor [58] has been taken to be 0.3 as cell death products are removed from brain with substantial difficulty and it has been expressed as the sum of the cell loss factor due to necrosis (0.2) and the cell loss factor due to apoptosis (0.1). It is noted that estimates of the percentage of proliferating, dormant and dead cells in the various tumour layers are included in Stamatakos et al [17] and are based on a rather semi-quantitative representation of the cycling status of tumour cells depending on the imaging-based layer in which the considered tumour cells lie. For example, proliferating cells included in the dark tumour areas that appear on gadolinium enhanced T1 weighted MRI slices are expected to be scant. Concerning cell density, the standard assumption of 106 cells mm23 [46] has been made.

Validation and results In order to clinically evaluate the simulation model, several arm simulations of the Radiation Therapy Oncology Group (RTOG) study 83-02 [18] have been performed. The GBM imaging data considered throughout the whole paper have been used as the spatial basis for performing in silico experiments. The following typical parameter values have been adopted: OER53.0, clonogenic cell density 516CCD516104 cells mm23, cell cycle duration Tc530 h and the LQ parameters have been assumed to match the values of a GBM with mt p53 gene as previously mentioned. Figure 1a shows the total number of proliferating and dormant tumour cells as a function of time for the hyperfractionated (1.2 Gy twice daily to the dose of 81.6 Gy, ‘‘HF-81.6’’) and accelerated hyperfractionated (1.6 Gy twice daily to the dose of 54.4 Gy, ‘‘AHF-54.4’’) radiotherapy schedules. All schemes considered in this paper start on the first day of the radiotherapy course. HF-81.6 is completed on day 46 after initiation of treatment whereas AHF-54.4 is completed on day 23. Figure 1b depicts the total number of proliferating and dormant tumour cells as a function of time for the hyperfractionated (1.2 Gy twice daily to the dose of 76.8 Gy, ‘‘HF-76.8’’) and accelerated hyperfractionated (1.6 Gy twice daily to the dose of 48 Gy, ‘‘AHF-48’’) radiotherapy schedules. Both irradiation schedules start on the first day of the first week of treatment. HF-76.8 is completed on day 44 after initiation of treatment whereas AHF-48 is completed on day 19. According to the graphs, before completion of the AHF course, cell kill due to The British Journal of Radiology, May 2006


Figure 1. (a) Total number of proliferating and dormant tumour cells as a function of time for the hyperfractionated (1.2 Gy twice daily, 5 days per week to the dose of 81.6 Gy, ‘‘HF-81.6’’) and accelerated hyperfractionated (1.6 Gy twice daily, 5 days per week to the dose of 54.4 Gy, ‘‘AHF-54.4’’) radiotherapy schedules. HF-81.6 is completed on day 46 after initiation of treatment whereas AHF-54.4 is completed on day 23. In all fractionation schedules considered in this paper, no radiation is administered on Saturdays or Sundays. (b) Total number of proliferating and dormant tumour cells as a function of time for the hyperfractionated (1.2 Gy twice daily, 5 days per week to the dose of 76.8 Gy, ‘‘HF-76.8’’) and accelerated hyperfractionated (1.6 Gy twice daily, 5 days per week to the dose of 48 Gy, ‘‘AHF-48’’) radiotherapy schedules. Both irradiation schedules start on the first day of the first week of treatment. HF-76.8 is completed on day 44 after initiation of treatment whereas AHF-48 is completed on day 19.

AHF irradiation is more pronounced than cell kill induced by the HF scheme. This can be explained by the fact that a higher total dose has been administered to the tumour by the AHF scheme whereas, for the period under consideration, both schemes are characterized by the same time intervals between consecutive sessions. In cases where not all living cells have been killed by AHF irradiation, tumour repopulation is considerable so that, by the time the HF scheme is completed, living tumour cells and their progeny which have escaped AHF irradiation outnumber tumour cells which have escaped HF irradiation. Improved tumour control following HF irradiation in comparison with the AHF scheme is in agreement with the conclusions of the clinical trial RTOG 83-02.

Figure 2a shows the total number of proliferating and dormant tumour cells as a function of time for the hyperfractionated (1.2 Gy twice daily to the dose of 72 Gy, ‘‘HF-72’’) and accelerated hyperfractionated (1.6 Gy twice daily to the dose of 48 Gy, ‘‘AHF-48’’) radiotherapy schedules. HF-72 is completed on day 40 after initiation of treatment whereas AHF-48 is completed on day 19. Figure 2b shows the total number of proliferating and dormant tumour cells as a function of time for the hyperfractionated (1.2 Gy twice daily to the dose of 64.8 Gy, ‘‘HF-64.8’’) and accelerated hyperfractionated (1.6 Gy twice daily to the dose of 48 Gy, ‘‘AHF-48’’) radiotherapy schedules. HF-64.8 is completed on day 37 after initiation of treatment whereas AHF-48 is completed on day 19. Both irradiation

Figure 2. (a) Total number of proliferating and dormant tumour cells as a function of time for the hyperfractionated (1.2 Gy twice daily, 5 days per week to the dose of 72 Gy, ‘‘HF-72’’) and accelerated hyperfractionated (1.6 Gy twice daily, 5 days per week to the dose of 48 Gy, ‘‘AHF-48’’) radiotherapy schedules. HF-72 is completed on day 40 after initiation of treatment whereas AHF-48 is completed on day 19. (b) Total number of proliferating and dormant tumour cells as a function of time for the hyperfractionated (1.2 Gy twice daily, 5 days per week to the dose of 64.8 Gy, ‘‘HF-64.8’’) and accelerated hyperfractionated (1.6 Gy twice daily, 5 days per week to the dose of 48 Gy, ‘‘AHF-48’’) radiotherapy schedules. HF-64.8 is completed on day 37 after initiation of treatment whereas AHF-48 is completed on day 19. Both irradiation schedules start on the first day of the first week of treatment.


393


schedules start on the first day of the first week of treatment. According to the graphs, before completion of the AHF course, cell kill due to AHF irradiation is more pronounced than cell kill induced by the HF scheme. This can be explained by the fact that higher total dose has been administered to the tumour by the AHF scheme with the same fractionation as the one followed by the HF scheme for the period under consideration. In cases where not all living cells have been killed by AHF irradiation, tumour repopulation is considerable so that, by the time the HF scheme is completed, living tumour cells and their progeny which have escaped AHF irradiation outnumber tumour cells which have escaped HF irradiation. Figure 3 depicts the total number of tumour cells (proliferating, dormant and dead cells) as a function of time for the hyperfractionated (1.2 Gy twice daily to the dose of 76.8 Gy, ‘‘HF-76.8’’) and accelerated hyperfractionated (1.6 Gy twice daily to the dose of 48 Gy, ‘‘AHF48’’) radiotherapy schedules. Irradiation starts on the first day of the first week. HF-76.8 is completed on day 44 after initiation of treatment whereas AHF-48 is completed on day 19. Both irradiation schedules start on the first day of the first week of treatment. At the end of week 10, repopulation following AHF-48 treatment is more pronounced than repopulation following HF-76.8, in accordance with Figure 1b. All of the above mentioned observations are in agreement with the statistically significant results of the RTOG 83-02 study [18] which refer to the survival for the following radiotherapy fractionations: high dose hyperfractionation (HF, total doses: 76.8 Gy and 81.6 Gy), low dose hyperfractionation (HF, total doses: 64.8 Gy and 72.0 Gy) and accelerated hyperfractionation (AHF, total doses: 48.0 Gy and 54.4 Gy). It should be noted that carmustine (BCNU), which was also administered to all patients included in that study, was assumed not to significantly modify the effectiveness relations among the different radiotherapeutic schemes considered. The

Figure 3. Total number of tumour cells (proliferating, dormant and dead cells) as a function of time for the hyperfractionated (1.2 Gy twice daily, 5 days per week to the dose of 76.8 Gy, ‘‘HF-76.8’’) and accelerated hyperfractionated (1.6 Gy twice daily, 5 days per week to the dose of 48 Gy, ‘‘AHF-48’’) radiotherapy schedules. Irradiation starts on the first day of the first week. HF-76.8 is completed on day 44 after initiation of treatment whereas AHF-48 is completed on day 19. Both irradiation schedules start on the first day of the first week of treatment.

394

reason for this is that the same chemotherapy administration schedule was followed for all patients regardless of the specific radiotherapeutic scheme administered to each individual patient. All patients received BCNU 80 mg m22, intravenously, on days 1, 2 and 3 of the first week of radiotherapy and subsequently on 3 consecutive days every 8 weeks for a period of 1 year, to a maximum dose of 1440 mg m22. Only limited modifications of the schedule based on age and not on the radiotherapy scheme administered were made during the trial. Agreement with the results of the RTOG 83-02 study implies that the simulation model has successfully captured and integrated the critical biological aspects that determine the clinical outcome. Furthermore, it strengthens the proposition that advanced cancer integrative (radio)biology might explain and predict the therapy outcome, despite the complexity of the clinical setting. In the following, a parametric analysis is carried out in order to study the effect of the clonogenic cell density as well as that of the OER and cell cycle duration. The model code has been executed for a simulated period of up to 6 weeks, an interval which normally covers the treatment period of the radiotherapy course and may extend to some days after its completion. The cell cycle duration has been assumed to be 48 h, OER equal to 3.0, CCD equal to 16104 cells mm23 and the LQ parameters have been assumed to match the values of a GBM with known mt p53 gene, unless otherwise stated. The simulation results of Figure 4 demonstrate the ability of the algorithm to effectively simulate the tumour response to a standard irradiation scheme under different values of CCD (16104 cells mm23, 26104 cells mm23, 36104 cells mm23, 46104 cells 23 mm ). Higher values of CCD not only produce a higher number of proliferating tumour cells, but also affect the entire tumour composition. Such behaviour is in accordance with the previously described interdependence among the various metabolic layers of the tumour (dormant, dead, proliferating). Figure 5 provides a twodimensional visualization of the simulated response of a clinical glioblastoma multiforme tumour to the standard fractionation scheme for different CCD values. At the end of day 3, the tumour with 16CCD appears to be more radiosensitive compared with tumours with 26CCD, 36CCD and 46CCD as its population of proliferating cells is lower compared with the others (Figure 5a–d). At the end of day 5, all tumours are strongly affected by radiation treatment, whereas the highest number of proliferating cells is contained in tumour with 46CCD (Figure 5h). At the beginning of the first day of the second week (day 8), newly-produced proliferating cells are present in sufficient numbers to be apparent in Figure 5k,l. The population of the proliferating cells is larger in Figure 5l compared with Figure 5k while the majority of the slice of tumour Figure 5l is painted as ‘‘proliferating’’. For this specific case, the LQ model parameters of the tumour are in accordance with a GBM with known mt p53 gene. The simulation results in Figure 6 demonstrate the ability of the algorithm to effectively simulate the tumour response to a standard irradiation scheme for different OER values (1.0, 2.0, 3.0). Higher OER values describe a more radioresistant tumour. The British Journal of Radiology, May 2006


Figure 4. Simulation predictions of (a) the number of proliferating and (b) the total number of proliferating, dormant and dead mt p53 tumour cells in the case of the standard fractionation scheme (2 Gy once a day, 5 days per week, 60 Gy in total) for different clonogenic cells densities (CCD5104 cells mm23). Irradiation schedule starts on the first day of the first week of treatment. It should be stressed that reproductively dead cells and their offspring that are still cycling are considered proliferating until their ultimate biological death. The periodicity noticed on all graphs reflects the weekly irradiation periodicity. It is noted that no irradiation takes place during the weekend.

Figure 7 depicts the corresponding simulation results of a clinical glioblastoma tumour to the standard fractionation scheme for different Tc values. For the short period simulated (10 days) the tumour with Tc548 h appears to be more difficult to treat than the one with Tc560 h. It can nevertheless be noted that during most of the working days the tumour with Tc548 h shows a better response to irradiation whereas the tumour with Tc560 h shows a slower response to radiation due to its long Tc. Figure 8 is a three-dimensional reconstruction of the simulated response of a GBM with known mt p53 gene for Tc524 h (Figure 8a) and Tc572 h (Figure 8b) at the end of day 8. Figure 8a contains few geometrical cells which have been painted as ‘‘proliferating’’, in contrast to Figure 8b in which all geometrical cells have been painted as ‘‘dead’’. During the interval in which the tumour is not irradiated (days 6 and 7), the proliferating tumour cells with Tc524 h multiply prominently faster than in the case of the tumour with Tc572 h. Concerning the model sensitivity issue, Figures 4, 6 and 7 can also be used in order to quantitatively visualize the effect of perturbing critical model parameters on the simulation predictions. For example, according to Figure 4a, on the 30th day after initiation of the radiotherapy course the number of proliferating tumour cells for the case of CCD equal to 46CCD (CCD5104 cells mm23) is as high as 270% of the number of proliferating cells for the case of 16CCD on the same day. According to Figure 6a, on the 30th day after initiation of the radiotherapy course the number of proliferating tumour cells for the case of OER53.0 is as high as 127% of the number of proliferating cells for the case of OER51.0 on the same day. According to Figure 7a, on the 30th day after initiation of the radiotherapy course the number of proliferating tumour cells for the case of Tc548 h is as high as 339% of the number of proliferating cells for the case of Tc560 h on the same day. Concerning estimation of confidence intervals it is noted that, due to the limited information contained in The British Journal of Radiology, May 2006

the RTOG 83-02 clinical study data and the small number of complete sets of clinical data available by our group, no strict confidence limits can be estimated as yet. It is pointed out that cancer is inherently a partly stochastic phenomenon and, therefore, an accurate estimation of the confidence limits of its predicted in vivo behaviour is not always feasible. It is also noted that, although the simulation model predicts the obvious, it provides a means for quantitatively studying the response of real clinical malignant tumours to radiotherapeutic schemes.

Discussion and conclusion The results presented are in agreement with both qualitative clinical experience and the outcome of the RTOG 83-02 clinical study. Predictably, a whole tumour with shorter Tc tends to repopulate faster and therefore is more difficult to treat. Greater CCD produces a higher number of proliferating tumour cells, and therefore a tumour which is more difficult to treat, and eventually produces a lower tumour control probability (TCP) [59]. Webb and Nahum [56] report that the TCP is a complicated function of the variation in both dose and CCD and TCP will depend on the assumption made about CCD. It is pointed out that reproductively dead cells can cycle only for a very limited number of divisions (one to three) before they and their progeny die biologically. Therefore, re-growth of reproductively dead cells or reproductively ‘‘killed’’ clonogenic cells is unimportant for clinical outcome [60]. Although simulation prediction curves look quite similar (Figures 4, 6 and 7), a closer observation reveals several differences among them. Random numbers have actually been used, but the fact that all biological parameters for all curves in Figure 4 are the same except for the CCD has led to a pretty similar (analogous) behaviour of the composite biological system. Furthermore, a careful and successful application of the 395


Figure 5. Two-dimensional visualization of the simulated response of a radiosensitive clinical glioblastoma multiforme tumour to the standard fractionation scheme, for a range of clonogenic cell density (CCD). The figure shows a centrally located horizontal slice of a tumour with clonogenic cell density equal to 16CCD(5104 cells mm23) (a) 3 fictitious days after the beginning of the radiotherapy course (e) 5 fictitious days after the beginning of the radiotherapy course, (i) 8 fictitious days after the beginning of the radiotherapy course. A centrally located horizontal slice of a tumour with clonogenic cell density equal to 26CCD (b) 3 fictitious days after the beginning of the radiotherapy course (f) 5 fictitious days after the beginning of the radiotherapy course (j) 8 fictitious days after the beginning of the radiotherapy course. A centrally located horizontal slice of a tumour with clonogenic cell density equal to 36CCD, (c) 3 fictitious days after the beginning of the radiotherapy course, (g) 5 fictitious days after the beginning of the radiotherapy course, (k) 8 fictitious days after the beginning of the radiotherapy course. A centrally located horizontal slice of a tumour with clonogenic cell density equal to 46CCD, (d) 3 fictitious days after the beginning of the radiotherapy course, (h) 5 fictitious days after the beginning of the radiotherapy course, (l) 8 fictitious days after the beginning of the radiotherapy course. For this specific case, the LQ model parameters of the tumour are in accordance with a GBM cell line with known mt p53 gene. Irradiation schedule starts on the first day of the first week of treatment. Colour Code: dark grey: proliferating cell layer, light grey: dormant cell layer (G0), white: dead cell layer. The colouring criterion ‘‘99.8%’’ used to visualize the predictions has been defined as follows. ‘‘For a geometrical cell of the discretising mesh, if the percentage of dead cells is lower than 99.8% then {if percentage of proliferating cells . percentage of G0 cells, then paint the geometrical cell dark grey (proliferating cell layer)}, else paint the geometrical cell light grey (G0 cell layer) else paint the geometrical cell white (dead cell layer)’’. 396



Figure 6. Simulation predictions of the number of (a) proliferating and (b) dormant mt p53 tumour cells in the case of standard fractionation for different OER (1.0, 2.0, 3.0) values. It is stressed that reproductively dead cells and their offspring that are still cycling are considered proliferating until their ultimate biological death. Irradiation schedule starts on the first day of the first week of treatment.

Figure 7. Simulation predictions of the number of (a) proliferating and (b) dormant mt p53 tumour cells in the case of standard fractionation for different Tc (cell cycle time) values. It is stressed that reproductively dead cells and their offspring that are still cycling are considered proliferating until their ultimate biological death. Irradiation schedule starts on the first day of the first week of treatment.

Monte Carlo technique has led to a remarkably stable numerical behaviour of the simulation model. It has been experimentally demonstrated that most solid tumours contain resistant hypoxic cells, with estimates of the hypoxic fractions ranging from below 1% to well over 50% of the total viable cell population [33]. Additionally, evidence that hypoxia exists in human tumours to a degree that can influence radiation response comes from those clinical trials in which some form of hypoxia modification has been attempted and found to improve tumour response [61]. Knisely and Rockwell [62] have reported that the resistance of gliomas to treatment with radiation and antineoplastic drugs may result in part from the effects of the extensive and severe hypoxia that is present in such tumours. They have emphasised that the brain tumours contain extensive regions in which the tumour cells are subjected to an The British Journal of Radiology, May 2006

unphysiological degree of hypoxia, this being involved in the evolution of cells in low-grade malignancies to the resistant, aggressive phenotype characteristic of glioblastomas. Furthermore, the results of the parametric OER study agree with data presented by Horsman and Overgaard [33]. It is stressed that although the model confirms the obvious, it allows a quantitative study of the inter-relationship between the competing influences in a complex, dynamic tumour environment. A possible future refinement of the model would include a detailed description of the modulation of cell cycling by oxygen tension by taking into account the latest pertinent experimental observations. It is also noted that, according to our model, intrinsic sensitivity, at least as modified, by OER, is an important determinant of the poor outcome of glioblastoma multiforme irradiation, which appears to be in disagreement with Taghian et al [63]. 397


The CMA tumour shrinkage algorithm applied in this paper is more realistic than the one previously described by Stamatakos et al [17], as CMA is able to substantially conserve the tumour shape (conformal shrinkage) in accordance with Perez and Brady [45]. Additionally, the imaging based neovasculature distribution has been in agreement with Horsman and Overgaard [33]. Finally, the uniform cell distribution algorithm produces a realistic initial cell phase distribution. The predictions of all indicative simulations performed agree at least qualitatively with the clinical experience and with the data presented by Duechting et al [1, 5], Kocher et al [6, 7] and Horsman and Overgaard [33]. It is stressed that the model presented addresses the imageable gross tumour. In order to take into account the brain micrometastases (diffuse invasion), an approach similar to the one developed by Swanson et al [8, 10] and Mandonnet et al [9] or Kansal et al [1, 2] must be considered. It is also noted that the end points of most clinical trials have not been the volumetric and/or metabolic response of tumours, but rather the overall response to treatment such as survival and tumour relapse interval. Consequently, currently available clinical data can be exploited for the validation of the models only in a rough (approximate) way. This implies that a better survival following treatment scheme X compared with survival following treatment scheme Y can be roughly used as an indicator of better tumour control achievable with scheme X. Obviously this can be the case if radiation toxicity is within acceptable limits for both treatment schemes. Concerning fluctuations of radiosensitivity throughout the cell cycle, differing radiosensitivity of cells in the various phases of the cell cycle has been considered and successfully simulated by our group in the case of small in vitro or in vivo preangiogenetic/avascular tumour spheroids [17]. Nevertheless, although easily includable, this variation has not been addressed in the present version of the in vivo gross tumour model in order to keep computer memory and execution time requirements as low as possible. Modelling of the irradiation effects on the surrounding normal tissues in vivo, modelling of the tumour response to chemotherapy in vivo and further enhancement of our models with more genetic data are currently under way. Systematic comparison with clinical data is expected to lead to more clinically adapted parameter values. The clinical validation procedure is in progress and involves comparison of the model predictions with pertinent clinical data before, during and after the radiotherapy course. The easily adjustable, modular simulation model ‘‘follows’’ the clinical practice and activates a selfoptimization procedure. It should also be noted that very rare deviations from the poor standard prognosis of glioblastoma multiforme do exist. Nevertheless, the aim of the simulation model is to predict the most likely time course of the treatment response and, consequently, it is unlikely that an extreme scenario would be predicted. Execution of the computer code on supercomputer systems where a more dense discretizing mesh could be considered might refine the model’s prediction accuracy. The validation process in conjunction with generic parameter estimation techniques (neural networks 398

technique, taboo searching, etc.) can be used to achieve better estimates of the input parameters. Agreement with clinical observations strengthens the applicability of the model to real situations. An integrated and patientindividualized decision support and spatiotemporal treatment planning system is expected to emerge after completion of the necessary clinical adaptation and validation processes. Such a system could also serve as an educational platform for professionals and patients by means of virtual reality demonstrations of the likely natural development and treatment responsiveness of specific cancers so that all groups might positively contribute to the discussion about treatment procedure. Use of genotyping data might enhance the potential of the model, as more accurate estimates of the patient’s individualized linear quadratic radiobiological parameters might be achieved through the use of molecular interaction networks. A more quantitative validation can be achieved using the patient data to be collected and applying multiple parameter adaptation methods such as genetic algorithms or neural networks. The imaging (e.g. MRI, PET, etc.), histopathological and, eventually, genetic data are introduced into the simulation model and a candidate radiotherapeutic or chemotherapeutic scheme is defined. The output of the simulation run, which is the prediction of the tumour and the most affected normal tissue response to the treatment scheme, is then evaluated by the supervising doctor. If a further scheme is to be tested in silico, the simulation run is repeated with the same imaging and radiobiological/pharmacodynamic data as previously. In the end, the modelling platform might serve as a generic ‘‘decision-support system’’. In this way, the medical doctor might make his or her final decision on the selection of the most promising therapeutic scheme by taking into account both the predicted outcomes of all simulated regimens as well as his or her own medical knowledge and expertise. This computational platform does not therefore intend to replace the medical doctor’s input, but to add the possibility of investigating the impact of specific treatment-induced perturbations.

Acknowledgments The project has been supported by the Hellenic Ministry of Education and Religious Affairs under the program ‘‘Irakleitos: Fellowships for Research of the National Technical University of Athens’’. The project has been co-funded by the European Social Fund (75%) and National Resources (25%).

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A quantitative study of IMRT delivery effects in commercial planning systems for the case of oesophagus and prostate tumours J SECO,

PhD,

C H CLARK,

PhD,

P M EVANS,

DPhil

and S WEBB,

PhD, DIC

Joint Department of Physics, Institute of Cancer Research and Royal Marsden NHS Foundation Trust, Downs Road, Sutton, Surrey SM2 5PT, UK ABSTRACT. This study focuses on understanding the impact of intensity-modulated radiotherapy (IMRT) delivery effects when applied to plans generated by commercial treatment-planning systems such as Pinnacle (ADAC Laboratories Inc.) and CadPlan/ Helios (Varian Medical Systems). These commercial planning systems have had several version upgrades (with improvements in the optimization algorithm), but the IMRT delivery effects have not been incorporated into the optimization process. IMRT delivery effects include head-scatter fluence from IMRT fields, transmission through leaves and the effect of the rounded shape of the leaf ends. They are usually accounted for after optimization when leaf sequencing the ‘‘optimal’’ fluence profiles, to derive the delivered fluence profile. The study was divided into two main parts: (a) analysing the dose distribution within the planning-target volume (PTV), produced by each of the commercial treatment-planning systems, after the delivered fluence had been renormalized to deliver the correct dose to the PTV; and (b) studying the impact of the IMRT delivery technique on the surrounding critical organs such as the spinal cord, lungs, rectum, bladder etc. The study was performed for tumours of (i) the oesophagus and (ii) the prostate and pelvic nodes. An oesophagus case was planned with the Pinnacle planning system for IMRT delivery, via multiple-static fields (MSF) and compensators, using the Elekta SL25 with a multileaf collimator (MLC) component. A prostate and pelvic nodes IMRT plan was performed with the Cadplan/Helios system for a dynamic delivery (DMLC) using the Varian 120-leaf Millennium MLC. In these commercial planning systems, since IMRT delivery effects are not included into the optimization process, fluence renormalization is required such that the median delivered PTV dose equals the initial prescribed PTV dose. In preparing the optimum fluence profile for delivery, the PTV dose has been ‘‘smeared’’ by the IMRT delivery techniques. In the case of the oesophagus, the critical organ, spinal cord, received a greater dose than initially planned, due to the delivery effects. The increase in the spinal cord dose is of the order of 2–3 Gy. In the case of the prostate and pelvic nodes, the IMRT delivery effects led to an increase of approximately 2 Gy in the dose delivered to the secondary PTV, the pelvic nodes. In addition to this, the small bowel, rectum and bladder received an increased dose of the order of 2–3 Gy to 50% of their total volume. IMRT delivery techniques strongly influence the delivered dose distributions for the oesophagus and prostate/pelvic nodes tumour sites and these effects are not yet accounted for in the Pinnacle and the CadPlan/Helios planning systems. Currently, they must be taken into account during the optimization stage by altering the dose limits accepted during optimization so that the final (sequenced) dose is within the constraints.

Inverse IMRT treatment-planning optimization is usually performed based on a set of physical and/or biological constraints, independent of the delivery method. Leaf sequencing then creates a set of MLC leaf positions/motions to generate an actual delivered fluence which matches the optimal inverse fluence as closely as possible. When converting the optimal plan, as calculated by the planning system, into a deliverable

The work is supported by Cancer Research UK (under reference grant SP 2313|0201) and the Institute of Cancer Research.


Received 9 February 2005 Revised 7 September 2005 Accepted 7 September 2005 DOI: 10.1259/bjr/91588055 ’ 2006 The British Institute of Radiology

sequence of leaf positions (multiple-static fields, MSF or dynamic delivery, DMLC), the dose delivered to the planning target volume (PTV) and the organs at risk (OAR) will change depending on the MLC leaf characteristics taken into account during the leaf sequencing [1, 2]. In the present study, a quantitative analysis of the impact of IMRT delivery effects for 3 or 5 fields, uniformly distributed around the patient, for multiplestatic fields (MSF), dynamic MLC or compensator delivery techniques has been investigated, for the case of two commercial planning systems and for two anatomical sites. 401

J Seco, C H Clark, P M Evans and S Webb

Method Overview of IMRT delivery techniques: MSF, DMLC or Compensator In converting the optimal fluence plan from the planning system into a deliverable sequence of leaf positions (MSF or DMLC) or compensator thickness, different sequencing algorithms are used. To calculate the leaf sequences for the MSF technique the ADAC (now Philips) Pinnacle planning system uses the Kmeans clustering algorithm [3]. This breaks the optimal fluence into smaller groups or clusters of equal value. For the compensator, the density of the compensation material, width, height, resolution for the compensator and depth of the plane for the compensator shape optimization are defined. The energy fluence is then iteratively attenuated by the corresponding thickness of the modifier at each optimization fluence point. A ‘‘granular compensator’’ was selected with spatial resolution of 0.05 cm and material density of 4.9 g cm23 (approximate density of standard steel granulate compensator material). The source-tocompensator distance was 56.6 cm and the compensator was allowed a maximum thickness of 10 cm. In the case of the DMLC technique, the optimal fluences are converted to the actual fluences (Varian terms) using the leaf motion calculator (LMC), which designs the leaf motion patterns. The LMC takes into account the various MLC parameters such as maximum leaf span (i.e. the physical length of the leaf), leaf speed, transmission, rounded-end effects and minimum leaf gaps. Since the X and Y jaws do not move during beamon and hence cannot follow the trailing leaf, the maximum leaf span will determine how many MLC carriage positions (or static overlapping segments) will be required to deliver the fluence for a given field width (X jaws). The field is split into multiple overlapping fields of the appropriate number of carriage or jaw positions. Although the leaf motions are not fully synchronised, the time of travel across the field is the same for all leaf pairs, which helps to reduce the tongueand-groove effect [4].

Patient setup and treatment objectives Oesophagus tumour site and Pinnacle planning system A patient with oesophageal carcinoma was planned with an Elekta SL25 linac using 3- or 5-field IMRT plans (gantry angles: 0 ˚ (anterior field), 120 ˚, 240 ˚ and 0 ˚, 72 ˚, 144 ˚, 216 ˚ and 288 ˚), respectively, using P3IMRT (Pinnacle, ADAC version 6.0 g). In version 6.0 g, IMRT delivery effects have not been included in the optimization process. The clinical target volumes (CTV), spinal cord and lung parenchyma were outlined on each image. The CTV region included both the oesophagus tumour and adjacent lymph nodes. The PTV region was generated by adding a three-dimensional margin of 15 mm to the CTV to account for movement and target definition uncertainties. The goal of the plan was to deliver 55 Gy 402

to the PTV, while maintaining the spinal cord dose at less than 45 Gy and minimizing the dose to the lungs. The spinal cord dose constraint of 45 Gy is conservative. Martel et al [5] and Emami et al [6] have shown that, in head and neck cancers, the tolerance dose for the spinal cord is around 50 Gy, with a 5% chance of a complication occurring in 5 years. None of the patients treated by Martel et al [5] developed radiation myelitis, with the spinal cord receiving doses up to 50 Gy. In addition to the spinal-cord dose constraint, the volume of lung irradiated to 18 Gy has been used as a planning constraint at the Royal Marsden NHS Foundation Trust. In the present study, no more than 20% of the lung could receive more than 18 Gy (V18, 20%). The prediction of lung complications at the treatment-planning stage is not straightforward. There is no consensus on which dosimetric parameter should be used to reflect the clinical incidence of pneumonitis; however the volume of lung receiving 18 Gy was chosen [7]. Cardiac radiation toxicity for carcinoma of the oesophagus is not a major clinical concern because of the small number of long-term survivors. Therefore, the heart was not included as one of the OARs.

Prostate and pelvic nodes tumour site and Helios/ CadPlan planning system The patient with prostate carcinoma was planned with a 5-beam technique using CadPlan version 6.3.5, with the Helios inverse-planning module. Gantry angles of 180 ˚ (posterior), 270 ˚ (right lateral), 325 ˚ (right anterior oblique), 35 ˚ (left anterior oblique), 100 ˚ (left posterior oblique) were chosen such that the beams were approximately equispaced and were not opposing. These beam angles were found to provide good bowel sparing. The treatment was designed to deliver a dose of 70 Gy to the prostate and 50 Gy to the seminal vesicles and pelvic nodes. The prostate CTV was considered to be the entire visible prostate and was grown to a PTV with a 1 cm margin. However, if the overlap between the PTV and rectum was large, then the posterior margin was reduced to 8 mm. The nodal CTV was expanded to a PTV with a uniform 5 mm margin. The dose constraints used as goals for the prostate and pelvic node treatment were given by Clark et al [8]. For a prostate and pelvic node treatment with five gantry angles, typical beam lengths were 16–18 cm and beam widths were 10–18 cm. Typical monitor units (MUs) were 95 (for a section of the field) and 135 (for a maximum-width single field). The prescribed dose was 2.0 Gy per fraction to the median of the prostate PTV [8].

Clinical impact of IMRT delivery effects An ‘‘optimum’’ IMRT plan was obtained from each of the commercial planning systems used. The optimum IMRT fluence profiles were then leaf-sequenced in order to generate the leaf positions or compensator thicknesses, to allow the delivery of the planned profile. A step-andshoot K-means clustering algorithm from Pinnacle or the dynamic leaf motion calculators from CadPlan/Helios was used to leaf-sequence the IMRT profiles. The British Journal of Radiology, May 2006

Quantitative study of IMRT delivery effects

Each delivered IMRT plan would deliver a slightly different dose to the PTV, from that planned. The delivered IMRT fluence maps were then renormalized by the user such as to attain the prescribed dose to the median of the PTV, i.e. 55 Gy for the oesophagus tumour and 70 Gy for the prostate tumour. In the case of the oesophagus tumour site, the impact of clustering the IMRT fluence profiles was studied, where each fluence map was divided into equal fluence levels, using an error tolerance method equivalent to that described in Ba¨r et al [9]. The Pinnacle sequencing software allows the error tolerance, i.e. the maximum difference between the optimal and sequenced fluence maps, to be specified. The effect of this was studied by specifying 2%, 5% and 10% tolerance levels. In the case of the CadPlan/Helios planning system, only dose–volume values were compared before and after sequencing, for the secondary PTV and critical organs (rectum, bladder, etc.), for three prostate patients with nodal involvement because it was not possible to retrieve dose or fluence information before leaf sequencing, only dose–volume values. The dose–volume values compared correspond to the dose delivered to 90%, 75%, 50%, 25% and 10% of the volume of the bladder, rectum and small bowel and 95%, 75%, 50%, 25% and 5% of the volume of the PTV (prostate), right and left nodes.

Results and discussion Oesophagus tumour site and Pinnacle planning system Clinical impact of IMRT delivery effects on PTV An ‘‘optimum’’ IMRT plan was obtained with the Pinnacle planning system for a prescribed dose to the PTV region of 55 Gy. The IMRT profiles were delivered

using either the MSF or the compensator delivery techniques. In Figure 1, the DVH obtained for the delivered plans in case of 3 and 5 beams are compared with the optimal plan for the PTV (vertical axis starts at 45 Gy). In delivering the optimum fluence profile, the PTV dose has been ‘‘smeared’’ by both the MSF and the compensator techniques. In the case of the 3-beam delivery, the error tolerance of 2% has produced the dose distribution that better approximates the planned optimum (OPT_PTV). This is a consequence of the greater number of fluence levels allowed by the 2% than by the 5% and 10% cases (the compensator was not shown for the 3-beam case since the result was equal to the PTV_2% case). In the case of the 5-beam plan, the delivered plans for the MSF for 5% and 10% error tolerance and the compensator are shown to ‘‘smear’’ the dose distribution planned for the PTV. The 2% is not shown because the system would ‘‘crash’’, destroying any data previously calculated, possibly due to hardware and memory limitations. The compensator has produced the best delivered plan, of all the delivery methods represented for 5 beams. In Figure 2, the dose distribution obtained on a CT slice of the patient for the 3-beam case, obtained for (a) the optimum plan, and MSF delivery with (b) 2% and (c) 5% error tolerance is shown. In Figure 2a, the 98% isodose curve conforms closely to the PTV region represented by the dark red circle. After delivery with the MSF technique, the high 98% isodose covers part of the adjacent lungs. In addition to this, the 80% isodose has significantly changed from the planned distribution (green curve in Figure 2a) to the delivered (green curve in Figure 2b,c). The lungs have received increased dose due to the inclusion of head-scatter and transmission. The spinal cord is also receiving an increased dose as may be observed comparing the 50% isodose (blue) curves in Figure 2.

Figure 1. Dose–volume histograms (DVHs) of the optimum dose (OPT_PTV) or delivered dose to the planning target volume (PTV) by (a) 3 and (b) 5 intensity-modulated radiotherapy (IMRT) plans. The delivery techniques represented are the (i) multiple static fields (MSF) with 2%/compensator (red), 5% (blue) and 10% (green) error tolerance for the 3 beam and 5% (blue) and 10% (green) for the 5 beam and the (ii) compensator (orange) for the 5 beam case. The British Journal of Radiology, May 2006

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(a)

(b)

Impact of IMRT delivery effects on OARs The dose–volume histogram (DVH) for the spinal cord is presented in Figure 3 for both the 3- and 5-beam cases. In the case of the 3 beam, the spinal cord has received significantly more dose than initially planned, due to delivery effects. The increase in the maximum dose delivered to the spinal cord is of the order of 2–3 Gy, in addition to an overall increase of 5–10 Gy in the dose delivered to 50% of the total volume of the spinal cord. In the case of the 5 beams, a similar increase in the spinal cord dose is observed. In Figure 4, the DVH obtained for the left and right lungs are presented. The maximum dose delivered to the lungs increases, respectively, by 1 Gy and 3 Gy for the 3 and 5 beams. For the case of 5 beams, the dose delivered to 50% of the lung volume varies dramatically with the tolerance error used to segment modulated beam profiles. However, for the three beams, there is no significant difference between the tolerance errors in terms of the delivered dose to the lungs by the MSF technique. In the case of the 5 beams, the delivered doses to the lungs by the MSF (5% and 10%) or the compensator are distinctly different. As shown in these results, the IMRT delivery techniques strongly influence the delivered dose distributions to the tumour site (oesophagus) and the surrounding OARs (spinal cord and lungs). The overall increase in 1–3 Gy in the dose delivered to the spinal cord and lungs corresponds to a fractional increase of approximately 1.8–5.4% of the total planned PTV dose of 55 Gy. This increase is a consequence of the additional head scatter and transmission fluence that is given to the patient after delivering the optimum IMRT profile. The additional head scatter and transmission fluence due to these MLC fields is not accounted for in the optimization. This increase in the delivered IMRT fluence due to head scatter and transmission is consistent with that observed by Seco et al [1] when delivering optimum profiles for prostate tumours with the ELEKTA SL25 MLC. In Seco et al [1], the head scatter and transmission contribution for the MSF technique led to an increase of 3% in the delivered IMRT fluence to the prostate tumour and surrounding OARs.

Prostate and pelvic nodes and Helios planning system

(c) Figure 2. A standard oesophageal dose distribution obtained from the (a) optimal and actual fluence delivered with MSF for (b) 2% and (c) 5% error tolerance. The 98%, 80%, 50% and 20% isodose levels are shown, respectively, as red, green, blue and yellow and the planning target volume (PTV) is thick red.

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An ‘‘optimum’’ IMRT plan was obtained with the Helios planning system with a prescribed dose to the PTV region of 70 Gy to the primary prostate tumour and 48 Gy to the nodes. The ‘‘prescribed’’ dose to the nodes was set at 48 Gy (not 50 Gy) at the planning stage, in order to compensate for the increase in the nodal dose after leaf-sequencing the optimum fluence. Therefore, the final dose delivered to the nodes will be approximately equal to 50 Gy. The optimum profiles were then leaf-sequenced in order to generate the leaf positions for the DMLC delivery. The plan was then calculated to allow the evaluation of the delivered profile and dose distribution. Each delivered IMRT plan would yield a different dose to the primary PTV (c.f. that determined by the optimization), with subsequent increase in dose to The British Journal of Radiology, May 2006


Figure 3. Dose–volume histograms (DVHs) obtained for the spinal cord in the case of (a) the 3 and (b) the 5 beam plans. The delivery techniques represented are the (i) multiple static fields (MSF) with 2% (red), 5% (blue) and 10% (green) error tolerance for the 3 beam and 5% (blue) and 10% (green) for the 5 beam and the (ii) compensator (orange) for the 5 beam case.

the secondary PTV (nodes) and OARs (bladder, small bowel and rectum). The mean dose (over the group of three prostate patients) delivered to the volumes of interest ‘‘before’’ and ‘‘after’’ leaf-sequencing are presented (Table 1), for the various dose/volume points studied (the variation in the delivered dose, between the ‘‘after’’ and ‘‘before’’ case, is also given after each volume of interest in brackets). In addition to this, a typical DVH for a single patient is also presented in Figure 5 to illustrate the effect of leaf-sequencing on the final patient DVH. In the case of the OARs there is an increase of up to 3.32 Gy in the delivered dose to 50% of the total volume of the small bowel (c.f. Table 1 and Figure 5). The bladder and rectum are subject to 2.15 Gy and 3.03 Gy more dose, respectively, delivered to 50% of their total volume. The majority of the OARs volumes (bladder, small bowel and rectum) have received significantly more dose delivered to 90% of their volume. In the case of the rectum, this increase was the largest, being in the order of 5 Gy to 90% of its total volume. In the case of the bladder and small bowel, the increases in the dose delivered to 90% of the total volume were 3.78 Gy and 3.06 Gy, respectively. The greater increase in dose delivered to the OARs as compared with the increase to the PTVs is due to the extra transmission and scatter delivered to the OAR during the times when they are shielded by the MLCs. The proportion of time when the OARs are covered by the MLCs is greater than for the targets and therefore the increase in dose to the OARs is also greater. For the secondary PTV (right and left nodes), the leafsequencing of an IMRT ‘‘optimum’’ fluence profile (to be delivered by an MLC) leads to an average increase of, approximately, 2 Gy to 50% of the volume of the nodes, while the dose to the prostate tumour is not significantly altered. In addition, the dose to 90% of the nodal volumes is increased by 3 Gy. The increase is due mainly to head scatter and transmission radiation associated with the MLC delivery equipment and that The British Journal of Radiology, May 2006

is not accounted for by the Helios inverse planning module. In Van Esch et al [10], Clark et al [11] and Hong et al [12] the transmission fluence for MLC leaves measured contributed an additional 1.7%, approximately, to the delivered fluence from Varian Linacs using the dynamic mode of delivery. The increase of approximately 2 Gy in the dose delivered to 50% volume of the pelvic nodes (Table 1) is mainly constituted of approximately 1.7% of transmission fluence (leading to an extra 0.85 Gy in deposited dose) and 2.3% of additional head scatter fluence (leading to an extra 1.15 Gy in deposited dose). The overall increase in the delivered dose is observed for all the volumes of interest: bladder, small bowel, rectum and nodes but not the primary PTV/ prostate, to which the dose is renormalized ‘‘after’’ leaf-sequencing.

Conclusions The impact of IMRT delivery effects on two commercial treatment-planning systems: Pinnacle planning system (ADAC Laboratories Inc.) and CadPlan/Helios (Varian Medical Systems) planning system, was evaluated. The study was performed separately for the oesophagus and the prostate (with pelvic nodes) tumour sites, where multiple static fields and compensator delivery techniques were evaluated for the first tumour site and DMLC was evaluated for the second tumour site. In the case of the oesophagus tumour site and using the Pinnacle planning system, the IMRT delivery effects were shown to produce a ‘‘smearing’’ of the dose to the PTV (oesophagus) and an increase of 2–3 Gy in the dose delivered to the spinal cord. For the prostate (and pelvic nodes) tumour site planned with the CadPlan/Helios (Varian Medical Systems) system, an increase in delivered dose of approximately 3 Gy was observed for 405


Figure 4. Dose–volume histograms (DVHs) obtained for the left and right lungs in the case of the 3 and 5 beams intensitymodulated radiotherapy (IMRT) plan.

Table 1. The mean dose (over a group of three patients with prostate and nodal planning target volumes (PTVs)) delivered to the volumes of interest before and after leaf-sequencing. The difference in delivered dose, to the volumes of interest, between before and after leaf-sequencing is also presented in brackets in Table 1B (A) Mean dose ‘‘before’’ leaf sequencing Dose delivered to volumes of interest (Gy) Vol.

Bladder

Bowel

Rectum

5% 10% 25% 50% 75% 90% 95%

69.73 62.37 48.83 39.70 32.07

41.17 31.50 24.17 13.67 2.67

68.00 51.00 40.83 36.10 31.17

406

PTV prostate

PTV right node

PTV left node

71.00

53.33

65.33

70.00

50.17 48.30 46.43

50.50 48.30 46.50

67.33

43.67

43.83



(B) Mean dose ‘‘after’’ leaf sequencing Dose delivered to volumes of interest (Gy) Vol.

Bladder

Bowel

5% 10% 25% 50% 75% 90% 95%

68.24 61.85 50.98 42.59 35.85

41.80 33.74 27.49 15.24 5.73

(21.49) (20.52) (+2.15) (+2.89) (+3.78)

Rectum

(+0.63) (+2.24) (+3.32) (+1.58) (+3.06)

66.16 52.46 43.86 39.34 36.23

(21.84) (+1.46) (+3.03) (+3.24) (+5.07)

PTV prostate

PTV right node

PTV left node

72.28 (+1.28)

53.62 (+0.29)

63.14 (22.20)

69.77 (20.23)

51.62 (+1.45) 50.34 (+2.04) 48.90 (+2.47)

51.60 (+1.10) 49.98 (+1.68) 48.56 (+2.06)

66.36 (20.98)

46.76 (+3.10)

46.48 (+2.65)

Figure 5. A dose–volume histogram (DVH) of an example patient with prostate carcinoma and nodal involvement (with right (Node_RT) and left (Node_LT) nodes considered separately). The organs at risk considered are the bladder, small bowel and rectum. The dose values are presented for the two cases: (i) before (BEF) shown as points and (ii) after (continuous lines) leaf sequencing with the planning target volume (PTV) dose being normalized to 70 Gy. Arrows indicate the increase in the dose delivered to the volume of interest due to leaf-sequencing effects.

50% of the total volume of the OARs: bladder, rectum and small bowel. In addition to this, an increase of approximately 2 Gy was observed in the dose delivered to the 50% of the total volume of the pelvic nodes, the secondary PTV. The results obtained show that if delivery effects are not accounted for at the planning/ optimization stage, then an increase in delivered dose to the volumes of interest may be expected, after leaf sequencing the optimum IMRT profiles. Currently, they must be taken into account during the optimization stage by altering the dose limits accepted during optimization so that the final (sequenced) dose is within the desired constraints. The British Journal of Radiology, May 2006

References 1. Seco J, Evans P, Webb S. Analysis of the effects of the delivery technique on an IMRT plan: comparison for multiple static field, dynamic and NOMOS MIMiC collimation. Phys Med Biol 2001;46:3073–87. 2. Seco J. Comparison of the efficacy of intensity modulated radiotherapy delivered by competing technologies. PhD thesis, University of London, 2002. 3. Hartigan JA. Clustering algorithms. New York, NY: Wiley, 1975. 4. Essers M, de Langen M, Dirkx MLP, Heijmen BJM. Commissioning of a commercially available system for intensity-modulated radiotherapy dose delivery with

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dynamic multileaf collimation. Radiother Oncol 2001;60:215–24. Martel MK, Eisbruch A, Lawrence TS, et al. Spinal cord dose from standard head and neck irradiation: implications for three-dimensional treatment planning. Radiother Oncol 1997;47:185–9. Emami B, Lyman J, Brown A, et al. Tolerance of normal tissue to therapeutic irradiation. Int J Radiat Oncol Biol Phys 1991;21:109–22. Martel MK, Ten Haken RK, Hazuka MB, et al. Dose-volume histogram and 3-D treatment planning evaluation of patients with pneumonitis. Int J Radiat Oncol Biol Phys 1994;28:575–81. Clark CH, Mubata CD, Meehan CA, et al. IMRT clinical implementation: prostate and pelvic node irradiation using Helios and a 120-leaf MLC. J Appl Clin Med Phys 2002;3:273–84.

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9. Ba¨r W, Alber M, Nusslin F. A variable fluence step clustering and segmentation algorithm for step and shoot IMRT. Phys Med Biol 2001;46:1997–2008. 10. Van Esch A, Bohsung J, Sorvari P, et al. Acceptance tests and quality control (QC) procedures for the clinical implementation of intensity modulated radiotherapy (IMRT) using inverse planning and the sliding window technique: experience from five radiotherapy departments. Radiother Oncol 2002;65:53–70. 11. Clark CH, Bidmead AM, Mubata CD, et al. Intensitymodulated radiotherapy improves target coverage, spinal cord sparing and allows dose escalation in patients with locally advanced cancer of the larynx. Radiother Oncol 2004;79:189–98. 12. Hong L, Alektiar K, Chui C, et al. IMRT of large fields: whole-abdomen irradiation. Int J Radiat Oncol Biol Phys 2002;54:278–89.



Effect of ticlopidine in the prevention of radiation enteropathy 1

S AKYUREK, MD, 2L ATAHAN, MD, 2M CENGIZ, MD, 3C SOKMENSUER, MD, 4I HABERAL, MD, 2F YILDIZ, 1 MD and C ONAL, MD 1

Department of Radiation Oncology, Ankara University School of Medicine, Ankara, Departments of 2Radiation Oncology and 3Pathology, Hacettepe University School of Medicine, Ankara and 4 ENT and Head and Neck Surgery Department, Health Ministry Ankara Education and Training Hospital, Ankara, Turkey

ABSTRACT. Impairment of vascular function is considered to play an important role in chronic radiation enteropathy. In this experimental study, the role of ticlopidine, an inhibitor of ADP-induced platelet aggregation, was investigated in radiation enteropathy. 80 male Wistar albino rats, each weighing 170–200 g, were divided into four groups: (a) radiation alone (n520); (b) radiotherapy plus ticlopidine (n520); (c) ticlopidine control (n520) and (d) control (n520). Both radiation groups received 19 Gy radiation to the exteriorized intestinal segments in a single fraction. Ticlopidine or vehicle was administered 12 h after radiotherapy and continued for 1 month. Rats from every group were euthanized randomly at intervals of 6 weeks from 2 weeks to 26 weeks. Histopathological radiation injury was assessed using radiation injury scoring (RIS). Radiation with ticlopidine or radiation alone groups showed significant RIS deterioration compared with controls in all time points studied. Comparison of median RIS of radiotherapy and radiotherapy+ticlopidine groups at the 2nd, 14th and 26th weeks yielded statistically significant RIS in favour of radiotherapy+ticlopidine group (p50.05). However, these differences were less pronounced at the 8th and 20th week (p50.07). Both radiation groups had poor weight gain when compared with control and ticlopidine groups. The weight gain in radiotherapy+ticlopidine group was significantly superior to only radiation group between 10th and 20th weeks (p50.05). This study showed that inhibition of platelet aggregation with ticlopidine might be useful in radiation enteropathy. However, the precise role of antiaggregant therapies on radiation enteropathy should be comprehensively studied before clinical consideration.

Normal tissue damage is the main dose-limiting factor in clinical radiotherapy. The small intestine is one of the most radiosensitive organs and unavoidably included in most abdominal or pelvic radiotherapy fields [1]. Despite efforts to keep doses below tolerance doses, severe chronic radiation enteropathy (CRE) occurs in about 10–15% of patients and has an adverse effect on the quality of life of long term survivors [2, 3]. To date there has been neither an effective method of preventing radiation-induced intestinal injury nor the treatment once it has occurred. Intestinal radiation injury (radiation enteropathy) is classified as either acute or chronic. Acute radiation enteropathy is usually a self-limiting condition, which is the result of the cell kill of the rapidly renewing epithelial cells of the mucosa leading to a temporary breakdown of the mucosal barrier and inflammation. CRE, on the other hand, is associated with high rates of morbidity and mortality [3, 4], and the underlying molecular mechanisms have not been clearly identified yet. The characteristic histopathological findings after radiotherapy in acute phase are reported to take the form The British Journal of Radiology, May 2006

Received 17 March 2005 Revised 22 June 2005 Accepted 30 August 2005 DOI: 10.1259/bjr/16265085 ’ 2006 The British Institute of Radiology

of endothelial cell swelling, increased permeability and interstitial fibrin deposition, and development of platelet-fibrin thrombi in submucosal capillaries and microvasculature has been accused of being the main target for radiation damage. As the injury progress to chronic radiation enteropathy, the prominent histopathological changes appear to be intestinal wall fibrosis and vascular sclerosis [5–8]. It has previously been reported that CRE is associated with significant upregulation of the transforming growth factor b (TGF-b) and fibrogenic cytokine, and downregulation of endothelial cell surface protein, thrombomodulin (TM) [9–12]. Thrombomodulin plays a key role in protein C anticoagulant pathway. Reduced endothelial TM levels and increased prothrombotic properties of endothelium seem to contribute to hypercoagulation with increased fibrin formation, tissue factor activity and platelet aggregation [8, 13–15]. Platelets are the most important sources of fibrogenic factors and cytokines, such as TGF-b which is critically involved in radiation enteropathy [16, 17]. Adenosine 59-dipohosphate (ADP), on the other hand, is a potent and specific mediator of platelet aggregation 409

S Akyurek, L Atahan, M Cengiz et al

and activation that is released by damaged endothelial cells, red blood cells and activated platelets [18]. There are some encouraging data in the literature which show that anticoagulation therapy, such as heparin, warfarin, or acetyl salicylic acid (ASA) may be beneficial in the treatment of radiation injury [19–21]. Ticlopidine is a powerful platelet aggregation inhibitor recently used with encouraging success in human clinical trials. It has been reported to decrease recurrence and severity of cerebrovascular and cardiovascular accidents, peripheral arterial disease, and to decrease the occlusion rates of coronary saphenos vein bypass grafts [22]. Ticlopidine achieves antiplatelet efficacy by blocking activation of platelets by ADP. It selectively and irreversibly inhibits the binding of ADP to its receptor on platelets, thereby effecting ADPdependent activation of the glycoprotein IIb/IIIa complex, the major receptor for fibrinogen present on the platelet surface [23]. The study reported here was designed to determine whether ticlopidine has ability to prevent or retard the development of radiation induced enteropathy if it is used prophylactically.

Methods and materials Animals 80 male Wistar rats, weighing 170–200 g (aged 8–12 weeks) at the time of irradiation, were used in this experimental study. They were housed with five to a cage and had free access to water and commercial food pellets. All animals were kept under the same experimental conditions.

Irradiation Before irradiation, all rats including the controls were anaesthetized with 80 mg kg21 ketamin hydrochloride (Ketalar; Parke-Davis, Morris Plains, NJ) and 10 mg kg21 xylazin (Rompun; Milles Lab, Shawnee, KS). A surgical procedure which has been described by Hauer-Jensen et al [24] was performed to all rats, including shamirradiated animals, before radiotherapy. Briefly, a 1.5– 2 cm midline incision was made and 10 cm mid-small intestinal segment, 30–40 cm proximal to the caecum, was exteriorized and marked with silk ligatures in the adjacent mesentery. Rats were placed in a specially designed box, with the exteriorized intestinal segment lying on gauze compress soaked in 0.9% saline, and care was taken to avoid tension of the mesentery during irradiation. Only the exteriorized intestinal segment was irradiated and the rest of animal was protected by a 5 mm lead shield. The intestine was again placed into the abdomen at the end of the procedure. The intestinal segment was irradiated with 6 MeV electron beams using a Philips SL-25 linear accelerator (Philips, Best, The Netherlands). Before irradiation, animals were randomly assigned to one of the four groups. A single fraction of 19 Gy was delivered through a 6 cm66 cm anterior portal in group 1 (RT, n520). Same dose was given in group 2 (RT+T, n520) with 100 mg kg21 daily oral ticlopidine HCl (Ticlid-Sanofi) for 1 month starting from 12 h after irradiation. Similar 410

procedures except irradiation were performed in 2 agematched control groups, a ticlopidine control group (T, n520) and no-drug control group (C, n520). All animals were followed up weekly for weight loss and for complications. Experiments were performed in accordance with the national regulations for animal experimentation, and experimental protocols were approved by the local animal ethics committee before the start of studies.

Morphological assessment Four animals from each group were randomly selected and euthanized by bleeding under ether anaesthesia at intervals of 6 weeks from 2 (early inflammatory phase) to 26 weeks (chronic, fibrotic phase). Specimens suturated before irradiation procedure from irradiated and shamirradiated animals were excised and fixed in 4% formaldehyde solution, dehydrated and embedded in paraffin. After that cross sections and longitudinal sections of the intestine were performed then stained with haematoxylin-eosin for histopathological evaluation. Each specimen was examined with light microscopy and histopathological changes were scored by a single pathologist (CS) blinded to the study. Scoring (RIS) was done by recording the macroscopical and 7 histopathological changes: mucosal ulcerations (0–2), epithelial atypia (0–3), thickening of subserosa (0–3), vascular sclerosis (0–3), intestinal wall fibrosis (0–3), ileitis cystica profunda (0–3), and lymphatic congestion (0–1) within the irradiated segment. The RIS has been shown to be a reliable indicator of the severity of intestinal radiation injury [3, 24–28]. Total RIS was calculated by adding the scores of the individual parameters.

Statistical methods The Mann-Whitney U-test was used to compare median RIS values among the treatment groups. Student t-test was used to compare mean weights¡ standard errors. A significance level of p,0.05 was selected as the minimal level for significance.

Results All animals were followed-up to 6 months after treatment. All but eight rats survived during this period. Five deaths (14th , 28th, 49th, 126th, 161st days) from RT group and three deaths (14th, 42nd, 147th days ) from RT+T group were found dead during the follow-up time. Autopsies revealed intestinal perforation in three rats, obstruction in two rats and haemoperitoneum for the rest. There is no difference regarding to haemoperitoneum between RT and RT+T groups.

Weight changes Time-related weight changes for each group are shown in Figure 1. Both radiation groups had significantly poor The British Journal of Radiology, May 2006

The prevention of radiation enteropathy with ticlopidine

Figure 1. Weight changes as a function of time after irradiation.

weight gain when compared with C and T groups. There was no statistically significant difference between irradiated groups during the first 20 weeks, but a significant difference (p,0.05) in favour of RT+T group was observed between 10th and 20th weeks.

Morphological changes There were no significant changes in control groups except minor histopathological alterations, which were considered to be due to ageing. The median RIS value, as a function of time, is given in Figure 2 and Table 1. At 2 weeks after irradiation, the RT group showed deterioration in the crypt morphology, increase in epithelial proliferation, superficial and deep ulceration, and serosal thickening. Total RIS score was assessed as 11 in this group. The RT+T group on the other hand only showed minimal crypt epithelial changes and subserosal thickening due to increased amount of fibrin deposits and oedema. Median RIS score was 6 and the difference was statistically significant between these 2 groups (p,0.05). At 8 weeks after irradiation, macroscopic thickening of the bowel wall was apparent in the RT group and

Figure 2. Median radiation injury scores as a function of time after irradiation.


prominent vascular hyalinization in addition to histopathological changes in the crypt morphology was recorded. Median RIS score was 11. For the RT+T group, macroscopic thickening of the bowel wall was similar to the RT only group, whereas light microscopic evaluation revealed only superficial mucosal ulcerations and serosal thickening. Median RIS was 7, which was not statistically significant, although there was a trend in favour of RT+T (p50.07). At the 14th week after irradiation, prominent bowel wall thickening, stenosis and adhesions were recorded macroscopically for RT only group. Severe mucosal ulcerations, vascular sclerosis, hyalinization and apparent serosal thickening were observed by light microscopy and the median RIS was 14. RT+T group, on the other hand, showed only wall thickening with superficial ulceration, minimal serosal thickening and increase in vascular hyalinization at this week. Median RIS was 6, which was significantly better than the RT only group (p,0.05). At the 20th week, histopathological appearance in the RT group was not remarkably different from the 14th week of evaluation. The RT+T group on the other hand showed some improvement on the mucosal crypt morphology, although vascular hyalinization and sclerosis continued. There was a significant difference both macroscopically and microscopically between the two groups at the 26th week of evaluation. Apparent wall thickening, adhesions and stenosis with similar light microscopic findings to the previous weeks were recorded for the RT group. However, although minimal, some improvement was shown compared with the 20th week for the RT+T group (Figures 3–6).

Discussion The present study demonstrates that ticlopidine, an inhibitor of ADP-induced platelet aggregation, has an ability to reduce radiation induced damage in intestine both in early and late phases in the animal model. It is important to remember that the animal model does not exactly reflect its human counterpart. Radiation damage in different animal models can be quite different in terms of severity and the presentation of the damage. However, basic physiopathological chain of events and the final appearance looks similar [29, 30]. The histopathological appearance of chronic enteropathy in patients is described as atropic mucosa, thick and fibrotic submucosa and subserosa [31], which were similar to animal model findings. We therefore assume that animal model might be taken as representative of human gastrointestinal toxicity. Radiation induced morphological changes in the intestine were reported to begin just after irradiation. Progressive crypt shrinkage during the first 24–36 h after irradiation was shown by Potten et al [32]. Recently, radiation-induced apoptosis of endothelial cells were reported to play an essential role in the pathogenesis of acute gastrointestinal syndrome [33, 34] and it was prevented when endothelial cell apoptosis was inhibited [34]. Although vascular damage was not prominent at 2 weeks after irradiation in our study, vascular hyalinization and sclerosis was apparent at 8 weeks in the RT 411

S Akyurek, L Atahan, M Cengiz et al Table 1. Median RIS values stratified by groups

RT RT+T T C

2nd week

8th week

14th week

20th week

26th week

10 6 0 0

11 (10–12) 8 (7–9) 0 (0–0) 0 (0–0)

13 (12–14) 8 (4–10) 1 (0–1) 0 (0–0)

13 (12–14) 9 (8–9) 0 (0–0) 0 (0–0)

12 (11–13) 6 (6–6) 1 (0–2) 0 (0–0)

(9–12) (6–7) (0–0) (0–0)

RT, radiotherapy alone group; RT+T, radiotherapy plus ticlopidine group; T, ticlopidine control group; C, control group

only group. Although endothelial cell injury was not assessed separately, we think that vascular damage at 8 weeks supports the endothelial cell damage in the early phase after irradiation. Our findings are consistent with those of Jensen et al, who reported that vascular changes were almost absent 2 weeks after irradiation, increased during the following 8 weeks and stabilized 8 weeks after irradiation [35].

Figure 3. Radiotherapy (RT) alone group. Specimen exhibits extensive mucosal ulceration and vascular sclerosis, 26 weeks after irradiation (vs, vascular sclerosis; ul, ulcer; s, serosa).

Figure 5. Macroscopic appearance of the radiotherapy (RT) alone group 26 weeks after irradiation demonstrates apparent bowel wall thickening, adhesion and stenosis. Arrows show suturated specimen (vs, vascular sclerosis; ul, ulcer; s, serosa).

Figure 6. Macroscopic appearance of the radiotherapy+tiFigure 4. Radiotherapy+ticlopidine (RT+T) group. Specimen exhibits mild thickening of muscularis mucosa, 26 weeks after irradiation (vs, vascular sclerosis; m, muscularis mucosa).

412

clopidine (RT+T) group 26 weeks after irradiation demonstrates mild bowel wall thickening. Arrows show suturated specimen (vs, vascular sclerosis; m, muscularis mucosa).


The prevention of radiation enteropathy with ticlopidine

Radiation induced damage to the vasculature has long been considered an important pathophysiological mechanism of CRE [36]. The predominant damage after irradiation has been reported to be in the microvasculature causing endothelial cell swelling, detachment of the endothelium from underlying matrix, increased capillary permeability, progressive loss of endothelial cells, and platelet adhesion and aggregation [37, 38]. It has also been shown that radiation enhances von Willebrand factor release, down-regulation of endothelial TM and prostacyclin, which promote prothrombotic properties. Some encouraging results were reported, that administration of anticoagulation therapy such as heparin, warfarin, or ASA may be effective against radiation injury in some organs [19–21]. Glantz et al studied heparin and warfarin in radiation induced nervous system injury [20] and reported that anticoagulation therapy may reverse small vessel endothelial injury, which is the fundamental lesion of radiation necrosis, and produce clinical improvement in some patients. Furthermore, Ludgate et al demonstrated antiplatelet properties of ASA in radiation enteropathy, whereas 5-amino salicylic acid, a potent anti-inflammatory agent with minimal effect on platelet aggregation, was found to be ineffective and possibly even harmful in radiation enteropathy [39, 40]. These observations support the hypothesis that decreasing the platelet adhesiveness by antiaggregant therapy may produce the decrease of microvascular thrombosis and improve the microcirculation, finally decreasing late damage. Antiplatelet agents in the prevention of CRE, on the other hand, have been studied in the literature with promising results [22, 23]. The role of these agents in prevention are assumed to be directed on their activities against thrombocytes, since a large number of studies have shown that these drugs selectively inhibit ADPinduced platelet activation by antagonizing the binding of ADP to platelet receptors [23]. ADP elicits various effects on platelets after binding to the specific receptor for ADP, such as inhibition of adenyl cyclase, mobilization of calcium from internal stores and expression of fibrinogen receptor [23, 41, 42]. In a study by Wang et al, short-term clopidogrel (a recently developed analogue of ticlopidine), administration starting 2 days before until 10 days after irradiation, produced protection against early and, to a lesser extent, delayed radiation enteropathy [43]. In this particular report, the authors suggested that temporary inhibition of platelet aggregation is not sufficient to permanently interrupt the fibrogenic cycle responsible for intestinal fibrosis, and that prolonged treatment or combination therapies may be necessary. Contemporary to this study we administered ticlopidine for 1 month. Median RIS values in our study showed statistically significant differences between RT and RT+T groups in the 2nd, 14th and 26th weeks of evaluation, which leads us to think that the drug not only attenuates the early, but also the late effects of irradiation. Platelets are the first cellular elements that involved and initiate the haemostatic and inflammatory responses and release a number of proinflammatory and fibrinogenic mediators, and the main effect of antiaggregants in prevention are assumed to be directed by their antiplatelet activity. In recent years, however, it was shown that ticlopidine and The British Journal of Radiology, May 2006

clopidogrel increase endothelial nitric oxide and prostacyclin production [42, 44–46] and modulate the contractile response of vascular smooth muscles [46]. So it may be rational to think that the main effect is not only due to antiaggregant property but also positive modulation of endothelial cell function. We used ticlopidine as a daily dose of 100 mg kg21 based on the preliminary experiments, in which Yamamoto et al studied antithrombotic effect of ticlopidine on He-Ne laser induced thrombus formation in rat mesenteric microvessels [47]. In that experimental study, Yamamoto et al reported that at a dose of 100 mg kg21, ticlopidine inhibited thrombus formation both in arterioles and venules and appeared to be more potent than ASA. Our results confirm the efficacy of this dose in terms of preventing radiation induced enteropathy.

Conclusion This study supports that ticlopidine can ameliorate early and chronic radiation enteropathy. The effect of antiaggregant therapies on radiation enteropathy should be comprehensively studied in phase II and III studies before clinical considerations.

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Radioprotective effects of hesperidin against gamma irradiation in mouse bone marrow cells 1

S J HOSSEINIMEHR,

PhD

and 2A NEMATI,

MSc

1

Department of Medicinal Chemistry, Faculty of Pharmacy, Mazandarn University of Medical Sciences, Sari and 2Novin Medical Radiation Institute, Tehran, Iran ABSTRACT. The radioprotective effects of hesperidin (HES), a flavonone glucoside, were investigated by using the micronucleus test for anticlastogenic and cell proliferation activity. A single intraperitoneal (ip) administration of hesperidin at doses of 10 mg kg21, 20 mg kg21, 40 mg kg21, 80 mg kg21 and 160 mg kg21 45 min prior to gamma irradiation (2 Gy) reduced the frequencies of micronuleated polychromatic erythrocytes (MnPCEs). All five doses of HES significantly reduced the frequencies of MnPCEs and increased PCE/PCE+NCE ratio in mice bone marrow compared with nondrug-treated irradiated control (p,0.0001). There was a drug dose–response effect of HES in reducing MnPCE and increasing the PCE/PCE+NCE ratio in bone marrow cells. The maximum reduction in MnPCE was observed in mice treated with HES at a dose of 80 mg kg21. The total MnPCE values were 2.85 fold less in the 80 mg kg21 HES group after being exposed to 2 Gy of c-rays than those in the respective irradiated control. Our study demonstrates that hesperidin has powerful protective effects on the radiation-induced DNA damage and on the decline in cell proliferation in mouse bone marrow.

Ionizing radiation generates reactive oxygen species in the cells. These free radicals can induce damage to critical macromolecules such as DNA. The cellular DNA damage leads to mutation and cancer [1]. High levels of gamma irradiation can induce mortality in mammals. With respect to radiation damage to humans, it is important to protect biological systems from radiationinduced genotoxicity or lethality. The main radioprotective class is thiol synthetic compounds such as amifostine. Amifostine is a powerful radioprotective agent compared with other agents, but this drug is limited in the use in clinical practice due to side effects and toxicity [2–4]. The search for less-toxic radiation protectors has spurred interest in the development of natural products. Recently, we reported that citrus extract protects mouse bone marrow cells against gamma irradiation. The citrus extracts contained high amounts of flavonoids [5]. Flavonoids are a family of polyphenolic compounds found in fruits and vegetables. Flavonoids have wide biological properties including antibacterial, antiviral, anticancer, immunostimulant and antioxidant effects [6]. Hesperidin (HES) is a flavonone glycoside, belonging to the flavonoid family. This natural product is found in citrus species. HES was reported to have many biological effects including anti-inflammatory, antimicrobial, anticarcinogenic and antioxidant effects, and decreasing capillary fragility [7]. HES, in combination with a flavon called diosmin is used as DaflonH (Servier, France) to treat chronic venous insufficiency in Europe [8]. Other biological effects of HES are unknown. In continuation of this line of investigation, the in vivo radioprotective activity of HES was investigated by using gamma rays as an oxidative DNA damaging agent, and evaluating any reduction in the The British Journal of Radiology, May 2006

Received 22 April 2005 Revised 20 September 2005 Accepted 12 October 2005 DOI: 10.1259/bjr/40692384 ’ 2006 The British Institute of Radiology

frequency of micronuclated polychromatic erythrocytes (MnPCEs) in mouse bone marrow exposed to gamma rays.

Materials and methods Animals Male NMRI mice weighing 25¡3 g were purchased from the Razi Institute (Tehran, Iran). Mice were housed in accordance to the principles outlined in ‘‘The Guide for The Care and Use of Laboratory Animals’’ prepared by Tehran University of Medical Sciences in the university animal house, and given standard mouse pellets and water ad libitum. All animals were kept under controlled lighting conditions (light:dark, 12:12 h) and temperature (22¡1 ˚C).

Chemicals and treatment HES was from Aldrich (USA). HES was dissolved in phosphate buffered saline (pH 7.6). Mice were injected intraperitoneally for all experiments. For selection of the optimum dose of HES for radioprotection, five doses (10 mg kg21, 20 mg kg21, 40 mg kg21, 80 mg kg21, 160 mg kg21) of HES were administrated to the experimental animals 45 min before gamma irradiation. The control animals received the same volume of normal saline or buffer phosphate. Five mice were used for each treatment group. 415

S J Hosseinimehr and A Nemati

Irradiation

Statistical analysis

Whole-body irradiation was performed with a cobalt60 c-radiation source (Teratron 780, Canada). Mice were placed in ventilated PlexiglasH cages and irradiated in groups of five mice, simultaneously. The source-to-skin distance was 80 cm with a dose rate of 1.03 Gy min21 at room temperature (23¡2 ˚C). The mice were irradiated with a total dose of 2 Gy c-rays.

The data are presented as mean¡standard deviation (SD). One-way analysis of variance (ANOVA) analysis and Tukey’s HSD test were used for multiple comparisons of data.

Micronucleus assay The mouse bone marrow micronucleus test was carried out according to the method described by Schmid for evaluation of the chromosomal damage in experimental animals [5, 9]. The animals were euthanized by cervical dislocation 24 h after irradiation. The bone marrow from both femurs was flushed in the form of a fine suspension into a centrifuge tube containing fetal calf serum (FCS). The cells were dispersed by gentle pipetting and collected by centrifuge at 2000 rpm for 5 min at 4 ˚C. The cell pellet was resuspended in a drop of FCS and bone marrow smears were prepared. The slides were coded to avoid observer bias. After 24 h air-drying, the smears were stained with May-Grunwald/Giemsa, as described by Schmid. With this method, polychromatic erythrocytes (PCEs) stain reddish-blue and normochromatic erythrocytes (NCEs) stain orange, while nuclear material is dark purple. For each experimental point, five mice were used and a total of 5000 PCEs were scored for each experimental point to determine the percentage of micronucleated polychromatic erythrocytes (MnPCEs), micronucleated normochromatic erythrocytes (MnNCEs) and ratio of PCE to (PCE + NCE). The ratio of PCE to (PCE + NCE) was determined for each experimental group to assess radiation effects with or without HES on bone marrow proliferation [10].

Results The effect of gamma irradiation with or without HES on the induction of MnPCEs and the PCE/PCE+NCE ratio in bone marrow cells, 24 h after c-irradiation, is shown in Table 1. The frequency of micronuclei was increased in all groups of mice irradiated with 2 Gy cirradiation compared with the control treated with normal saline or phosphate buffer (p,0.0001). The frequencies of MnPCE found in the HES treated groups were significantly lower than that of the group treated with radiation alone. The total MnPCE values were 1.15, 1.36, 1.66, 2.85 and 1.8 fold less in the 10 mg kg21, 20 mg kg21, 40 mg kg21, 80 mg kg21 and 160 mg kg21 HES group after being exposed to 2 Gy of c-rays, respectively, than those in the respective irradiated control (Figure 1). All five doses were effective in significantly reducing (p,0.0001) the frequency of MnPCE induced by 2 Gy irradiation and there was a significant difference between the effects of various doses of HES. There was a drug dose–response effect of HES in the reduction of MnPCE in bone marrow cells. The maximum reduced MnPCE was observed in mice treated with HES at a dose of 80 mg kg21 (Figure 1). The frequency of MnPCE in the latter group was 5.01¡0.32%, much lower than in the group receiving radiation alone (14.29¡0.5). With a further increase in the HES dose to 160 mg kg21, there was a reduced effect of HES on the frequency of MnPCE induced by c-irradiation. The ratio of PCE/PCE+NCE reduced significantly after exposure to 2 Gy of c-irradiation (p,0.0001). Determination of

Table 1. Effects of hesperidin (HES) on the formation of radiation-induced micronulei polychromatic erythrocytes (PCEs) and normochromatic erythrocytes (NCEs) and the ratio of PCE/PCE+NCE in mice bone marrow exposed to 2 Gy c-irradiation Treatment

MnPCE/PCE (%)

MnNCE/NCE (%)

PCE/PCE+NCE (%)

Control Normal saline Phosphate buffer Irradiation 10 mg kg21 HES +irradiation 20 mg kg21 HES +irradiation 40 mg kg21 HES +irradiation 80 mg kg21 HES +irradiation 160 mg kg21 HES +irradiation 40 mg kg21 HES 80 mg kg21 HES 160 mg kg21 HES

0.236¡0.108 0.482¡0.239 0.674¡0.295 14.292¡0.503 12.422¡0.674

0.700¡0.300 0.718¡0.285 1.43¡0.319 3.200¡0.661 2.861¡0.440

62.872¡2.139 62.716¡3.160 58.744¡2.643 39.010¡1.537 42.194¡1.063

10.943¡0.496

3.644¡0.760

44.800¡1.659

8.604¡0.768

3.248¡1.096

50.046¡1.302

5.012¡0.317

2.71¡0.549

54.658¡1.633

7.957¡0.537

3.187¡0.489

49.974¡1.674

0.755¡0.208 0.786¡0.266 2.405¡0.492

0.654¡0.234 0.838¡0.229 1.575¡0.338

57.338¡1.878 55.796¡1.152 49.910¡1.668

MnPCE, micronucleated polychromatic erythrocyte; MnNCE, micronucleated normochromatic erythrocyte.

416


Radioprotective effects of hesperidin

Figure 1. Effect of various doses of hesperidin (HES) on the frequency of micronucleated polychromatic erythrocytes (MnPCEs) in the bone marrow of mice exposed to c-radiation (R) at a dose of 2 Gy.

ratio of PCE/PCE+NCE in the gamma irradiated mice showed a pronounced cytotoxic effect of radiation on bone marrow proliferation. Treatment of mice with HES arrested the radiation-induced decline in the PCE/ PCE+NCE ratio (Table 1), and this increase in the PCE/ PCE+NCE ratio in the HES+irradiated group (at doses 20 mg kg21, 40 mg kg21, 80 mg kg21, 160 mg kg21) was higher than that of the irradiated-alone group (p,0.001). The highest PCE/PCE+NCE ratio was observed in HES treated mice with 80 mg kg21 before c-irradiation (Figure 2). There was no significant difference between phosphate buffer-control and HES treated mice at this dose before c-rays. In this study, HES did not indicate any genotoxic and toxic effects at 40 mg kg21 and 80 mg kg21, but genotoxicity was observed at 160 mg kg21 compared with phosphate buffer (p,0.001).

Discussion We previously reported that citrus extracts could protect mice bone marrow cells against gamma irradiation and cyclophosphamide, when injected prior to exposure and cyclophosphamide treatment. These citrus extracts contain high levels of flavonoids [5, 11]. The results of this study demonstrated the protective effects of HES, a flavonone, against genotoxicity and toxicity induced by

c-irradiation in mouse bone marrow cells. However, synthetic compounds, mainly thiol compounds, have good radioprotective effects, but they are limited in their use by side effects. Natural compounds, including flavonoids, may play a role in scavenging free radicals, such as hydroxyl radicals generated by c-rays in cells. Ionizing radiation generates free radical damage in DNA and induces genotoxic effects and death in the cells [1, 12]. There is a possibility that pre-treatment with flavonoids could induce protection against oxidative stress. Orientin and Vicenin, two flavonoids, protect mice against chromosomal aberration induced by cirradiation when administrated before 2 Gy c-rays [13]. HES is the predominant flavonoid in lemons and oranges. The peel and membranous parts of these fruits have the highest HES concentrations. Sweet or navel oranges (Citrus sinensis) and lemons (Citrus limon) are the richest dietary sources of HES [14]. In this study, HES had a dose-dependent protective effect on reducing MnPCEs induced by gamma irradiation. The greatest protective effect was observed at a dose of 80 mg kg21. HES did not show any genotoxic or toxic effects at doses up to 80 mg kg21 in mice bone marrow cells. Treating mice with HES at dose 80 mg kg21 before exposure to 2 Gy radiation reduced the frequency of MnPCEs almost 2.85 fold. The percentage PCE/PCE+NCE ratio declined in irradiated mice, since this ratio gives a direct index of cell division. HES protected mice against radiationinduced decline in cell proliferation, as evidenced by the increased PCE/PCE+NCE ratio. The molecular mechanism of the radioprotective effects of HES is not clear. It has been reported that flavonoids have antioxidant and chelating properties. These polyphenols are excellent scavengers of free radicals due to the high reactivity of their hydroxyl substituents [12]. Free-radical scavenging is apparently responsible for the inhibitory effect of flavonoids such as rutin, morin, quercetin and genestin on the clastogenic activity induced by cirradiation in mice [15]. Among naturally occurring flavonoids, HES has been pharmacologically evaluated as a potential anticarcinogenic agent because of its antioxidant activity [8, 16]. Other biological effects include immune-modulation, treatment of venous insufficiency and scavenging of peroxynitrite as a reactive oxidant [7, 17]. HES has also protected against photoinduced breakage of DNA [18]. It is possible that HES protects bone marrow cells with its antioxidant activity. However, the antioxidant capacity of HES is not as high as that of other flavonoids such as quercetin and myricetin [19, 20]. Other mechanisms probably contribute to its radioprotective effects. Thus, further experiments are needed to explain the molecular mechanism of HESs protective effects. In conclusion, our results demonstrate that HES gives significant protection to mice bone marrow against the clastogenic effects of gamma irradiation.

Acknowledgments Figure 2. Effects of various doses of hesperidin (HES) on the radiation induced polychromatic erythrocyte (PCEs)/PCE+ normochromatic erythrocytes (NCEs) ratio in the bone marrow of mice exposed to c-radiation (R) at dose 2 Gy.


We would like to thank Mr Mahmoudzadeh, Dr Zahmatkesh, Dr Akhlaghpour and Dr Abassi from the Novin Medical Radiation Institute for their assistance. 417

S J Hosseinimehr and A Nemati

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Characteristics and predictive factors of early-onset diarrhoea during pelvic irradiation 1,2

E Y HUANG,

MD,

1

C J WANG,

MD,

1

H C HSU,

MD

and 1L M SUN,

MD, MS

1

Department of Radiation Oncology, Kaohsiung Chang Gung Memorial Hospital, 123 Ta-Pei Road, Niao-Sung Hsiang, Kaohsiung Hsien, Taiwan and 2School of Traditional Chinese Medicine, Chang Gung University, Taiwan ABSTRACT. This study reported characteristics and predictive factors of early-onset diarrhoea in patients receiving pelvic irradiation. We retrospectively reviewed 229 patients undergoing radiotherapy alone for cervical or uterine cancer. Oral barium was taken 90 min before simulation. According to contrast medium within small intestine only or colon in simulation films, we categorised patients as normal and rapid transit groups. Small or large volume of small-bowel was also evaluated according to barium distribution of simulation films. Whole-pelvic irradiation (39.6–45 Gy/22–25 fractions) was delivered to all patients initially. We recorded the onset of diarrhoea during pelvic irradiation. The rates of early-onset diarrhoea (,10 Gy) were compared between these two groups. The incidence of diarrhoea before 10 Gy was 7% and 17% (p50.138) in patients with normal and rapid transit, respectively. In multivariate analysis, interaction among rapid transit, prior abdomen operation and large small-bowel volume (p50.019) were noted for early-onset diarrhoea. Further subgroup analysis revealed that rapid transit (p50.046) was a significant factor in patients with both prior abdominal operation and large small-bowel volume. The incidence of early-onset diarrhoea was as high as 40% in this particular group. Patients experiencing early-onset diarrhoea had a higher incidence of moderate to severe diarrhoea (65%) than those without early-onset diarrhoea (23%) (p,0.001). In multivariate analysis, early-onset diarrhoea was the only factor of moderate to severe diarrhoea (p50.001). In conclusion, rapid small-bowel transit may be predisposed to early-onset diarrhoea during pelvic radiotherapy in patients with both prior abdominal operations and large small-bowel volume. Early-onset diarrhoea is considered as a predictive factor of diarrhoea of a higher grade.

Radiation-induced acute diarrhoea is common during abdominal and pelvic irradiation. It results from a variety of different pathophysiological mechanisms, including malabsorption of bile salts and lactose, imbalances in local bacterial flora and changes in the intestinal patterns of motility [1, 2]. Although villus atrophy [3] of the small bowel is a key factor, the cause of diarrhoea is not only pathological, but also a functional change of the intestine. Changes in gastrointestinal (GI) motility during fractionated irradiation precede the appearance of histopathological lesions in the GI tract, and the symptoms experienced particularly within the first week of radiotherapy are directly related to changes in bowel motility [4]. The onset of diarrhoea during pelvic radiotherapy usually takes place in the second to third week of radiotherapy with conventional fractionation. However, unusually and inexplicably early onset of diarrhoea was noted in a few patients. Morphological change of small bowel can not explain the phenomenon. It is known that laxatives such as castor oil can rapidly induce diarrhoea by changing the small bowel motility [5]. Similarly, we hypothesize that patients with rapid small-bowel transit can be predisposed to early-onset diarrhoea during pelvic radiotherapy. Interestingly, we used multivariate analysis The British Journal of Radiology, May 2006

Received 4 May 2005 Revised 8 August 2005 Accepted 17 August 2005 DOI: 10.1259/bjr/51376226 ’ 2006 The British Institute of Radiology

and noted interaction among rapid small-bowel transit, prior abdominal operation and large small-bowel volume within pelvis. The result is that it is worth investigating the mechanism of early-onset diarrhoea in the future.

Materials and methods Patients’ characteristics Between January 1996 and January 2004, patients with histologically proven cervix or uterine malignancies were reviewed. The selection criteria of patients were four-field whole-pelvic irradiation and oral barium contrast taken 90 min before simulation [6]. The exclusion criteria of patients were: (1) receiving antispasmodic drugs, laxatives or chemotherapy before or during the first week of radiotherapy; (2) recent history of gastroenteritis; (3) history of diabetes mellitus; or (4) history of pelvis irradiation. Characteristics of 229 patients and treatment are shown in Table 1. Orthogonal films including anterior– posterior (AP) and lateral view were taken during simulation. In the AP view, at source–axis distance (SAD) 100 cm, x-axis size is the width (median: 15.5; range 13–18 cm) and y-axis size is the cephalocaudal distance 419

E Y Huang, C J Wang, H C Hsu and L M Sun Table 1. Characteristics of patients (n5229) Parameters

Age ,62 years > 62 years Diagnosis Cervical cancer Stage IB–IIA Stage IIB–IIIA Stage IIIB–IVB Post-operative Recurrent Endometrial cancer Stage I Stage II Stage III No operation Uterine sarcoma Prior abdomen operation Yes No Small-bowel volume Small volume Large volume Small-bowel transit Normal Rapid

Number

112 117

76 73 27 14 21 4 3 6 2 3 83 146 138 91 211 18

(median: 17; range 15–21 cm). In the lateral view, at SAD 100 cm, z-axis size is the ventrodorsal distance (median: 13; range 11–18 cm). The upper margin of AP film was located at L4–5. The lateral margin was 1–1.5 cm beyond the widest part of the pelvic brim. For lateral film, the anterior margin was anterior part pubic symphysis. The posterior margin is behind S2–3 at least. Before radiotherapy, an interview of each patient was performed for information of pelvic radiotherapy. We had explained the possible toxicities of radiotherapy. We informed each patient that she could ask for medication at the onset of symptoms. After the beginning of radiotherapy, we inspected patients weekly if no progression of symptoms was noted, especially during the first 1–2 weeks of radiotherapy. When symptoms attacked or progressed, immediate medication was prescribed to patients according to the bowel habits and stool consistency recorded in the charts. If patients had no improvement of initial diarrhoea with minor antispasmodic drug such as mebeverine, we could change to major antispasmodic medication such as loperamide and the onset of moderate to severe diarrhoea was recorded in the chart. Whole-pelvic irradiation (39.6–45 Gy) of four-field technique was initially administered with 10 MV or 15 MV photons through supine position. The daily fraction dose was 1.8 Gy, five fractions weekly. We reviewed the onset of diarrhoea in chart record. The severity of diarrhoea could be measured according to common toxicity criteria (CTC) [7]. Grade 1 is an increase of 2–3 stools per day over pre-treatment. Grade 2 is an increase of 4–6 stools per day, or nocturnal stools. Grade 3 is an increase of >7 stools per day or incontinence; or need for parenteral support for dehydration. Grade 4 is physiological consequences requiring intensive care; or haemodynamic collapse. It was considered moderate to severe diarrhoea when patients had any of the following 420

conditions: (1) CTC Grade 2 or greater diarrhoea; (2) no improvement of initial diarrhoea with minor antispasmodic drug and change to major antispasmodic medication; (3) watery diarrhoea. The dose of onset of diarrhoea was recorded for actuarial analysis.

Definition of small-bowel movement and volume distribution When barium filled the small intestine only, we defined this as normal bowel transit. If barium could be noted within the colon, rapid bowel transit was categorised. In order to evaluate small-bowel volume/distribution within the whole pelvis target, we used a simple volumetric method. We scored barium-filled amount of small bowel within the radiation fields in orthogonal films. If caudal small-bowel was below inferior aspect of sacroiliac (SI) joint and dorsal small-bowel was behind vertical line of sacral promontory, large-volume distribution was defined. Otherwise, the distribution was classified as small volume. Small or large volume presents corresponding upper/ anterior or lower/posterior position of small bowel within pelvis. The early-onset of diarrhoea was defined as diarrhoea before 10 Gy.

Statistics Univariate analysis of early-onset and moderate to severe diarrhoea rate was performed using Chi-square/ Fisher’s exact test. For multivariate analysis of earlyonset and moderate to severe diarrhoea, we used a logistic regression model with stepwise forward procedure. Age, rapid transit, small-bowel volume, and prior abdominal surgery were evaluated as categorical data. A p-value of less than 0.05 was considered to be statistically significant. The risk of diarrhoea or large-volume distribution was presented as odds ratio (OR) with 95% confidence interval (CI). Data processing and statistics were carried out on a personal computer using the software SPSS 12.0 for Windows (SPSS Inc., Chicago, IL).

Results Early-onset diarrhoea between normal and rapid small bowel transit groups There were 17 patients (7.4%) with diarrhoea before 10 Gy. No moderate to severe diarrhoea was noted before 10 Gy. The result of univariate analysis is shown in Table 2. The incidence of early-onset diarrhoea was 7% and 17% in patients with normal and rapid small bowel transit (p50.138), respectively. The corresponding rates were 4% and 40% (p50.068) in patients with both large volume and prior abdominal operation.

Multivariate analysis of early-onset diarrhoea among patients during pelvic irradiation We considered the possible predisposed factors for the analysis of early-onset diarrhoea. The results are shown The British Journal of Radiology, May 2006

Early-onset diarrhoea after pelvic irradiation Table 2. Univariate analysis of factors associated with early-onset and moderate to severe diarrhoea during pelvic irradiation Parameters

Early-onset

p-value

Moderate to severe

p-value

Age > 62 years (yes vs no) Prior abdominal operation (yes vs no) Large small-bowel volume (yes vs no) Rapid small-bowel transit (yes vs no) Prior abdomen operation and large volume (both vs another 3 groups) Prior abdomen operation and rapid transit (both vs another 3 groups) Rapid transit and large volume (both vs another 3 groups) Prior abdomen operation, large volume, and rapid transit (all vs another 7 groups) Early-onset diarrhoea (yes vs no)

7% 8% 10% 17% 10%

8% 7% 6% 7% 7%

0.730 0.660 0.248 0.138 0.459

24% 33% 29% 28% 38%

28% 22% 24% 26% 24%

0.517 0.078 0.430 0.785 0.109

29% vs 7%

0.087

43% vs 25%

0.378

18% vs 7%

0.192

27% vs 26%

1.000

40% vs 7%

0.045*

40% vs 25%

0.605

2

65% vs 23%

,0.001*

vs vs vs vs vs

2

vs vs vs vs vs

*Statistically significant.

in Table 3. If we did not consider interaction, no significant factors were noted. Interestingly, we noted interaction among operation, large volume and rapid transit (OR: 9.29, 95% CI: 1.44–59.93) (p50.019). Subgroup analysis revealed that rapid small-bowel transit was a significant factor of early-onset diarrhoea in patients with simultaneous large volume and prior abdominal operation (OR: 15.33, 95% CI: 1.05–224.78) (p50.046), but not another combinations.

Early-onset diarrhoea predicts moderate to severe diarrhoea among patients during pelvic irradiation Incidence of moderate to severe diarrhoea was 65%, and 23% in patients with and without early-onset diarrhoea (p,0.001), respectively. In multivariate analysis, early-onset diarrhoea was the only factor of moderate to severe diarrhoea (p50.001) (OR: 6.26, 95% CI: 2.20– 17.82) while age, small-bowel volume, prior abdominal operation, and small-bowel transit were adjusted.

Discussion One of the most serious complications of abdominal or pelvic radiotherapy is bowel damage, leading to nausea, vomiting, diarrhoea and abdominal cramps during or shortly after therapy. These symptoms are probably due to a rapid modification of the intestinal motility and to the structural alteration of the intestinal mucosa (cell loss and altered crypt integrity) [2]. Clinically, fractionated radiotherapy of the pelvis usually results in diarrhoea while dose is around 10–20 Gy. The mechanism is villus atrophy [3] caused by impaired crypt proliferation. According to our results, the incidence of early-onset diarrhoea was about 7%. It is difficult to explain diarrhoea before 10 Gy with the model of cell loss and altered crypt integrity because structural alteration of the intestinal mucosa may not appear before 10 Gy. Hence, small-bowel hypermotility may be considered for the mechanism of early-onset diarrhoea. The radiation-induced change of small-bowel motility had been reported in some investigations. Yeoh et al conducted a series of human and animal studies about

Table 3. Multivariate analysis of factors associated with early-onset and moderate to severe diarrhoea during pelvic irradiation Parameters

Age > 62 years Prior abdominal operation Large small-bowel volume Rapid small-bowel transit Interaction between operation and large volume Interaction between operation and rapid transit Interaction between rapid transit and large volume Interaction between operation, large volume, and rapid transit Early-onset diarrhoea

Early-onset

Moderate to severe p-value

p-valuea

0.668 0.900 0.495 0.882 0.600

0.562 0.087 0.606 0.846 2

0.562 0.087 0.606 0.846 0.137

2

0.704

2

0.609

2

0.506

2

0.810

2

0.019*b

2

0.978

2

2

0.001*c

0.001*c

p-value

p-value

0.730 0.660 0.248 0.119 2

a

*Statistically significant. a Consideration of interaction. b Odds ratio (95% confidence interval)59.29 (1.44–59.93). c Odds ratio (95% confidence interval)56.26 (2.20–17.82).


421

E Y Huang, C J Wang, H C Hsu and L M Sun

small-bowel motility during or after pelvis/abdomen irradiation. During radiotherapy, increased stool frequency (p,0.001) was associated with more rapid smallintestinal (p,0.01) and whole-gut (p,0.05) transit [8]. 1–6 years after radiotherapy, small intestinal transit (p,0.05) was inversely related to stool frequency [9]. In an animal study, there was an initial increase in frequency followed by a non-significant reduction, but not in the amplitude of ileal pressure waves following fractionated abdominal irradiation (22.5 Gy in 9 fractions at 3 fractions/week). There was no relationship between motility and histology [10]. Erickson et al used an animal model with the same fractionation that dramatically increased the frequency of giant migrating contractions (GMCS) of the small intestine, with the incidence peaking after the second dose. The increased frequency of GMCS associated with abdominal cramps and diarrhoea occurred as early as a few hours after the first radiation fraction, and returned to normal within days of cessation of radiation. Altogether, hypermotility of small bowel is associated with diarrhoea during fractionated irradiation. Our multivariate analysis revealed interaction among rapid transit, prior abdominal operation and large smallbowel volume in pelvis. Subgroup analysis showed that rapid transit was a significant factor of early-onset diarrhoea in patients with both prior abdominal operation and large volume. Prior abdominal operation [11– 13] and small-bowel volume [11–15] are associated with radiation-induced enteropathy. Our previous study noted volume effect of small bowel for acute diarrhoea in patients with prior abdominal operation [16]. When patients have both factors, the possibility of large fulldose volume is high. If small-bowel transit is rapid before radiotherapy, there may be hypermotility within a large volume of the small bowel irradiated. While water can not be absorbed in this segment of small bowel, early-onset diarrhoea may appear. Jankovic et al noted that X-radiation produced acute sensitization (2 h) of rat GI tract to acetylcholine and histamine [17]. Hence, radiation-induced hypermotility effect of neurotransmitters may be enhanced in patients with rapid transit. The effect of abdominal operation in small bowel distribution is important and associated with enterotoxicity. If small bowel is located in posterior pelvis (i.e. large-volume distribution), patients with prior abdominal operation may have a fixed loop of small bowel that could be repeatedly exposed by fractionated radiotherapy. Gallagher et al [11] reported that pelvic smallbowel volume was larger after pelvic surgery, especially abdominoperineal resection. In post-operative patients with rectosigmoidal and endometrial cancers, 26–65% of fixed small bowel was noted in the posterior pelvis [14, 18, 19]. Operation could exacerbate radiation-induced acute diarrhoea in rectal cancer because small bowel could be fixed in posterior pelvis by operation [12, 20]. We could calculate comparable large-volume distribution (35%) in our study. Limited movement of small bowel within posterior pelvis caused by operation may result in repeated large-volume and full-dose exposure of small bowel. If these patients have rapid small-bowel transit before radiotherapy, hypermotility of large volume can result in early-onset diarrhoea. Some methods of measurement for transit time of small intestine were reported. Radioisotopes allow 422

accurate quantitation of the pattern and effectiveness of the transit of chyme through the small and large intestines [21, 22]. Birkebaek et al measured healthy adult test subjects using a scintigraphic method with 111 In marked single unit tablets. The small intestine transit time was 5 h (median) with an interquartile range of 4–7 h [23]. Benmair et al used a magnetic method for the measurement of small intestinal transit time in a group of 20 normal subjects [24]. The mean transit time was 157.5¡63.9 min. Pia de la Maza et al used hydrogen breath test [25] and noted mean intestinal orocecal transit time decreased from 1.7¡0.9 h to 1.36¡0.9 h after 5 weeks of radiation (p,0.05) [26]. The variations may depend on the methodology. Simulation at 1.5 h after barium intake seems to be applicable for mimicking transit time in our study. Although oral contrast is not the optimal method for transit time of small bowel, we could still note increased incidence of early-onset diarrhoea in patients with short transit time. The severity of early-onset diarrhoea was mild in our results. The role of early-onset diarrhoea in moderate to severe diarrhoea was tested in our study. We noted that early-onset of diarrhoea was predisposed to moderate to severe diarrhoea in both univariate and multivariate analysis. Patients experiencing early-onset of diarrhoea may be radiosensitive. Hence, more accumulative damage and more severe diarrhoea developed in those who received further fractionated radiotherapy after early-onset diarrhoea. The result suggests that more severe diarrhoea may appear if early onset diarrhoea is noted. There are some flaws in the present study. The smallbowel volume and transit were evaluated by orthogonal simulation films. Although they are not standard methods, they are acceptable for prediction. Because diabetes mellitus [27–31] and chemotherapy [12, 32] could influence diarrhoea, we excluded patients with these factors and sample size was decreased. Furthermore, the incidences of rapid transit and earlyonset diarrhoea were low. Hence, a large sample size is needed to present statistical differences. Regardless, the role of small-bowel transit was established while interaction was considered for multivariate analysis. This is the first clinical result investigated about the association between small-bowel transit and early-onset diarrhoea. Prospective study with more sensitive evaluation for small-bowel volume and transit is encouraged. Although we focused on acute diarrhoea during whole pelvis irradiation, late bowel complications were also important because quality of life was impaired [33, 34]. Andreyev et al stated that most studies have used inadequate assessments of GI toxicity [35]. The population of our patients included those receiving operations or not. Hence, the radiation dose, field size, and boost technique varied after 39.6–45 Gy whole pelvis irradiations. These factors limited effective assessment of influence of early-onset diarrhoea in late bowel complications. However, the prospective study using more comprehensive records of symptoms, such as diary and questionnaire, is encouraged. Furthermore, our previous study used early-onset diarrhoea as part of severity of diarrhoea and showed correlation between acute and late toxicity [36]. If early-onset diarrhoea appears, the patient may have more severe diarrhoea during The British Journal of Radiology, May 2006

Early-onset diarrhoea after pelvic irradiation

subsequent irradiation and increased incidence of late bowel complications. More supportive care and dose/ target volume modification to avoid late complications may be considered. In conclusion, early-onset diarrhoea may be associated with rapid small-bowel transit in patients with simultaneous prior abdominal operation and large volume of small-bowel irradiated. Patients with early-onset diarrhoea are predisposed to more severe diarrhoea during subsequent pelvic radiotherapy.

16.

17.

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Optical density variations in CT films and their effect on image quality 1

I A TSALAFOUTAS,

PhD,

2

G V PAPOUTSIS,

RT,

3

P N MANIATIS,

MD

and 4K A GOGOS,

MSc

1

Medical Physics Unit, ‘Konstantopoulio - Agia Olga’ Hospital, 3-5 Agias Olgas, Nea Ionia, 142 33, Athens, 2Field Service Engineering Department, General Electric Medical Systems, 156 Cyprou Avenue and 91 Konstantinoupoleos Str, Argyroupolis, 164 51, Athens, 3Computed Tomography Department, ‘Konstantopoulio - Agia Olga’ Hospital, 3-5 Agias Olgas, Nea Ionia, 142 33, Athens and 4Nuclear Medicine Department, ‘Hygeia’ Hospital, 4 Erythrou Stavrou, 151 23, Maroussi, Athens, Greece ABSTRACT. It was recently reported that optical density (OD) variations were observed in CT films printed with a laser camera, depending on the printing format and the frame position within the film. The purpose of the present study was to investigate if these variations are common to both laser and dry-film printers and if the different OD settings along with day-to-day and frame-to-frame variations may affect the image quality. Eight laser and five dry-film printers installed at 12 different CT facilities were tested. For each one, the SMPTE test pattern was printed on all frames of a film using the same printing format. The ODs of the 0%, 10%, 40% and 70% patches of the 11step greyscale of the SMPTE patterns were measured with a densitometer in all frames, while all films were examined on a viewing box to assess subjectively the image quality by visual inspection of the test pattern. A wide range of OD settings and variations were recorded. Frame-to-frame variations in the same film of up to 0.19, 0.15 and 0.21 OD, were observed for contrast index (CI, the OD difference of patches 10% and 70%), speed index (SI, the OD of patch 40%) and maximum OD (ODmax, the OD of patch 0%), respectively. The variations were not always of the same magnitude, nor always followed the same pattern, even for printers of the same model. Considering all films and frames, the CI ranged from 1.26 to 1.74, the SI from 0.68 to 1.43 and the ODmax from 2.5 to 3.11 OD, well beyond the proposed settings and tolerances of 1.55¡0.15, 1.15¡0.1 and 2.45¡0.1 given in the literature for CI, SI and ODmax, respectively. Despite these large differences, the various problems that were identified in image quality from the visual inspection of the films could not be directly attributed to OD settings, as films with similar CI, SI and ODmax presented quite different image quality levels. Therefore, for routine quality control, thorough visual inspection of the SMPTE test pattern provides all the necessary information about the imaging chain status.

CT was first introduced into clinical practice in the early 1970s and since then has become one of the most useful diagnostic tools. The technology of CT scanners is continuously and rapidly evolving while significant developments have also occurred in CT film printers, as laser and dry printers superseded CRT based multiformat cameras (hard-copy cameras). Nowadays, in CT departments, diagnosis is more and more frequently performed on workstation monitors. However, film still remains an important means for diagnosing, storing and transmitting CT images. Indeed, whilst monitors and films are both currently used for the interpretation of CT images, in many cases diagnosis is based solely on film reading. Since for viewing a CT image, the window width (WW) and window level (WL) are adjusted so that the structures of interest are best imaged on the monitor, it is consequently important to ensure that the image characteristics are maintained when printed on film. The British Journal of Radiology, May 2006

Received 4 May 2005 Revised 9 September 2005 Accepted 13 September 2005 DOI: 10.1259/bjr/28579947 ’ 2006 The British Institute of Radiology

In order to test medical imaging systems and hardcopy devices, the Society of Motion Picture and Television Engineers (SMPTE) formed a committee that developed and tested a number of patterns, until they agreed on the final version of the SMPTE monochrome test pattern that provides both qualitative and quantitative information [1]. The characteristics of SMPTE pattern, its use for acceptance and quality control (QC) purposes and relevant technical details on the design of film printers are summarized in the report of the American Association of Physicists in Medicine (AAPM) Diagnostic X-ray Imaging Committee Task Group No 1 [2]. In the AAPM report [2], reference values are given for video and laser cameras concerning the optical densities (OD) that certain patches of the 11-step greyscale of the SMPTE pattern should have when printed on film, along with the OD variation tolerances for the QC of these systems. However, in a recent study [3] it has been 425

I A Tsalafoutas, G V Papoutsis, P N Maniatis and K A Gogos

reported that the ODs of these patches vary with printing format and with frame position within the same film. In the present study, the frame-to-frame OD variations within the same film were further investigated for a number of laser and dry printers, to identify if these variations are dependent on the type of film printer and whether being combined with different OD settings and day-to-day variations can affect the image quality in any way.

Materials and methods In this study, 12 different CT facilities were included; 7 equipped with a laser and 4 with a dry-film printing device, while in one facility two film printers (one laser and one dry) were interfaced to the same CT scanner. All CT scanners were manufactured by the same company (GE Medical Systems, Milwaukee, WI), while film printers were manufactured by Agfa (Agfa-Gevaert A.G., Mu¨nchen, Germany) and Kodak (Eastman Kodak Company, New York, NY). In each facility, the SMPTE test pattern was viewed on the CT monitor, setting the WW to 100 and the WL to 0 (except for the Hispeed models where the WL was set to 1024 automatically when the SMPTE was imaged) [2]. The SMPTE test pattern is shown in Figure 1 and a detailed description of its characteristics is given in the AAPM report [2]. According to the suggestions of the AAPM report, the brightness and contrast of the TV monitor can be considered adequately adjusted when the 5% and 95% average picture level patches inset, respectively, in the 0% and 100% patches (located at the two ends of the 11-step greyscale) and the low and high contrast bar-patterns (located at the centre and the

four corners of the SMPTE test pattern) are clearly discernible. In all TV monitors examined, these criteria were fulfilled and the adjustment of contrast or brightness was not considered necessary. Subsequently, and without changing the window settings, the SMPTE pattern was printed so as to cover the 20 frames of a film (5 rows by 4 columns). The OD values of the 0%, 10%, 40% and 70% patches of the 11step greyscale SMPTE patterns of all frames were measured with a calibrated optical densitometer (RMI 331 Densitometer; X-Rite, Grandville, MI). The OD of patch 40% is defined as the speed index (SI), the OD difference of patches 10% and 70% as the contrast index (CI), while ODmax is the OD of patch 0% [2]. In two CT facilities, the above procedure was repeated a few months after the first test in order to identify possible day-to-day variations. In one CT facility, the test procedure was carried out three times with the same laser camera and three different film-types (after applying the necessary sensitometric adjustments), while for one film type the test was also repeated a few months after the first test. In the latter facility, in order to adjust the photographic processor after film changes, the procedure described in the user manual of the laser camera was carried out. The laser camera produced a 21step greyscale setting the target density at the value of 3.3 OD, as proposed by the manufacturer. The maximum OD (the OD of the 21st step) was measured with the optical densitometer and was manually fed into the camera. This procedure was repeated until a film with maximum OD within ¡0.05 of the target density was produced. Next the ODs of all steps were measured and manually fed into the laser camera for updating the lookup tables [2].

Figure 1. The SMPTE test pattern. 426


Optical density variations in CT films and image quality

In addition to the OD measurements, all the films were observed on the same viewing box to assess the qualitative characteristics of the printed SMPTE patterns. Two observers graded in consensus the images for contrast (in both low and high ODs) and resolution (in both low and high contrast) using an arbitrary five-point scale ranging from 1 (not discernible) to 5 (perfectly discernible) and commented on the presence of any type of artefacts. For contrast grading, the visibility of the 5% and 95% insets to the 0% and 100% patches, respectively, was assessed. For low contrast resolution grading, the visibility of the low contrast bar-patterns was assessed; similarly, for high contrast resolution grading, the high contrast bar-patterns located beside the respective low contrast bar-patterns were evaluated.

Results Overall, 18 films acquired from 13 different printers were examined. The mean values and the maximum variation (max2min) of the CI, SI and ODmax recorded in the 20 frames of each film are given in Table 1, along with the results of the qualitative assessment. The tests with decimal numbers indicate those tests made on the same film printer, in order to investigate the impact of film changes and day-to-day variations on frame-toframe OD variations and image quality. Most of the observed OD variation patterns are given in Figures 2–6. The variation patterns depicted in Figure 2 refer to the same laser printer (Scopix LR 3300) for 4 tests made with 3 different films. The corresponding sensitometric curves (used for updating the lookup tables) are depicted in Figure 7. In Figure 3, the OD variation patterns of 6 laser printers of the same model (Scopix LR 3300) are given (for printers 1 and 2, the results from the tests that exhibited the largest variations are depicted). In Figure 4, the OD variation patterns of 3 different laser printers (Scopix LR 3300, Scopix LR 5200, Ektascan 2180) are given (for Scopix LR 3300, the results from the test that exhibited the largest variations is depicted). In Figure 5, the OD variation patterns of 5 dry printers (three DryView 8100, one DryView 8700 and one Drystar 3000) are given. Finally, in Figure 6, the OD variation patterns of the laser and the dry printers that exhibited the largest frame-to-frame variations (one Scopix LR 3300 and one DryView 8100, respectively) are given for comparison. The OD variation patterns of CI, SI and ODmax given in Figure 2 present some similarities. All these parameters obtain their maximum values in the upper left frame and then, when moving to the right, decrease to a minimum (at the 3rd frame) and then increase again. This pattern is repeated for each row and in some cases a slight decrease from the first to the last row is also observed. The mean values of CI, SI and ODmax are quite similar even if the respective sensitometric curves are quite different. However, the maximum variations in CI and ODmax ranged from 0.07 to 0.15 and 0.06 to 0.16, respectively, with no obvious correlation with the sensitometric curves. It is important to note that when the film was changed from Agfa to Ferrania, the sensitometric curve (shown in Figure 7 with the dashed line) was The British Journal of Radiology, May 2006

significantly faster than that of Agfa, obtaining a maximum OD of 3.72 instead of the 3.3 target OD. After the look-up tables were updated, the sensitometric curve of the Ferrania film was found to be significantly slower than that of Agfa film. However, the CI, SI and ODmax were almost equal to those obtained with the Agfa film. It can also be deduced from the tests repeated with the same film (tests 1.3–1.4, 2.1–2.2, 8.1–8.2) that day-to-day variations did not significantly change the OD variation pattern. However, differences in the maximum frame-toframe OD variations and the mean values of CI, SI and ODmax were observed, as in the case of the Ektascan 2180 (tests 8.1–8.2) where the mean value of SI changed between tests from 1.02 to 1.39. In Figure 3, it is apparent that not all Scopix LR3300 presented the same OD variation pattern or the same maximum variations in CI, SI and ODmax. In Figure 4, the more extreme pattern observed for Scopix LR3300 is compared with the less extreme patterns observed for Scopix LR5200 and Ektascan 2180 laser printers. In Figure 5, one can appreciate the variety of OD variation patterns observed for the various dry printers tested. Finally, in Figure 6 it is obvious that the variation patterns of the laser and dry printers that exhibited the largest CI, SI and ODmax frame-to-frame variations are completely different. As far as the qualitative assessment of images is concerned, films with reduced contrast and resolution, and various artefacts were identified. These problems were common to all film frames and no frame-to-frame variations in image quality were observed, with one exception discussed in the end of this section. No problems were observed in the contrast in the high ODs, as the 5% inset patch was always perfectly (grade 5) or well (grade 4) discernable from the 0% background. However, for low ODs there were three films (tests no. 5, 2.1 and 4) where the 95% inset patch was not discernable at all (grade 1) or was poorly discernable (grade 3) from the 100% background. While this could be attributed to the very low SI of all three films, for the film from test no. 2.2, which also exhibited a very low SI, the respective contrast grade was 4. On the other hand, the deterioration of low and high contrast resolution observed in some of the films could not be associated with OD settings. While in some cases a reduced resolution could be attributed to other printer problems causing geometrical distortion and background mottle (as in tests no. 2.1, 2.2, 9 and 11), a reduced low contrast resolution was also observed in the films from tests no. 8.1 and 8.2 with no obvious reason. Finally, it is worth mentioning that the only exception in the image quality homogeneity among frames was the film from test no. 4. In this film, alternating bands of white-black and white-grey line pairs in the high contrast bar patterns (an artefact we termed ‘‘banding’’) were apparent in all frames. However, this artefact was more intense in the outer left frames (the first frame of each row) and specifically in the upper left and lower left high contrast bar patterns. This was indeed the only case where a definite variation of image quality among frames was observed. 427

428 Table 1. Results of the optical density (OD) measurements and visual assessment of image quality in the 18 films studied. The mean values of contrast index (CI), speed index (SI) and maximum OD (ODmax), and the maximum frame-to-frame differences (max2min) observed are shown. Mean values outside the ranges (1.55¡0.15, 1.15¡0.1 and 2.45¡0.1 for CI, SI and ODmax, respectively) proposed for the quality control (QC) of a laser camera [2], are given in italics. The maximum frame-to-frame differences observed for laser (tests 8.1– 8.2) and dry printers (tests 9–13) are given in bold. The contrast grading in the high and low ODs (around the ODs of inset patches 5% and 95%, respectively) and the low contrast (LC) and high contrast (HC) resolution grading are shown, as well as comments on the presence of artefacts CT

Cameraa

Filmb

CI

max2min

SI

max2min

ODmax

max2min

Contr. Contr. 5% 95%

LC HC Res. Res.

Comments

1.1 1.2 1.3 1.4 2.1 2.2 3 4

Prospeed Prospeed Prospeed Prospeed Sytec Sytec Hispeed Prospeed

Scopix Scopix Scopix Scopix Scopix Scopix Scopix Scopix

LT 2B daylight Lifeimager HNB KAP LP 670 A KAP LP 670 A LT 2B daylight LT 2B daylight LT 2B daylight LT 2B daylight

1.53 1.56 1.65 1.58 1.29 1.36 1.58 1.36

0.15 0.07 0.09 0.1 0.08 0.08 0.19 0.17

1.05 1.06 1.07 1.08 0.86 0.88 1.19 0.78

0.08 0.08 0.09 0.06 0.04 0.04 0.15 0.07

2.87 2.90 2.91 2.90 2.59 2.68 2.99 2.59

0.16 0.09 0.06 0.12 0.1 0.1 0.21 0.19

5 5 5 5 5 5 5 5

5 5 5 5 3 4 4 3

4 4 5 5 2 2 5 4

5 5 4 4 3 3 4 5

5 6 7 8.1 8.2 9

Prospeed Hi-Speed Prospeed Prospeed Prospeed Prospeed

Scopix LR3300 Scopix LR3300 Scopix LR5200 Ektascan 2180 Ektascan 2180 DryView 8100

LT 2B daylight LT2B daylight LT 2B daylight EHG-10 EHG-10 DVB

1.44 1.67 1.69 1.43 1.46 1.49

0.16 0.10 0.08 0.1 0.09 0.13

0.69 1.31 1.18 1.02 1.39 1.24

0.05 0.04 0.04 0.04 0.07 0.1

2.57 2.93 3.02 2.85 2.89 2.82

0.18 0.10 0.1 0.18 0.1 0.2

5 4 4 5 5 4

1 5 5 5 5 4

4 4 4 2 3 1

5 4 5 4 4 4

10 11f

Hi-Speed Prospeed

DryView 8100 DryView 8100

DVB DVB

1.63 1.48

0.06 0.05

1.23 1.16

0.1 0.04

3.01 2.72

0.08 0.06

4 5

5 5

4 3

4 4

12 13

Prospeed Hispeed

DryView 8700 Drystar 3000

DVB DT1B Dry

1.65 1.53

0.07 0.12

1.02 1.35

0.1 0.07

2.66 2.87

0.07 0.12

4 4

4 5

4 4

5 4

– – ‘‘Banding’’c – Large GDd + ‘‘banding’’ Large GD + ‘‘banding’’ ‘‘Banding’’ Variable ‘‘banding’’, ‘‘overshoot’’e – – Large ‘‘overshoot’’ – – Bgndg mottle (horizontal white stripes), ‘‘banding’’ ‘‘Banding’’ Bgnd mottle (thin vertical white and grey lines) Large ‘‘overshoot’’ ‘‘Banding’’

a

LR3300 LR3300 LR3300 LR3300 LR3300 LR3300 LR3300 LR3300

Scopix LR 3300, Scopix LR 5200 and Drystar 3000 printers are manufactured by Agfa, while Ektascan 2180, DryView 8100 and DryView 8700 are manufactured by Kodak. LT2B daylight and DT1B Dry films are manufactured by Agfa (Agfa–Gevaert N.V., Mortsel, Belgium), EHG-10 and DVB by Kodak (Eastman Kodak Company, New York, NY), Lifeimager HNB by Ferrania (Ferrania S.p.A, Ferrania (SV), Italy) and KAP LP 670 A by Konica (Konica Corporation, Tokyo, Japan). c ’’Banding’’ stands for the presence of alternating bands of white-black and white-grey line pairs in the high contrast bar patterns in horizontal, vertical or both directions. d GD stands for geometrical distortion (straight lines appear as curves). e In films with ‘‘overshoot,’’ the high contrast bar-patterns extend beyond their limits casting their shadow within the adjacent areas. f Test 11 was carried out in the same CT facility as tests 8.1 and 8.2. g Bgnd is an abbreviation for ‘‘Background’’ of the SMPTE test pattern where the OD is equal to that of the 50% patch. b



Test


Figure 2. The variation of the contrast index (CI), speed index (SI) and maximum optical density (ODmax) values with frame numbers for tests 1.1 (Agfa film), 1.2 (Ferrania film), 1.3 and 1.4 (Konica film) performed on different days with the same laser printer (Scopix LR3300, Agfa) at the same CT installation. The frames have been numbered starting from the upper left corner, and vertical lines indicate the start of each of the 5 rows with 4 frames each.

Discussion The AAPM report [2] gives mean values and acceptable tolerances of CI, SI and ODmax as 1.55¡0.15, 1.15¡0.1 and 2.45¡0.1, respectively. These values, however, refer to the periodic quality control (QC) of a laser camera where the same format and frame position are monitored using the Kodak SO-497 film. These serve as a reference for the accepted OD settings and day-today variations but are not applicable to frame-to-frame variations. For the latter variations, the AAPM report [2] recommends that OD differences of ¡0.05 or slightly wider within one sheet of film may be accepted. However, the report does not state whether exceeding the proposed tolerances could have an effect on image quality. Furthermore, as has been noted in a previous study [3], limits have not been set for frame-to-frame variations indicating a certain malfunction of the film printer. The reason for frame-to-frame variations has not yet been identified. While some possible reasons have been proposed by Agfa technicians concerning the The British Journal of Radiology, May 2006

Figure 3. The variation of the contrast index (CI), speed index (SI) and maximum optical density (ODmax) values with frame numbers for tests performed at different CT installations which all had the same laser printer model (Scopix LR3300, Agfa) and used the same film (LT 2B daylight, Agfa).

results of our previous paper [3], in view of the variations also observed in dry printers, where different technologies are used for film exposing and processing, none seem very convincing. The other image quality problems identified in the films studied could be attributed to a number of factors (e.g. the geometrical distortion is most probably due to a worn polygon mirror) and can usually be identified and repaired by the camera manufacturer technicians, who usually utilize other types of test patterns that the cameras are able to produce. As written in the AAPM report [2], the SMPTE test pattern can be used to identify image quality problems, but not to trace the exact origin of the problem within the imaging chain. This requires further investigation by the camera (and perhaps the CT) technicians and it is outside the scope of this paper. Whilst we are eager to study the above issues further, this is not currently possible because of the lack of detailed information on design and sources of problems provided by the companies concerned. The mean CI values given in Table 1 (except for tests 2.1, 2.2 and 4) are within the reference range. However, many of the SI and all the ODmax values given in Table 1 are outside the range of values given by the AAPM 429


Figure 4. The variation of the contrast index (CI), speed index (SI) and maximum optical density (ODmax) values with frame numbers for tests performed at different CT installations, using three laser printer models (3: Scopix LR3300Agfa; 7.1: Scopix LR5200-Agfa; 8.1: Ektascan 2180-Kodak).

report [2]. Thus, the printers and films investigated in our study exhibit a larger variability in performance. The maximum differences in CI, SI and ODmax observed in the present study due to frame-to-frame variations reached 0.19 OD, 0.15 OD and 0.21 OD, respectively (film no. 3). Thus, if both day-to-day and frame-to-frame variations are taken into account, the CI, SI and ODmax on any frame for a specific film printer could well vary within ¡0.25, ¡0.2 and ¡0.2, respectively, from the central OD settings. It must be also noted that considering all the frames of all films, the CI ranged from 1.26 to 1.74 (0.48 OD difference), the SI from 0.68 to 1.43 (0.75 OD difference) and the ODmax from 2.5 to 3.11 (0.61 OD difference). Despite these large OD variations, image quality as assessed by visual inspection of the SMPTE test pattern could not be correlated with the OD values. The results of this study confirmed that frame-toframe OD variations were common among the different printers tested, though were not always of the same magnitude or pattern. However, an OD variation large enough to reduce image quality only for certain frames of the same film was not identified. Therefore, whatever the frame-to-frame differences between ODs, if the SMPTE test pattern characteristics remain clearly and equally discernable in all frames during visual 430

Figure 5. The variation of the contrast index (CI), speed index (SI) and maximum optical density (ODmax) values with frame numbers for tests performed at different CT installations, using different dry printer models (9,10,11: DryView 8100-Kodak; 12: DryView 8700-Kodak; 13: Drystar 3000Agfa).

inspection, it should be safe to assume that the quality of printed clinical CT images remains adequate for diagnosis. In conclusion, we propose that for a certain CT facility, once the reference values for the OD of all the patches in the 11-step grey scale of the SMPTE test pattern and the CI, SI and ODmax have been determined, as well as the frame-to-frame OD variations (at least for the most commonly used printing format), periodic visual inspection of the SMPTE test pattern on the TV monitor and on all the frames of a film should be adequate for routine QC purposes. However, when from the visual inspection a loss of image quality is suspected, measurements of the ODs in a film fully covered with the SMPTE test pattern should be carried out to determine whether the camera should be readjusted or repaired.

References 1. Specifications for medical diagnostic imaging test pattern for television monitors and hard-copy recording cameras. SMPTE Recommended Practice 1986, RP 133-1986. SMPTE Journal 1986;95:693–5.



Figure 7. The sensitometric curves used for updating the look-up tables before acquiring the SMPTE films (tests 1.1 to 1.4). The dashed line is the sensitometric curve that resulted when the film changed from Agfa to Ferrania before updating the look-up tables. The latter curve is included in the graph to demonstrate the influence of look-up tables on the sensitometric characteristics of CT films.

Figure 6. The variation of the contrast index (CI), speed index (SI) and maximum optical density (ODmax) values for the laser and dry printers that exhibited the largest frame-toframe variations (3: Scopix LR3300-Agfa; 9: DryView 8100Kodak).


2. Gray JE, Anderson WF, Shaw CC, Shepard J, Zeremba LA, Lin PP. Multiformat video and laser cameras: history, design considerations, acceptance testing and quality control. Report of AAPM Diagnostic X-ray Imaging Committee Task Group No1. Med Phys 1993;20:427–38. 3. Tsalafoutas IA, Tsapaki V, Koulentianos E, Triantopoulou C. Quality control of a laser camera with the SMPTE test pattern: optical density variations with printing format and frame position. Br J Radiol 2004;77:52–6.

431


SHORT COMMUNICATION

Assessment of environmental disturbances to the static magnetic field in magnetic resonance installations M A SCHMIDT,

PhD

Department of Medical Physics, St George’s Hospital, Blackshaw Road, London SW17 0QT, UK ABSTRACT. The static magnetic field of MRI scanners can be affected by environmental factors. Magnetic resonance spectroscopy and functional imaging with single-shot echo-planar imaging (EPI) are particularly vulnerable to the movement of lifts, vehicles, trains and other large metallic masses in the vicinity. This work investigates the sensitivity of two different imaging techniques to assess disturbances of the static magnetic field: (i) phase changes in gradient-echo images of a uniform test object; and (ii) image displacement along the phase encoding direction in single-shot EPI images. For the latter a hexane sample was used, and the separation between CH2 and CH3 signals was taken as a reference. Both techniques were evaluated in a site known to be free of any significant environmental disturbances and validated by inducing a magnetic field disturbance. Both techniques provide valuable information in acceptance tests, allowing MRI users to evaluate and manage the environmental conditions surrounding a scanner. The single-shot EPI technique was found to be highly sensitive, being expected to detect magnetic field fluctuations down to 0.005 parts per million (ppm). The phase images method was found to be less sensitive (0.02 ppm) but is more easily available. The single-shot EPI technique was used in acceptance tests and environmental disturbances to the magnetic field of the order of 0.04 ppm were measured at the isocentre on two separate occasions.

MRI scanners are known to be affected by external environmental factors. Examples are movement of lifts, vehicles, trains and other large metallic objects, often having some ferromagnetic content, which disturb the static magnetic field surrounding the magnet isocentre. Functional imaging and other techniques based on echoplanar imaging (EPI) are particularly vulnerable: Durand et al [1] have shown that changes in currents associated with a rail line caused significant deterioration on functional brain studies at 1.5 T. Peaks in magnetic resonance spectra can also suffer considerable broadening, reducing magnetic resonance spectroscopy (MRS) sensitivity. MRI equipment suppliers advise on site planning and on minimum distances between the magnet isocentre and various sources of disturbance to the magnetic field. Some MRI scanners are placed relatively close to possible sources of disturbance, and checking for any environmental effects then becomes a necessary part of the acceptance test. It may also be necessary to verify the stability of the magnetic field to aid the management of the environment surrounding an MRI system throughout its life-span. At any given point within the magnet bore, a transient environmental disturbance to the static magnetic field will be perceived as a fluctuation of the Larmor frequency. A second effect is a transient reduction of the field homogeneity within a given volume. MRI 432

Received 19 April 2005 Revised 26 October 2005 Accepted 26 October 2005 DOI: 10.1259/bjr/76396327 ’ 2006 The British Institute of Radiology

scanners always adjust the central frequency prior to imaging, but the timescale of this automatic measurement is invariably slow, not allowing the user to measure a transient magnetic field fluctuation reliably. A localized measurement of the magnetic field fluctuation could be made with a small sample, simply by repeatedly acquiring a free induction decay (FID) signal. However, in practice, this is extremely difficult to undertake in acceptance tests, as the required pulse sequence is often unavailable. This work investigates the sensitivity of widely available imaging techniques to evaluate disturbances of the static magnetic field. Current shimming standards allow the routine acquisition of brain 1H spectra with 2 Hz water line width at 1.5 T, and thus measurement techniques must be sensitive to magnetic field variations at least of the order of 0.01 parts per million (ppm). To be truly useful in acceptance testing, a technique for the measurement of changes to the static magnetic field must not depend on any detailed knowledge of the pulse sequences involved or other information to be provided by the manufacturer. Also, it must not depend on specific software packages, which may not be available. It is also desirable to produce an objective record of any transient changes in the static magnetic field intensity. The latter is particularly relevant for disturbances which are not very frequent, but are expected to be significant (an infrequent train service in the vicinity, for example). The British Journal of Radiology, May 2006

Short communication: Assessment of disturbances to the magnetic field in MR installations

Methods Two separate imaging techniques were considered for the quantification of magnetic field disturbances caused by environmental factors, and their sensitivity was evaluated. This work was undertaken at the Radiology Department, St George’s Hospital (GE Signa/Echo-Speed gradients, Milwaukee, WI). A 1.5 T MRI scanner has been in operation on this site for 10 years and the site is known to be free of any significant disturbances to the static magnetic field due to the surrounding environment.

Phase images method Changes in the magnetic field value are demonstrated as phase changes on gradient-echo images, since the term c.TE.b(t) is added to the image phase in the presence of magnetic field fluctuations (c is the gyromagnetic ratio, TE is the echo-time and b(t) is an unknown magnetic field fluctuation). This technique can be used in any scanner, but some of the main manufacturers require a research agreement or a service password to produce phase, real and imaginary images. Manufacturers also use different approaches for the scaling of the phase images, sometimes using a threshold from the magnitude images to avoid displaying the random background phase. When the phase scaling for a given scanner is not known, the phase must be calculated from real and imaginary images. Another solution to eliminate any ambiguity is to impose a known shift to the central frequency, as this will cause a known phase change, and acquire two images (before and after shift) for reference. A self-loaded uniform cylindrical test object with 19 cm diameter was employed (NiCl solution, 1.66 g l21, pH 4). A relatively short T1 was preferred to maximize the signal-to-noise ratio (SNR) in rapid imaging techniques. The phase change associated with field fluctuations can be maximized by increasing TE, but this slows down the measurement and leads to poorer SNR. As a compromise, a series of 100 fast spoiled gradient-echo images were acquired with the following parameters: TE53.3 ms (shortest TE for a symmetrical echo), repetition time (TR)57.3 ms, 24 cm field of view (FOV), 10 mm slice thickness, 2566128 data acquisition matrix and ¡32.25 kHz receiver bandwidth. These parameters produce a 2p phase shift for a 4.4 ppm shift on the Larmor frequency. The first phase image acquired was taken as a reference image, and all subsequent phase images were subtracted from the reference image. A region of interest (ROI) was chosen at the isocentre and the ROI mean phase value was computed over the resulting series of subtracted phase images. The range of ROI phase values was used to estimate the sensitivity of the method.

the interecho delay. However, details of the EPI pulse sequence may not be known at the time the acceptance test is performed, since manufacturers differ on the amount of pulse sequence information provided as standard. Again, a known shift in the central frequency will produce a reference displacement, which can be used to calibrate any other images where a displacement along the PE direction occurs. Another solution is to use a sample which contains more than one chemical shift, and use the known separation of the two peaks as a reference. The use of the single-shot EPI pulse sequence is widespread, but it is not necessarily available in every scanner as it is usually sold as a separate software package. An 8 mm diameter test tube containing hexane (C6H14) was employed, and the known chemical shift difference between the CH2 and CH3 peaks (0.39 ppm) was used as a reference. Both peaks have similar intensity and are close enough to produce two clear images if the central frequency is chosen as the midpoint between the two peaks. A series of 100 single-shot spin-echo EPI images was acquired at the rate of one image per second with the following parameters: TE5100 ms, bandwidth ¡31.25 kHz, slice thickness 10 mm, FOV 24 cm624 cm, producing a 2566256 image using a half-Fourier acquisition. These parameters were chosen to make the technique as sensitive as possible by increasing the PE bandwidth. The two locations of the sample tube in the images were determined by calculating their centre of gravity, after thresholding. The distance between the two images corresponding to the CH2 and CH3 peaks was also calculated to provide a reference. Both techniques described above can generate at least one image per second, as this rate was considered the minimum requirement to monitor magnetic field disturbances associated with traffic, trains and lift movement. Images can be acquired at a higher rate by reducing the data matrix size along the PE direction. When it was possible to control the potential source of magnetic field disturbance, this control was exercised (by controlling lift movement, for example). The images were processed with in-house software written in IDL (IDL 6.1; Research Systems Inc., Boulder, CO). Both techniques were validated prior to use in acceptance testing by bringing a strongly paramagnetic solution of dimeglumine gadopentetate (concentrated Magnevist, 0.5 mol l21; Schering, Berlin, Germany) closer to the measurement point and then removing it, creating a magnetic field disturbance. After validation of the imaging techniques, the stability of the magnetic field was investigated at two other 1.5 T MRI scanner sites as a part of acceptance testing in new installations (Philips Intera/Explorer gradients, Eindhoven, Netherlands and Siemens Symphony/Sprint gradients, Erlangen, Germany).

Results Single-shot echo-planar imaging (EPI) method

Sensitivity assessment

Magnetic field fluctuations cause image shifts along the phase encoding (PE) direction in a series of single-shot spin-echo EPI images. The magnitude of the image shifts depends on the PE bandwidth and, as a consequence, on

Phase images method Figure 1a shows the first phase image of the testobject, taken as a reference image, and Figure 1b shows


433

M A Schmidt

the difference between one of the phase images in the series and the reference image. The maximum phase change for any pixel over the object within the series of 100 images was under 0.25 ppm. Considering the

average phase over the small central ROI indicated (Figure 1b), the range of variation was up to 0.02 ppm (0.003 ppm standard deviation). This method is therefore not expected to be sensitive to field fluctuations below

Figure 1. (a) Reference phase image and (b) subtracted phase image of cylindrical uniform test object, with region of interest (ROI) used for measurements indicated. The variation of the average ROI phase value over a series of subtracted phase images is shown in (c), both in the absence of any magnetic field disturbance, and in the presence of a magnetic field disturbance caused by bringing a strongly paramagnetic solution to the test object vicinity and then removing it.

434


Short communication: Assessment of disturbances to the magnetic field in MR installations

0.02 ppm. When a strongly paramagnetic solution is brought closer to the test object, the ROI phase value changes as shown in Figure 1c.

EPI method Figure 2a–g shows sections of several single-shot EPI images of the hexane sample. CH2 and CH3 peaks produce separate images of the sample tube, of similar intensities. The image shift along the phase encoding direction can be seen from Figure 2d–f, as those were acquired when a strongly paramagnetic solution was brought close to the hexane sample, and then removed. In the absence of any magnetic field fluctuations, the position of the geometric centre of each tube image changes by no more than 0.005 ppm (Figure 2h). The standard deviation is as low as 0.0015 ppm. Figure 2i shows how the position of the tube centre changes, for both CH2 and CH3, when the magnetic field is disturbed by moving a tube containing a strongly paramagnetic solution in the vicinity. Despite some progressive blurring associated with the central frequency offset, the distance between the two peaks is kept constant in the range of the measurement. Therefore the EPI method is expected to be sensitive to magnetic field variations of the order of 0.01 ppm, as required for acceptance testing.

Acceptance testing The EPI method with the hexane sample was employed in two separate acceptance tests where the site conditions were to be evaluated. Both sites were equipped to perform functional brain studies with single shot-EPI, and one of them was equipped with a spectroscopy package. Magnetic field fluctuations of approximately 0.04 ppm were detected at the magnet isocentre during both tests. At one site the variation was associated with the movement of heavy equipment (image intensifier) along a shielded wall in the corridor adjacent to the MRI scanner room. A sign was placed to prevent that area from being used in a way that could disturb MRI data acquisition. At the other site, the variation was associated with car parking on the pavement close to the external wall on the side of the magnet room. The erection of a barrier was suggested.

Discussion and conclusions The single-shot EPI technique is the most sensitive of the imaging techniques discussed as it is expected to detect magnetic field fluctuations down to 0.005 ppm. However, single-shot EPI may not be available in every system, and it may not always be possible to decrease the PE bandwidth to the same level of sensitivity used in testing on the GE scanner as those parameters are clearly undesirable for clinical applications. The toxicity of hexane must also be mentioned, as it needs to be transported to the site of the acceptance test with due care. The ‘‘phase images’’ method is less sensitive but more easily available, and deserves further investigation and The British Journal of Radiology, May 2006

Figure 2. (a–g) Sections of several single-shot echo-planar imaging (EPI) images of hexane, showing the image shift along the phase encoding (PE) direction associated with a magnetic field disturbance. (d–f) A paramagnetic solution is brought to the vicinity of the test object, and (g) then removed. In the absence of any magnetic field disturbances any measured image shifts are under 0.005 ppm, as shown in (h) for the image associated with CH3. (i) The separation between the CH3 and CH2 images remains constant even when there is movement of a paramagnetic solution in the vicinity of the isocentre. 435

M A Schmidt

optimization. It could be made more sensitive with a higher SNR and using longer echo-times. If higher sensitivity is achieved, the method could be used in large uniform test objects (spherical or cylindrical) to detect disturbance to the magnetic field homogeneity over larger volumes. Both imaging techniques discussed could be further optimized, trading spatial resolution for a higher time resolution. However, standard MRI pulse sequences often impose restrictions to the minimum size of the data acquisition matrix. The interpretation of the results obtained with both methods assumes that an actual drift of the Larmor frequency is the only cause of the measured phase shifts and offsets. However, these effects can easily be caused by a number of different sources related to the scanner hardware (electronic components warming up, for example). For a scanner operating within specification, hardware-related phase drifts and offsets should be of very small magnitude, if detectable at all. Because of the possibility of simple hardware malfunction, it is essential to associate the measurements obtained with specific environmental sources and to perform reproducible measurements.

436

Both techniques discussed allow MRI users to evaluate the environmental conditions surrounding an MRI scanner, and can provide the basis for informed discussion on the management of the area surrounding it. The final decision on how much disturbance to the static magnetic field can be tolerated is complex and depends not only on the nature of the work to be performed in a particular scanner, but also on the pattern of the local magnetic field disturbance and its associated frequency. The single-shot EPI method, in particular, was proven to be highly sensitive and is strongly recommended for acceptance testing.

Acknowledgment The author would like to thank Dr Franklyn Howe for very helpful discussions on MR spectroscopy.

References 1. Durand E, van de Moortele P-F, Pachot-Clouard M, Le Bihan D. Artifact due to Bo fluctuations in fMRI: correction using the k-space central line. Magn Reson Med 2001;46:198–201.



SHORT COMMUNICATION

An investigation of search pattern extent in the threshold contrast detection task C J KOTRE,

PhD

Regional Medical Physics Department, Newcastle General Hospital, Newcastle-upon-Tyne NE4 6BE, UK ABSTRACT. The threshold contrast-detail diameter test is used as a semi-quantitative measure of image quality in radiology. This observation task is called ‘‘signal known exactly/background known exactly’’ because the signals are usually low contrast disks in known positions, and the background is uniform except for noise fluctuations. The performance of the observer undertaking this task can to some extent be predicted from knowledge of the noise power in the image background, and adoption of the assumption that the noise is sampled through an aperture of the same area as the test feature being observed. In order to extend this approach to optimization of clinical images, the effect of the cluttered anatomical background on the detection task must be quantified. To study the effect on detection of nearby structure, a series of contrastdetail tests was carried out using a progressively restricted background area of Gaussian noise, and a range of object diameters. It was found that the observer’s ability to detect low contrast objects is progressively reduced as the area of the search area is reduced, the difficulty of the task increasing rapidly as the diameter of the restricted search area falls to less than twice that of the target disk. The results suggest the presence of a search pattern that scales in proportion with the size of the test feature.

The threshold contrast-detail diameter test has been used for many years as a standard semi-quantitative test of image quality especially in fluoroscopy, radiography and mammography [1–3]. In this technique, the observer is shown a set of disks imaged against a uniform background. The disks vary both in contrast and diameter and are laid out in a fixed pattern within the test object. This type of stimulus is known as a ‘‘signal known exactly/background known exactly’’ (SKE/BKE) task. For each disk diameter the observer works along a sequence of disks of decreasing contrast and decides which is the last visible disk. The contrast of the last visible disk is defined as the threshold contrast at that disk diameter. The process is repeated over a range of disk diameters and the results can be plotted as a contrast-detail curve, which marks the transition from what combination of size and contrast can be seen in the image and what cannot be seen because it is masked by noise. The most widely accepted model of the behaviour of the human visual system in undertaking the SKE/BKE contrast-detail task is that the threshold contrast obtained is equal to the noise observed in the background, multiplied by a constant, the threshold signal-tonoise ratio, the value of which is thought to be around 2.5 [4]. The noise perceived by the observer is taken to be the standard deviation of a number of luminance samples of the background, where each sample is the mean luminance within an aperture of the same area as the disk under observation [5, 6]. This model broadly fits The British Journal of Radiology, May 2006

Received 7 September 2005 Revised 2 November 2005 Accepted 14 November 2005 DOI: 10.1259/bjr/13489819 ’ 2006 The British Institute of Radiology

experimental results when modified to take account of further limitations to detection at small disk sizes due to the finite point spread function of the eye, and at large disk sizes due to the maximum extent of the photoreceptor field over which summation of the noise sample may occur [4, 6]. Although the SKE/BKE contrast-detail task has been shown to be very useful in comparing the performance of radiological image receptors, where the noise in the uniform background is principally that from X-ray quantum fluctuations plus components of system noise, such as electronic amplification noise and the fixed noise pattern of phosphor screens, there are difficulties in extending the results to the prediction of detectability in clinical images. The background in clinical images consists of a wide variety of object shapes, sizes and contrasts all of which tend to inhibit the detection of the diagnostic signal required in any given examination. The presence of the background can be thought of as a noise source, sometimes termed structure noise or clutter, and it is the removal of this noise component that allows small contrast signals to be detected in digital subtraction angiography. Although meaningful results can be obtained in observational studies of contrast-detail test objects overlying a complex anatomical type background [7], it is difficult to extend this to the general problem of dose optimization of clinical images. In order to undertake optimization it will be necessary to be able to quantify background structure as a noise source, and to be able to predict its effect on observer performance. One 437

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means of achieving this may be to design algorithms to sample the radiological image in a similar way to a human observer, so that the detectability of a low contrast object can be analysed on a point-by-point basis against the background of real clinical images. Although the present work is based on the contrast-detail test, other approaches to quantifying signal detection in realistic radiological backgrounds have been explored, notably alternative forced choice methods [8–10] and receiver operating characteristic studies [11, 12]. As a starting point for the design of algorithms to mimic an observer undertaking a contrast-detail task, it is necessary to extend the knowledge of the human

(a)

observer noise sampling to investigate not only the area of the sampling aperture, but the spatial distribution of the samples and the effect on observer performance of the proximity of adjacent structures.

Experimental observations A series of static contrast-detail test patterns were presented to six experienced observers on a monochrome TV display monitor. The display was first calibrated in terms of luminance using a Hagner S2 photometer (Hagner AB, Solna, Sweden) whose response is matched

(b)

(c) Figure 1. Examples of test images. (a) The 16 pixel diameter disks with no search restriction (other than the limits of the image frame). (b,c) The same diameter with Ds/Dd53 and 1.5, respectively. 438


Short communication: Search pattern extent for threshold contrast tests

to that of the human eye, and the monitor settings were then left fixed for the duration of the experiment. The photometer was calibrated using a source of known luminous intensity in terms of illuminance, from which the luminance calibration of its internal detector was derived using the relationship between luminance of a surface and illuminance at a point for a Lambert source given by Guibelalde et al [13]. Each test pattern consisted of a 464 matrix of 16 circular test features decreasing in contrast in steps of 0.84 (reciprocal cube-root of 2). Computer-generated Gaussian noise was added to the whole image sufficient to produce a threshold cut-off within the contrast range displayed. The test images consisted of 5126512 8-bit pixels displayed in a 20 cm616 cm format at a viewing distance of approximately 50 cm, although the observer was not restricted in viewing distance. The observations were carried out with room lights dimmed, and observers were allowed to score disks as ‘‘half seen’’ where detection was uncertain, following common practice in scoring contrast-detail tests in radiological image assessment. Four detail diameters were used (32, 16, 8 and 4 pixels, equivalent to 11.3 mm, 5.6 mm, 2.8 mm and 1.4 mm, respectively) and each of these was displayed within a larger circular background area of mid-grey with noise added to both disk and background. Outside the area of mid-grey, the screen was blanked to black with no noise to limit the search area of the observer. The combinations of disk diameter (Dd) and search area diameter (Ds) were varied to give no search restriction, then Ds/Dd54, 3, 2, 1.5, 1.25 and 1.125 for each of the four disk diameters. The test patterns were presented in random order with no time restriction on the observation. Three example test patterns are reproduced in Figure 1. Figure 1a shows the 16 pixel diameter disks with no search restriction (other than the limits of the image frame), Figure 1b,c shows the same diameter with Ds/Dd53 and 1.5, respectively.

Figure 2. Contrast-detail plots of the results for the unrestricted search (#), and Ds/Dd53 (%), 1.5 (n) and 1.125 (,). The error bar illustrates a typical¡1 standard error on the mean value of threshold contrast for the experiment.


Figure 3. The ratio of threshold contrast for the limited search, normalized to that of the unlimited search, plotted against Ds/Dd, for all four disk diameters: 1.4 mm ( ), 2.8 mm (&), 5.6 mm (m) and 11.3 mm (.).

N

At the conclusion of all observations, the recorded scores were converted to luminance threshold contrasts (CT) and averaged across observers.

Results Figure 2 shows conventional contrast-detail plots of some of the results with quadratic curve fits on logarithmic axes. To avoid obscuring the figure, only plots for the unrestricted search and Ds/Dd53, 1.5 and 1.125 are shown. As Ds/Dd is reduced, the threshold contrasts increase as expected. The error bar illustrates a typical¡1 standard error of the mean value of threshold contrast for the experiment. Standard errors for each point were calculated, but are not shown to avoid obscuring the figure. Figure 3 shows the ratio of threshold contrast for the restricted search, normalized to that of the unrestricted search, plotted against Ds/Dd, for all four disk diameters. Although the values are subject to considerable uncertainties, the trend with Ds/Dd appears to be similar for all diameters, suggesting the presence of a search pattern that scales directly with the size of the object being observed. Taking the similarity of the results of Figure 3 as sufficient to support the hypothesis that the search pattern scales in proportion to the size of the object, Figure 4 shows the data from Figure 3 averaged across all disk diameters. It is notable that the mean ratio of threshold contrasts already has value greater than unity (although unity is within the error bar) at Ds/Dd54, indicating that some interference with the observer’s search pattern is already present with this background. The ratio increases quickly for values of Ds/Dd , 2, as the visual task changes from finding a disk on a background to detecting the difference between the intensity of the disk and the narrow band of noisy background surrounding it. This finding is in agreement 439

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systems, and also to the broader investigation of the effect of background anatomical structure on the detection of diagnostic features in radiology.

References

Figure 4. The ratio of threshold contrast for the limited search, normalized to that of the unlimited search, plotted against Ds/Dd for the data from Figure 3 averaged across all disk diameters. The error bars show¡1 standard error on the mean.

with previous results observed for a design of mammographic contrast-detail phantom, where a marker circle at 2.7 times the diameter of the test feature was found to reduce detection, but marker circles at greater relative diameters (.6) had little effect [14].

Conclusions It had been shown that for the SKE/BKE contrastdetail detection task undertaken against a background of Gaussian noise, the observer’s ability to detect low contrast objects is progressively reduced as the area of the search area around the object is reduced, the difficulty of the task increasing rapidly as the diameter of the restricted search area falls to less than twice that of the target disk. The results further suggest the presence of a search pattern that scales in proportion with the size of the test feature. These results are relevant to the design of contrast-detail test phantoms for radiological imaging

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1. Hay GA, Clarke OF, Coleman NJ, Cowen AR. A set of X-ray test objects for quality control in television fluoroscopy. Br J Radiol 1985;58:335–44. 2. Bijkerk KR, Lindeijer JM, Thijssen MAO. The CDMAM phantom: a contrast-detail phantom specifically for mammography. Radiology 1993;185(P):395. 3. Institute of Physics and Engineering in Medicine. Recommended standards for the routine performance testing of diagnostic X-ray systems. (IPEM Report 77) York: IPEM, 1997. 4. Schnitzler AD. Image-detector model and parameters of the human visual system. J Opt Soc Am 1973;63:1357–68. 5. Rose A. The sensitivity performance of the human eye on an absolute scale. J Opt Soc Am 1948;38:196–208. 6. Chesters MS, Hay GA. Quantitative relation between detectability and noise power. Phys Med Biol 1983;28: 1113–25. 7. Kotre CJ. The effect of background structure on the detection of low contrast objects in mammography. Br J Radiol 1998;71:1162–7. 8. Ohara K, Doi K, Metz CE, Giger ML. Investigation of basic imaging properties in digital radiography. 13. Effect of simple structured noise on the detectability of simulated stenotic lesions. Med Phys 1989;16:14–21. 9. Bochud F, Valley J-F, Verdun FR. Estimation of the noisy component of anatomical backgrounds. Med Phys 1999;26: 1365–70. 10. Burgess AE, Jacobson FL, Judy PF. Human observer detection experiments with mammograms and power-law noise. Med Phys 2001;28:419–37. 11. Kundel HL, Nodine CF, Thickman D, Carmody D, Toto L. Nodule detection with and without a chest image. Invest Radiol 1985;20:94–9. 12. Reverez G. Conspicuity and uncertainty in the radiographic detection of lesions. Radiology 1985;154:625–8. 13. Guibelalde E, Vano E, Llorca AL. Quality assurance of viewing boxes: proposal for establishing minimum requirements and results from a Spanish quality control programme. Br J Radiol 1990;63:564–7. 14. Kotre CJ, Porter DJT. A printed image quality test phantom for mammography. Br J Radiol 2005;78:746–8.



SHORT COMMUNICATION

Imaging pulmonary embolism in pregnancy: what is the most appropriate imaging protocol? S MATTHEWS,

MBBS, MRCP, FRCR

Radiology Department, Northern General Hospital, Herries Road, Sheffield, South Yorkshire S5 7AU, UK ABSTRACT. Pulmonary embolism is the leading cause of death in pregnancy. Despite the difficulties in clinical diagnosis and the concerns regarding radiation of the fetus, the British Thoracic Society guidelines for imaging pulmonary embolism do not specifically address the issue of imaging for pulmonary embolism in this group. This communication discusses the difficulties of diagnosis and imaging pulmonary embolism in pregnancy and proposes a suitable imaging protocol. Clinical exclusion of patients from further imaging is recommended if the patient has a low pre-test probability of pulmonary embolism and a normal d-dimer. It is advised that all remaining patients undergo bilateral leg Doppler assessment. If this test is positive, the patient should be treated for pulmonary embolism; if negative, all patients should be referred for CT pulmonary angiography. Ideally, informed consent should be obtained prior to CT scanning. All neonates exposed to iodinated contrast in utero should have their thyroid function tested in the first week of life due to the theoretical risk of contrast induced hypothyroidism.

Pulmonary embolism (PE) is the leading cause of maternal death [1]. The rate of PE in pregnancy is five times greater than that for non-pregnant women of the same age and is about 1 in 1500 deliveries; the risks are even higher in the puerperium. The clinical diagnosis of PE is difficult in the general population, but is further complicated in pregnancy as some of the clinical symptoms of PE can be normal/expected symptoms of pregnancy. Precise PE diagnosis in pregnancy is vital to prevent unnecessary treatment of PE as treatment is associated with side effects for both the mother and fetus. Accurate imaging is essential, but there are frequently anxieties relating to fetal radiation exposure during diagnostic procedures. Despite the complexities of the clinical scenario, the British Thoracic Society (BTS) guidelines on the management of PE do not address the issue of imaging for PE in the pregnant patient [2]. There are an array of clinical, biochemical and radiological tests available for the investigation of PE, some of which have not been validated in pregnancy. In general terms, the BTS guidelines recommend clinical assessment of the pre-test probability of PE for each patient with d-dimer assessment for patients with a low or intermediate pre-test probability followed by either a ventilation/perfusion scan or CT pulmonary angiogram (CTPA; depending on the availability of nuclear medicine scans locally and the presence of a chest X-ray abnormality). The Wells criteria are the most frequently used tool for assessing the clinical probability of PE. However, pregnant patients were excluded from the analysis group for validation of the criteria [3]. The d-dimer is known to increase in pregnancy. The d-dimer The British Journal of Radiology, May 2006

Received 1 August 2005 Revised 17 November 2005 Accepted 19 January 2006 DOI: 10.1259/bjr/15144573 ’ 2006 The British Institute of Radiology

is usually normal in the first trimester of uncomplicated pregnancy, starts to rise during the second trimester and returns to baseline levels at 4–6 weeks post partum [4–6]. The d-dimer is not affected by bleeding, breast feeding or heparin prophylaxis, but is elevated in association with many pregnancy-related complications such as pre-eclampsia [5, 7]. Pregnant women with ongoing thrombosis have been shown to develop a significant rise in d-dimer [8]. It is estimated that 70% of patients with a proven PE have proximal deep venous thrombosis (DVT). However, the proportion in pregnant women with PE is unknown. Therefore, Doppler ultrasound of the leg veins is recommended for the investigation of PE if the patient has symptoms and signs suggestive of DVT [2]. The diagnosis of DVT in pregnant women can be problematic. There is increased lower extremity vein diameter and decreased flow secondary to hormonal effects and the direct compressive effect of the enlarged uterus on pelvic veins, hence the legs are frequently swollen in the absence of DVT [9]. There is an increased risk of iliac vein thrombosis, which is not routinely assessed by leg Doppler studies. In addition, the accuracy of Doppler ultrasound, including iliac vein assessment, has not been validated in pregnant patients. However, Doppler ultrasound does not involve ionizing radiation and, as there are concerns about the radiation exposure to mother and fetus, bilateral leg Doppler has been proposed as the initial investigation of suspected PE in pregnancy [9]. This would conform to the general principle of maintaining doses ‘‘as low as reasonably achievable’’ whilst still offering a valuable test that may 441

S Matthews

preclude the need for further investigations associated with a significant radiation dose [10]. Ventilation/perfusion (V/Q) imaging is well established for imaging PE and, in a survey relating to imaging practice for the investigation of PE in pregnancy in 1998, was the most frequently employed test for this sub group [11]. In pregnant women the radiation dose can be minimized by using a half-dose perfusion scan and only proceeding to ventilation imaging if a defect is identified on the perfusion scan [12]. However, for the general population 50–70% of V/Q scans are indeterminate. For pregnant women the proportion of patients falling into each of the reporting categories (high probability, normal and indeterminate) is different: fewer pregnant women have high probability scans (less than 5%) and many more have a normal scan (75%) [13– 15]. Therefore, only 20% of this patient group have indeterminate scans. This change in distribution of patients within the probability groups is thought to reflect the younger average age and reduced presence of co-morbidities compared with the general population. PE can be confidently excluded with a normal V/Q scan, but this test throws up a relatively high proportion of indeterminate results and, in the high probability group, up to 20% of the patients may not have PE. A small study involving 113 de novo cases of potential PE showed that withholding anticoagulation in pregnant women with normal or indeterminate scans is probably safe [13]. However, larger studies are required to confirm this finding, especially as the incidence of PE in the proportion of pregnant women investigated is so low. It has been suggested that CTPA has a greater discriminatory power than V/Q scanning with a low pre-test probability, but that CTPA and V/Q scanning have a similar discriminatory power in those with a high pre-test probability [16]. CTPA is now a well-validated investigation with a sensitivity and specificity between 94% and 100% [17, 18]. The negative predictive value of a normal CTPA is over 99% [19]. The clinical validity of a CT scan to rule out PE is similar to that reported for conventional pulmonary angiography [19]. Anticoagulants can be safely withheld if the CTPA is negative for PE. CTPA is advantageous as the emboli are directly visualized (unlike for V/Q scanning) and alternative causes for the patient’s symptoms may be diagnosed. However, there are concerns regarding the radiation dose received from CT scanning, particularly to the fetus. Recent studies have shown that the fetal radiation exposure for CTPA varies from 3.3 mGy to 130.0 mGy; the dose increasing during each trimester as the fetus enlarges and approaches the imaged area in the thorax [20]. However, the estimated fetal radiation dose for V/Q scanning is estimated as 100–370 mGy, i.e. the dose may be more than 3 times greater than for CTPA. In addition, CTPA has a superior sensitivity and specificity for PE [21]. All radiation to the fetus carries a potential risk. This risk must be balanced against the risk to the mother/ fetus if PE is not diagnosed or treated and against the risk of treatment of non-confirmed PE. Everyone is exposed to radiation all the time from the atmosphere, ground and from ingested food and drink. The average ‘‘background radiation’’ for an individual in the UK is 442

2.7 mSv per year, which equates to about 1000 mGy for a fetus in utero for 9 months [22]. The worst estimated absorbed dose for the fetus in the third trimester undergoing CT pulmonary angiography is 130 mGy, i.e. approximately 7 times less than the natural background radiation. Central nervous system malformations can be associated with excess radiation [23]. A threshold fetal dose greater than 100 000–200 000 mGy is required to cause such a problem. A dose of 100 mGy to the fetus is associated with an excess death from cancer up to the age of 15 years of 1 in 300 000. This can be put into perspective by comparison with everyday activities and their relative annual risk of death in the UK, e.g. smoking 10 cigarettes each day has a risk of 1 in 200 and relative risk of death from uncomplicated pregnancy for the mother is 1 in 170 000 [24]. The annual risk of death in the UK for all cancer is 1 in 400 and for death from all causes at the age of 40 years 1 in 700 [24]. Therapeutic termination would not be considered for a fetal dose below 100 000 mGy. However, it is recommended that the patient is given more detailed information regarding the risks of radiation for procedures where the fetal dose is expected to be greater than 1000 mGy [23]. Therefore, although the radiation risk to the fetus cannot be ignored, the risk is very low. The risk of fetal death is much greater if the mother has untreated PE. In addition to the radiation risk to the fetus, the breast radiation dose from CTPA must be considered. The female breast is extremely radiosensitive, and it has been shown that a sufficiently large radiation dose can cause breast cancer [25]. The exposure of the immature breast during early development and around the time of menarche carries a higher risk than at other times of a woman’s life. There is little evidence that radiation exposure after 45 years of age increases the incidence of breast cancer. The effect of radiation on the breast in pregnancy is unclear. However, there has been shown to be a relatively linear relationship between radiation dose and subsequent breast cancer, although this relationship does not extend into the highest radiation dose exposures used for radiotherapy. The vast majority of people exposed to radiation do not develop a cancer related to that exposure. The 25 000 female atomic bomb survivors in Japan have been followed for over 50 years, but only 173 breast cancer deaths have occurred of which 41 were attributed to the radiation received in 1945. A radiation dose of 100 cGy is associated with an increased risk of breast cancer of 40% in young Western women [25]. This is the same risk a woman experiences secondary to several common conditions/lifestyle choices, e.g. never being pregnant, menarche before the age of 11 years or a late menopause. Epidemiological studies have not detected a significantly increased risk of breast cancer below a dose of 20 cGy. Parker et al have specifically investigated female breast radiation exposure during CTPA and calculated an effective minimum dose of 20 mGy (2 cGy) [26]. This dose concurs with other data estimating a dose between 20 mGy and 50 mGy (2–5 cGy) for a standard chest CT scan [27]. These estimates are significantly below the level of 20 cGy, below which no effect on the breast can be demonstrated, but significantly higher than the estimated breast radiation dose of 0.28 mGy associated with ventilation/perfusion scanning [28]. Although this The British Journal of Radiology, May 2006

Short communication: Imaging pulmonary embolism in pregnancy

radiation exposure is associated with an immeasurable low malignancy risk, this exposure should not be ignored and the development of breast shields that may reduce this dose by up to 73% may be considered in the future [29]. A 2003 survey of members of the Society of Thoracic Radiology found that 53% of responding radiologists would use CTPA as a first line investigation for excluding PE in pregnant patients, but only 60% of radiologists obtained informed consent from any pregnant patient undergoing CTPA, only 16% of departments had a written policy for this group and only 40% modified the imaging protocol in an attempt to reduce radiation dose [30]. The risks of iodinated contrast media on the fetus have not been fully investigated. Many pregnant patients have received intravenous contrast for the investigation of other problems such as renal colic. However, there are no reports in the literature of any ill effects, despite the theoretical risk of contrast induced hypothyroidism. In summary, d-dimer concentration is known to rise in normal pregnancy, but a normal value is clinically helpful. Doppler ultrasound of both legs has a low pick up rate, but does not involve ionizing radiation. If DVT is found, the patient should be treated as for PE. This test requires further validation in pregnant women. Ventilation perfusion imaging is associated with a higher radiation dose to the fetus, but lower radiation dose to the mother/breast than CTPA. The majority of patients have a low or indeterminate probability of PE from the V/Q scan and it is probably safe to withhold anticoagulants in these groups, but this has not been validated in a large controlled study. V/Q scanning has a lower discriminatory threshold for PE than CTPA in patients with a low or intermediate pre-test clinical probability. The mother receives a higher radiation dose from CTPA, especially to the breast, but the fetal dose is lower than V/Q imaging. However, all quoted radiation doses to breast or fetus are below the thresholds estimated to be associated with any significant risk. CTPA has a greater sensitivity and specificity than V/Q scanning and is able to diagnose alternative causes for the patients’ symptoms in cases where PE is absent. However, the risk of iodinated contrast to the fetus is not known. Multiple clinicians may interpret these same facts and derive differing imaging protocols based on their perceived significance for each factor discussed. The following imaging protocol is proposed to optimize the accuracy of PE diagnosis whilst minimizing radiation dose to the fetus: All patients should be clinically assessed by a senior clinician and a pre-test probability assigned. Patients in the first two trimesters of pregnancy should have the d-dimer measured, provided there is no specific contraindication. A normal d-dimer and low pretest probability of PE can be used to exclude PE in the pregnant patient. Exclusion of patients from further imaging on the basis of intermediate pre-test probability and normal d-dimer needs further investigation. All pregnant patients requiring further investigation for PE should proceed to bilateral leg Doppler ultrasound. If this test is positive for venous thromboembolism, the patient should be treated for PE. If this test is negative, all patients should proceed to CTPA. Informed consent The British Journal of Radiology, May 2006

should be sought from the pregnant patient prior to CTPA. The patient should be given simple information explaining the risks of fetal and maternal radiation and risks to the mother and child of failing to accurately diagnose PE. The CTPA protocol should be modified to minimize the radiation dose; in particular the length of the thorax along the z-axis should be reduced. It is recommended that the baby has thyroid function testing within the week of birth due to the theoretical risk of contrast induced hypothyroidism [31]. Further research is advised to validate the above protocol and to clarify the in utero affects of iodinated contrast on the neonate.

References 1. Pabinger I, Grafenhofer H. Thrombosis during pregnancy: risk factors, diagnosis and treatment. Pathophysiol Haemost Thromb 2002;32:322–4. 2. British Thoracic Society guidelines for the management of suspected acute pulmonary embolism. Thorax 2003;58: 470–84. 3. Wells PS, Anderson DR, Rodger M, Stiell I, Dreyer JF, Barnes D, et al. Excluding pulmonary embolism at the bedside without diagnostic imaging: management of patients with suspected pulmonary embolism presenting to the emergency department by using a simple clinical model and d-dimer. Ann Int Med 2001;135:98–107. 4. Kline JA, Williams GW, Hernandez-Nino J. D-dimer concentrations in normal pregnancy: new diagnostic thresholds needed. Clin Chem 2005;51:825–9. 5. Epiney M, Boehlen F, Boulvain M, Reber G, Antonelli E, Morales M, et al. D-dimer levels during delivery and the postpartum. J Thromb Haemost 2005;3:268–71. 6. Frascalanci I, Comeglio P, Liotta AA, Cellai AP, Fedi S, Parretti E, et al. D-dimer concentrations during normal pregnancy, as measured by ELISA. Thromb Res 1995;78: 399–405. 7. Nolan TE, Smith RP, Devoe LD. Maternal plasma D-dimer levels in normal and complicated pregnancies. Obstet Gynecol 1993;81:235–8. 8. Bombelli T, Raddatz-Mueller P, Fehr J. Coagulation activation markers do not correlate with the clinical risk of thrombosis in pregnant women. Am J Obstet Gynecol 2001;184:382–9. 9. Chan W-S, Ginsberg JS. Diagnosis of deep vein thrombosis and pulmonary embolism in pregnancy. Thrombosis Res 2002;107:85–91. 10. Statutory Instruments. Ionising Radiation (Medical Exposure) Regulations 2000. London: HMSO, 2000, No. 1059. 11. Boiselle PM, Reddy SS, Villas PA, Liu A, Seibyl JP. Pulmonary embolus in pregnant patients: survey of ventilation-perfusion imaging policies and practice. Radiology 1998;207:201–6. 12. Balan KK, Critchley M, Vedavathy KK, Smith ML, Vinjamuri S. The value of ventilation-perfusion imaging in pregnancy. Br J Radiol 1997;70:338–40. 13. Chan WS, Ray JG, Murray S, Coady GE, Coates G, Ginsberg JS. Suspected pulmonary embolism in pregnancy: clinical presentation, results of lung scanning, and subsequent maternal and pediatric outcomes. Arch Intern Med 2002;162:1170–5. 14. Hull RD, Hirsh J, Carter CJ, Raskob GE, Gill GJ, Jay RM, et al. Diagnostic value of ventilation-perfusion lung scanning in patients with suspected pulmonary embolism. Chest 1985;88:819–28. 15. Value of ventilation/perfusion scan in acute pulmonary embolism. Results of the prospective investigation of

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pulmonary embolism diagnosis (PIOPED). The PIOPED investigators. JAMA 1990;263:2753–9. Hayashino Y, Gotto M, Noguchi Y, Fukui T. Ventilationperfusion scanning and helical CT in suspected pulmonary embolism: Meta analysis of diagnostic performance. Radiology 2005;234:740–8. Blachere H, Latrabe V, Montaudon M, Valli N, Couffinhal T, Raherisson C, et al. Pulmonary embolism revealed on helical CT angiography: comparison with ventilation-perfusion radionuclide lung scanning. AJR Am J Roentgenol 2000;174:1041–7. Remy-Jardin M, Remy J, Baghaie F, Fribourg M, Artaud D, Duhamel A. Clinical value in the thin collimation diagnostic workup of pulmonary embolism. AJR Am J Roentgenol 2000;175:407–11. Quiroz R, Kucher N, Zou KH, Kipfmueller F, Costello P, Goldhaber SZ, et al. Clinical validity of a negative computed tomography scan in patients with suspected pulmonary embolism: a systematic review. JAMA 2005;293:2012–7. Winer-Muram HT, Boone JM, Brown HL, Jennings SG, Mabie WC, Lombardo GT. Pulmonary embolism in pregnant patients: fetal radiation dose with helical CT. Radiology 2002;224:487–92. Bates SM, Ginsberg JS. How we manage venous thromboembolism in pregnancy. Blood 2002;100:3470–8. Watson SJ, Jones AL, Oatway WB, Hughes JS. Health Protection Agency. Ionising radiation exposure of the UK population: 2005 review. HPA – RPD – 001.

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23. Pregnancy and medical radiation: International Commission on Radiological Protection. Ann ICRP 2000;30: 1–43. 24. Living with radiation. (Fifth edition). Chilton: NRPB, 1998. 25. Ionising radiation and breast cancer risk. Sprecher Institute for Comparative Cancer Research. Fact Sheet 52, October 2004. 26. Parker MS, Hui FK, Camacho MA, Chung JK, Broga DW, Sethi NN. Female breast radiation exposure during CT pulmonary angiography. AJR Am J Roentgenol 2005;185: 1228–33. 27. Rehani MM. CT: caution on radiation dose. Ind J Radiol Imag 2000;10:19–20. 28. Cook JV. UK agency reports slight increase in radiation exposure. BMJ 2005;330:1229. 29. Ocker J. New Breast Shield reduces radiation to female breast during chest CT by up to 73%. Medical News Today, 16 May 2005. 30. Schuster ME, Fishman, JE, Copeland JF, Hatabu H, Boiselle PM. Pulmonary embolism in pregnant patients: a survey of practices and policies for CT pulmonary angiography. AJR Am J Roentgenol 2003;181:1495–8. 31. Webb JA, Thomsen HS, Morcos SK, and members of the Contrast Media Safety Committee of European Society of Urogenital Radiology. The use of iodinated and gadolinium contrast media for MRI during pregnancy and lactation. Eur Rad 2005;15:1234–40.



CASE REPORT

Primary subcutaneous sacrococcygeal ependymoma: a case report and review of the literature Y T MA,

MRCP,

P RAMACHANDRA,

FRCR

and D SPOONER,

FRCR

The Cancer Centre, Queen Elizabeth Hospital, Edgbaston, Birmingham B15 2TH, UK ABSTRACT. Extraspinal ependymomas are rare. The majority occur in the sacrococcygeal region. The subcutaneous variety accounts for approximately two thirds of cases, which are commonly misdiagnosed as a pilonidal cyst or sinus. Treatment is complete surgical resection. The role of coccygectomy is controversial. Adjuvant radiotherapy is of benefit to those with an incompletely excised tumour. Up to 20% metastasise, chiefly to the inguinal lymph glands, but pulmonary metastases are also reported. Palliative chemotherapy has not been shown to be of any benefit. Long term follow-up is important as metastases can occur up to 20 years after initial presentation. We report a 37-year-old woman with a subcutaneous sacrococcygeal ependymoma with iliac lymph nodal metastasis at presentation.

The ependymal cells form an epithelial sheet and line the ventricles in the brain and the central canal of the spinal cord. Ependymomas are tumours derived from ependymal cells and are usually confined to the central nervous system, occurring predominantly in children and young adults. They account for only 6% of intracranial gliomas, but in the spinal cord they are the most common tumours of glial origin, accounting for over 60%, mostly in the region of the cauda equina and the filum terminale [1]. Rarely, ependymomas occur outside of the central nervous system [1]. The majority occur in the sacrococcygeal subcutaneous tissue or the presacral regions [2– 10]. We present a case of an ependymoma arising from the sacrococcygeal subcutaneous tissue with iliac lymph nodal metastasis at presentation.

Received 5 May 2004 Revised 13 May 2005 Accepted 19 May 2005 DOI: 10.1259/bjr/61959899 ’ 2006 The British Institute of Radiology

pelvis demonstrated a residual lobulated soft tissue mass, about 3 cm in diameter, lying anterior and to the right of the tip of the coccyx. A second smaller lesion was also visible between the right internal and external iliac arteries (Figure 1). CT scanning of the thorax and abdomen demonstrated multiple liver cysts, but no evidence of lung or liver metastases. Isotope bone scanning revealed no evidence of skeletal metastases.

Case report A 37-year-old Caucasian lady presented with a 6month history of an enlarging nodule at the base of her coccyx. She had always been aware of the presence of a small nodule but had thought this was ‘‘normal’’ for her. During the 6-month period the nodule had increased in size from less than 0.5 cm to over 3 cm in diameter. There was no history of urinary or faecal problems. She was seen at her local hospital and a diagnosis of a pilonidal sinus was made which was subsequently excised. Histological examination revealed this to be a subcutaneous myxopapillary sacrococcygeal ependymoma and she was referred to the regional centre for musculoskeletal oncology. The full blood count, biochemical profile and tumour markers (alpha-fetoprotein (AFP), ß-human chorionic gonadotrophin (HCG), carcinoembryonic antigen (CEA) and CA-125) were normal. MRI of the abdomen and The British Journal of Radiology, May 2006

Figure 1. MRI abdomen and pelvis. Sagittal T2 short tau inversion recovery (STIR) sequence showing a 3 cm lobulated soft tissue mass lying anterior to the tip of the coccyx, and a second small rounded soft tissue mass superiorly in the pelvis.

445

Y T Ma, P Ramachandra and D Spooner

The residual mass was completely excised. Intraoperatively, it was seen near the tip of the coccyx but separate from it. Additionally, a 1 cm solid lesion with cystic elements was found above the right common iliac artery. This was thought to be a right iliac lymph node and was also completely removed. The histology of the specimens again confirmed a myxopapillary ependymoma (Figure 2). The majority of the tumour cells were positive for glial fibrillary acidic protein (GFAP) and S-100, but epithelial and other neural and neuroendocrine markers were negative. The resection margins were free of tumour cells and the smaller lesion was confirmed to be a nodal metastasis. She made a good post-operative recovery. No adjuvant therapy was offered, as there was no evidence of either macroscopic or microscopic disease. Re-staging CT scanning of the thorax, abdomen and pelvis at 3 months showed no residual or metastatic disease. 6 months following surgery she developed persistent right iliac fossa pain. CT scanning of the abdomen and pelvis now demonstrated a soft tissue mass 4.5 cm 6 4 cm 6 3.5 cm in the right pelvis and also the presence of an enlarged right internal iliac lymph node (1.5 cm), suggestive of recurrent disease. The right pelvic mass and iliac lymph node were completely excised. Histological examination of the pelvic mass revealed reactive tissue only. Histological examination of the right iliac lymph node confirmed metastatic myxopapillary ependymoma. Surgery was followed with adjuvant radiotherapy to the right pelvis (45 G in 25 daily fractions over 5 weeks). She has made a good recovery and is now being closely followed up in clinic.

Discussion Primary extraspinal ependymomas are rare and usually occur in the subcutaneous sacrococcygeal or presacral regions. Mallory reported the first case in 1902 [5]. Since then there have been approximately 75 reported cases of extraspinal ependymomas, with the subcutaneous variety accounting for about 50 cases (Table 1) [2–10]. Most of the published cases of sacrococcygeal ependymomas have occurred in children and young adults, although there have been reported cases ranging from 2 months to 67 years of age [2–10]. The male to female ratio of affected individuals is roughly equal [2, 6]. Subcutaneous sacrococcygeal ependymomas are thought to arise from the coccygeal medullary vestige, an ependymal lined cavity forming the remnant of the caudal portion of the neural tube [11]. This theory is supported by Bale who found post-coccygeal ependymal vestiges in 10 of 15 random necropsies in infants [12].

Presacral ependymomas are thought to arise either from the extradural remnants of the filum terminale or as an extension from the intradural filum terminale [2, 3]. The frequent finding of sacral bone destruction and sacral nerve dysfunction in patients with presacral ependymomas supports this theory [4]. A germ cell origin has also been proposed which may explain the rare occurrence of ependymomas in the ovary and the mediastinum [2]. Most subcutaneous sacrococcygeal ependymomas are of the myxopapillary variety [7]. The other histological types, according to the World Health Organization (WHO) classification, are papillary and subependymal [2]. Characteristically, subcutaneous sacrococcygeal ependymomas grow slowly and often reach a relatively large size prior to presentation. The majority of patients are asymptomatic. Consequently, these tumours are frequently misdiagnosed clinically as a pilonidal cyst or sinus [6, 8]. This was a feature in our case. Ependymomas confined to the central nervous system rarely metastasise outside of the cerebrospinal fluid (CSF) pathway. Conversely, up to 20% of subcutaneous sacrococcygeal ependymomas have been reported to develop distant metastases [2, 6, 8]. This has often been attributed to the easier access of these tumours to lymphatic and vascular vessels. Pulmonary, pleural, osseous, inguinal nodal and subcutaneous metastases have been documented [2–10]. Our patient was found to have a synchronous right iliac lymph nodal metastasis, which is an unusual finding. Currently, no reliable clinical or histological feature exists to help predict the development of metastases. Distant metastases can occur 10–20 years after initial presentation [6]. Consequently, patients need to be under regular follow-up for many years. Immunohistochemical staining helps to make the diagnosis. Typically, there are positive reactions to glial fibrillary acidic protein (GFAP), S-100 and vimentin [2]. This was seen in our case; the tissue reacted positively to GFAP and S-100. The treatment of subcutaneous sacrococcygeal ependymomas remains controversial because of their rarity. The favoured treatment is surgical excision. Complete excision appears to be essential to prevent local recurrence. Surgical removal of the coccyx may also be required if the tumour is attached to this bone [2]. Some authors have suggested routine coccygectomy as part of the primary surgical treatment, because of an apparent reduction in the recurrence rate. Aktug˘ et al carried out a review of all reported cases of sacrococcygeal extraspinal ependymomas [9]. They found that local recurrence did not occur if the coccyx was removed (four patients, mean follow-up period of 3 years). Conversely, if the coccyx was not excised, local recurrence occurred in 71% of patients (5 out of 7 patients). It appears that the likely

Table 1. Comparison of subcutaneous and presacral sacrococcygeal ependymomas Extraspinal ependymomas

No. of reported cases

Male:female ratio

Origin

Symptoms

Subcutaneous Presacral

49 28

Equal Equal

Coccygeal medullary vestige Extradural remnants of the filum terminale, or Extension from the intradural filum terminale

Asymptomatic mass Neurological symptoms (Bowel/bladder dysfunction; saddle anaesthesia; paraesthesia and weakness of lower limbs)

446


Case report: Primary subcutaneous sacrococcygeal ependymoma

Although extremely rare, it is important to consider subcutaneous sacrococcygeal ependymomas in the differential diagnosis of a sacrococcygeal mass. In the absence of metastatic disease, complete excision of the tumour with regional lymph node dissection will usually result in a prolonged disease-free survival. Once metastases develop, the disease tends to be slowly, but unremittingly, progressive.

References

Figure 2. Histology of the excised mass showing plump bipolar spindle cells and epithelial-looking cells arranged in loose fascicles with extensive myxoid degeneration forming pseudoacinar structures.

benefit for coccygectomy is in the group of patients with presacral ependymomas, with involvement of the coccyx. In our case, as the mass was found separate to the coccyx, coccygectomy was therefore not performed. Lymphadenectomy is also recommended in patients who present concurrently with inguinal lymph nodal metastases [8]. The role of lymphadenectomy in patients with iliac lymph nodal metastases is unknown as it is very rare. Although our patient had complete excision of the ependymoma and removal of the affected iliac lymph node, she developed local recurrence within 6 months of surgery suggesting a possible role for regional lymph node dissection. Intracranial ependymomas are radiosensitive; radiotherapy is an important local therapy [8]. However, radiotherapy has not been found useful as primary treatment for subcutaneous sacrococcygeal ependymomas. In those with incompletely excised tumour or in the presence of metastases, adjuvant radiotherapy is recommended [2, 8]. There is no evidence that these tumours respond to chemotherapy [2, 8].


1. Russell DS, Rubinstein LJ. Pathology of tumours of the nervous system (5th edn). London: Edward Arnold, 1989. 2. Johnson JM, Jessurun J, Leonard A. Sacrococcygeal ependymoma: case report and review of the literature. J Pediatr Surg 1999;34:1405–7. 3. Timmerman W, Bubrick MP. Presacral and postsacral extraspinal ependymoma: report of a case and review of the literature. Dis Colon Rectum 1984;27:114–9. 4. Morantz RA, Kepes JJ, Batnitzky S, Masterson BJ. Extraspinal ependymomas: report of three cases. J Neurosurg 1979;51:383–91. 5. Mallory FB. Three gliomata of ependymal origin: two in the fourth ventricle, one subcutaneous over the coccyx. J Med Res 1902;8:1–10. 6. Helwig EB, Stern JB. Subcutaneous sacrococcygeal myxopapillary ependymoma: a clinicopathologic study of 32 cases. Am J Clin Pathol 1984;81:156–61. 7. Maiorana A, Fante R, Fano RA. Myxopapillary ependymoma of the sacrococcygeal region. Pathologica 1989;81:471–6. 8. Kramer GWPM, Rutten E, Sloof J. Subcutaneous sacrococcygeal ependymoma with inguinal lymph node metastasis. J Neurosurg 1998;68:474–7. 9. Aktug˘ T, Hakgu¨der G, Sariog˘lu S, et al. Sacrococcygeal extraspinal ependymomas: the role of coccygectomy. J Pediatr Surg 2000;35:515–8. 10. Lynch J, Kelly N, Fitzpatrick B, Regan P. A sacrococcygeal extraspinal ependymoma in a 67-year-old man: a case report and review of the literature. Br J Plast Surg 2002;55:80–2. 11. Kernohan JW. Primary tumours of the spinal cord and intradural filum terminale. In: Cytology and cellular pathology of the nervous system. New York: Penfield W, 1932) 993–1025. 12. Bale PM. Ependymal rests and subcutaneous sacrococcygeal ependymoma. Pathology 1980;12:237–43.

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’ 2006 The British Institute of Radiology



Dose distribution to the mediastinum and heart The Editor—Sir, It is encouraging to read the detailed analysis of dose distributions to the mediastinum and heart during radical radiotherapy for oesophageal cancer in the UK [1]. This is a desirable consequence of better radiotherapy treatment planning using three-dimensional (3D) techniques and advanced software systems. Although the authors quote the Japanese experience of increased toxicity following chemoradiotherapy [2], there is important pioneering work from Japan using protons to reduce normal tissue doses while delivering a high tumour target dose using only two fields, with a respiratory gating technique linked to a pulsed output of protons. At The University of Tzukuba Proton Medical Research Centre, the potential indications for proton therapy have been extended by conducting a detailed clinical study of 46 oesophageal cancer patients using limited proton beam availability between 1985 and 1998. These mostly mixed protons and X-rays studies show that high total doses can be given with long term tumour control at least equivalent to X-ray therapy [3]. Toxicity reduced with experience, dose adjustment, increasing the component of protons and the use of respiratory gating. No late symptomatic complications were observed in the tracheobronchial tree, heart or spinal cord and no late grade 4 or 5 oesophageal toxicity has been observed since 1991. Severe and persistent oesophageal ulceration in some patients had previously occurred because of high doses per fraction (2.5– 3.7 Gy) related to limited beam availability. Since 2001, pure proton therapy treatments have been delivered in a purpose built Centre using rotating gantries and at fractionated doses of 1.8–2 Gy per fraction: the results of these studies are awaited with interest. Software for proton dose–volume histograms (DVH) have only been available recently at Tzukuba, so it is not possible to retrospectively analyse the patients treated in their publication. However, on a recently treated patient the DVH for a two field (anterior and posterior fields) 200 MV proton plan has been obtained and Table 1 shows a comparison of this plan with averaged values reported by Cominos et al [1] for their best case scenario of a 4 field plan (taken from Table 2 in their paper). The patient was a 61 year old man with a T1 N1 M0 squamous cell carcinoma of the mid-thoracic oesophagus. He had refused the offers of chemoradiotherapy and surgery. It should be noted that the UK prescribed dose is considerably lower than that used in Japan, yet the cardiac doses are markedly reduced in the latter, despite the anterior beam field traversing through the heart. Furthermore, the 4 field X-ray technique will inevitably result in higher lung exposures than for the 2-field proton technique, which were not reported.

448

Table 1. Dose distribution comparison for data in Cominos et al [1] and a single patient treated by proton therapy in Tzukuba

Prescribed dose (Gy) Mean cardiac dose (Gy) Volume of heart receiving > 30 Gy (%) Volume of heart receiving > 40 Gy (%) Volume of heart receiving > 45 Gy (%)

4 field X-rays (UK)

2 field protons (Japan)

54 (30 fractions) 41.4 77

72 (36 fractions) 18.7 15

51

10

26

8

Thus it has been demonstrated – as proof of principle in a small group of patients – that proton therapy can be successfully used in the treatment of oesophageal cancer with no evidence of late cardiac toxicity. The DVH analysis obtained in one patient is substantially better than with the megavoltage X-ray technique. It would be timely for the UK to acquire proton therapy facilities in order to improve therapeutic dose distributions in a wide variety of anatomical sites. Japan – with a population of c. 160 million – will soon have 8 centres capable of delivering proton or ion beam radiotherapy. To maintain a similar proportion, the UK would need to have 3 centres. Yours etc., 1

B JONES and 2 Y AKINE

1

University Hospital Birmingham Birmingham B15 2TH, UK 2 Proton Medical Research Centre University of Tzukuba Japan (Received 3 January 2006 and in revised form 18 January 2006, Accepted 18 January 2006) DOI: 10.1259/bjr/88462532

References 1. Cominos M, Mosleh-Shirazi MA, Tait D, Henrys A, Cornes P. Quantification and reduction of cardiac dose in radical radiotherapy for oesophageal cancer. Br J Radiol 2005;78:1069–74. 2. Satoshi I, Keiji N. Atushi O, Narikasu B. Long term toxicity after definitive chemoradiation of oesophageal and gastrooesophageal junction cancers. J Clin Oncol 2003;21:115–20. 3. Sugahara S, Tokuuye K, Okumura T, Nakahara A, Saida Y, Kagei K, et al. Clinical results of proton beam therapy for cancer of the oesophagus. Int J Radiat Oncol Biol Phys 2005;61:76–84.



CASE OF THE MONTH

A catastrophic complication D R VUMMIDI,

MRCP,

R S KASTHURI,

MRCS, FRCR

and R W BURY,

FRCR

Department of Radiology, Blackpool Victoria Hospital and University of Manchester, Manchester, Cheshire, UK Received 6 April 2005 Revised 12 May 2005 Accepted 27 June 2005 DOI: 10.1259/bjr/24873219 ’ 2006 The British Institute of Radiology

Case A 61-year-old Caucasian male was admitted as an emergency with central abdominal pain. On examination

he was in hypovolaemic shock, with a pulsatile lower abdominal mass. Based on a clinical diagnosis of ruptured abdominal aortic aneurysm, emergency surgical repair was undertaken. He was admitted to the

(a)

(b)

(c)

(d)

Figure 1. (a–d) Non-contrast CT brain (selected images). The British Journal of Radiology, May 2006

449

D R Vummidi, R S Kasthuri and R W Bury

Figure 3. An area of petechial haemorrhage in the cerebral parenchyma.

Figure 2. Magnified section from non-contrast CT (right frontal lobe).

intensive therapy unit following successful aneurysm repair. Renal dysfunction and gradual obtundation were features of a stormy post-operative period. His biochemical profile was significantly deranged and included elevated serum amylase. Worsening mentation was the immediate concern and a CT scan was organized to investigate this (Figures 1 and 2). What are the abnormalities demonstrated on the CT images? What is the differential diagnosis of this appearance? What case specific diagnosis would be appropriate?

Discussion The CT brain demonstrates marked attenuation of the grey/white matter differentiation. There is associated diffuse cerebral oedema with effacement of the sulci and basal cerebrospinal fluid (CSF) spaces. Furthermore, the magnified images also demonstrate multiple petechial haemorrhages (Figures 2 and 3). These are early findings of diffuse cerebral anoxia. There is often progression to a more diffusely low density brain, with possible ‘‘reversal’’ of the grey/white matter attenuation. At this stage, the differential for this appearance includes profound and persistent hypotension, asphyxiation and carbon monoxide fume inhalation [1]. In the given scenario, cerebral hypoperfusion would be most appropriate from the above mentioned list of differential diagnosis. Multiple cerebral fat emboli can also result in generalized cerebral hypoperfusion and result in similar appearances [1]. 450

Figure 4. Frozen section demonstrating arterial fat embolus (Oil Red O stain).

The described patient unfortunately succumbed to his illness. The post mortem examination revealed multiple cerebral fat emboli (Figure 4), secondary to acute necrotising pancreatitis. Fat embolism syndrome is an uncommon but recognized complication of acute necrotising pancreatitis [2]. Pathophysiologically, the chylomicrons and VLDL from the pancreatic fat necrosis are propagated into the vascular system resulting in cerebrovascular occlusion and consequent infarction. Multiple such embolisation into the intracranial vasculature results in global cerebral hypoperfusion. Cerebral CT scans demonstrate features as described above, but these are not specific to fat emboli. MRI is more sensitive in the neuroradiological diagnosis of intracranial fat embolism [3]. MRI findings include multiple intracranial infarcts, returning low signal on T1 weighted and relatively high signal on T2 weighted MRI. Similar areas are also seen to involve the corpus callosum and the basal ganglia. More diffuse grey and white matter abnormalities akin to CT imaging are demonstrated in the later stages. Diffusion weighted The British Journal of Radiology, May 2006

Case of the month: A catastrophic complication

MRI has also been reported to have role in the diagnosis of cerebral fat emboli [4]. Whilst the cerebral imaging features are specific to diffuse cerebral hypoperfusion, the aetiology is broad. This case besides highlighting an unusual cause of cerebral anoxia emphasises the importance of clinical correlation of the radiological differential diagnosis.

Acknowledgments The authors acknowledge Dr Patrick Shenjere, Specialist Registrar, Department of Histopathology, Blackpool Victoria Hospital, Blackpool, UK.


References 1. Osborn AG, Davis WL, Jacobs J. Cerebral vasculature: normal anatomy and pathology. In: Osborn AG, editor. Diagnostic neuroradiology. St. Louis, MO: Mosby, 1994:355–60. 2. Bhalla A, Sachdev A, Lehl SS, Singh R, D’Cruz S. Cerebral fat embolism as a rare possible complication of traumatic pancreatitis. JOP 2003;4:155–7. 3. Takahashi M, Suzuki R, Osakabe Y, Asai JI, Miyo T, Nagashima G, et al. Magnetic resonance imaging findings in cerebral fat embolism: correlation with clinical manifestations J Trauma 1999;46:324–7. 4. Parizel PM, Demey HE, Veeckmans G, Verstreken F, Cras P, Jorens PG, De Schepper AM. Early diagnosis of cerebral fat embolism syndrome by diffusion weighted MRI. Stroke 2001;32:2942.

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June 2006 Volume 79 Issue 942

June 2006, Volume 79, Issue 942 Three-dimensional MRI of the male urethrae with implanted artificial sphincters: initial results ●

Heel bone densitometry: device specific thresholds for the assessment of osteoporosis ●

The impact of a short course of study on the performance of radiographers when highlighting fractures on trauma radiographs: ‘‘The Red Dot System’’ ●

Portal vein embolisation prior to hepatic resection for colorectal liver metastases and the effects of periprocedure chemotherapy ●

Radiation benefit and risk at the assessment stage of the UK Breast Screening Programme ●

Radiation risks for the radiologist performing transjugular intrahepatic portosystemic shunt (TIPS) ●

Organ doses from prostate radiotherapy and associated concomitant exposures ●

Intensity-modulated radiation therapy in the treatment of gastric cancer: early clinical outcome and dosimetric comparison with conventional techniques ●

Changes in applicator positions and dose distribution high dose rate brachytherapy fractions in cervix patients receiving definitive radiotherapy ●

Clinical and cellular ionizing radiation sensitivity in a patient with xeroderma pigmentosum ●

Survival and initial chromatid breakage in normal and tumour cells exposed in vitro to gamma rays and carbon ions at the HIRFL ●

Usefulness of mini-tracheostomy and torque controlled insertion of applicator in fractionated endobronchial brachytherapy ●

●

MR findings of penile lymphoma

Deep tracheal laceration after balloon dilation for benign tracheobronchial stenosis: case reports of two patients ●

●

Correspondence

●

Neck pain: an unusual presentation of a common disease

●

Book reviews


Three-dimensional MRI of the male urethrae with implanted artificial sphincters: initial results 1,4

J DENG, MMed, PhD, 2M A HALL-CRAGGS, FRCR, 3M D CRAGGS, PhD, 1R RICHARDS, PhD, 3S L KNIGHT, 1 3 MSc, PhD, A D LINNEY, PhD and A R MUNDY, FRCS 1

Department of Medical Physics and Bioengineering, 2Department of Radiology and 3Institute of Urology, 4Department of Obstetrics and Gynaecology, University College London, Gower Street, London WC1E 6BT, UK ABSTRACT. The aim of this study was to develop a method for simultaneous 3D visualization of a new type of artificial urethral sphincter (AUS) and adjacent urinary structures. Serial MR tomograms were acquired from seven men after AUS implantation. 3D reconstruction was performed by thresholding original (positive) and inverted (negative) image intensity and by subsequently fusing positive and negative images. Results show that the bladder, cuff and balloons of the AUS of originally high intensity were imaged in 3D by thresholding the positive datasets. The urethrae and corpora cavernosa penis of originally low intensity were displayed in 3D by thresholding the negative datasets. Fusion of the positive and negative datasets allowed simultaneous visualization of the AUS complex and adjacent urinary structures. All the structures of interest were also clearly seen by interactive multiplanar reformatting. Coronal tomographic datasets provided better 3D and reformatted 2D images than sagittal and transverse datasets. This technique offers a simple means for evaluating the complex urethral anatomy and the AUS, and has potential for improved 3D visualization of many other complex morphological and pathological conditions.

Artificial urinary sphincters (AUS) have been used for many years to treat male stress urinary incontinence. The most widely-used type is the American Medical Systems AS800 AUS, but this is not without its problems, the most serious one being urethral erosion caused by the constant pressure of the inflated AUS on the urethra [1]. Postoperative imaging is important for assessing the position, configuration and function of an implanted AUS. The AS800 contains radio-opaque substances to make it visible on X-ray CT, but this exposes the patient to ionizing radiation. In addition, the contrast agents can degrade the device. To reduce the erosion risk, a new type of AUS with conditional occlusion has been developed by Craggs et al [2]. The Craggs AUS consists of four main parts (Figure 1). The perineal part is a cuff implanted around the bladder neck or urethra. The next two, intrapelvic or lower abdominal parts, are made up of two balloons, usually placed extraperitoneally close to the bladder. The fourth part is a scrotal pump connected to the balloons and cuff by tubing. The whole implant is made from silicone rubber and contains normal saline. This lends itself to MRI, a modality using no ionizing radiation. In JD is supported by an MRC Clinician Scientist Fellowship. We acknowledge some of the technical and equipment support from EPSRC-MRC’s MIAS IRC and RCR’s Pump Priming grant. The prototype artificial urinary sphincters were developed in collaboration with Bibby-Sterilin Limited and Isotron plc in a British Government Link project funded by the Department of Health and the Department of Trade & Industry.

The British Journal of Radiology, June 2006

Received 1 October 2004 Revised 19 April 2005 Accepted 13 October 2005 DOI: 10.1259/bjr/56511504 ’ 2006 The British Institute of Radiology

Figure 1. Diagram of Craggs artificial urinary sphincter. 1: urethral cuff; 2: primary reservoir/balloon; 3: additional reservoir (stress-relief balloon); 4: scrotal pump. The design allows a temporarily increased intra-abdominal pressure (caused by a stress such as coughing) to be transmitted from 3 into 1. This prevents stress incontinence, as well as urethral erosion that could be caused by a constant high pressure on the urethra even during rest when a conventional artificial sphincter was used (Artwork by Martin Knight). 455

J Deng, M A Hall-Craggs, M D Craggs et al

addition, MRI can embrace a large body volume without a need, unlike ultrasound, for a particular imaging window and direct contact with the body parts being examined. This is necessary for visualizing the Craggs AUS because it occupies both intrapelvic and perineal spaces, and its cuff is implanted around the delicate urethral tissue that should not be deformed for functional assessment. Consequently, MRI was chosen to visualize the AUS in this study. Ultrasound, CT and MR imaging have all been used for visualizing the pelvis in incontinent patients [3–5]. Conventional two-dimensional (2D) imaging requires expertise to compose a mental three-dimensional (3D) picture from a series of 2D images. Studies have demonstrated that 3D ultrasound [6, 7], CT [8] and MRI [9, 10] improve comprehension of the spatially complex morphology in the pelvis. To the best of our knowledge, there are no previous reports of the use of MRI for the assessment of AUS. Initially, we used cross-sectional images, directly available from standard MR scanners, for this assessment, and 2D MR appearances of the Craggs AUS will be described in a separate report. In brief, it is easy to comprehend the structures and positions of different parts (except the connecting tubes) of the AUS and their relationships to corresponding local body parts through serial 2D images. However, it is difficult to form a reliable 3D impression of all the AUS parts, the complex connections between them and the global spatial relationship to the relevant anatomies. This is clinically desirable, particularly when cross-sectional imaging planes are not perfectly aligned with ideal anatomical axes or imaginary axes of the AUS parts. Recently, 3D reconstruction has become available on some more advanced MR scanners. Three 3D display methods are often used on these systems. The first method is multiplanar reformatting. This requires further mental work from the operator to construct a 3D object. The second is volume 3D display, in which an object is rendered somehow transparent so that all the structures inside it become more or less visible, even when views to one structure may in reality be blocked by other structures in the front. The third, surface 3D display, probably offers the most realistic 3D images as structures of interest are represented as solid objects. Surface display is often achieved by first applying a threshold so that structures with signal intensities below it will not be rendered. This is sufficient for some clinical applications, such as for a CT skeletal examination; highintensity bone is visualized while low-intensity soft tissues are ‘‘removed’’. However, this approach is not sufficient for the imaging of complex biological structures where a simultaneous display of different intensity structures is necessary. As in this study, there is a need to show the AUS (mainly of high intensity) simultaneously with the bladder and urethra (mainly of high and low intensities, respectively). When a threshold appropriate for displaying the high-intensity structures (the bladder and AUS) is applied, the low-intensity structures (the urethra) are not seen. If a lower threshold appropriate for displaying the urethra is applied, irrelevant structures of high (fat) or intermediate (other soft-tissues) intensity are also shown, obscuring both low- and high-intensity structures of interest. In other words, simple thresholding 456

could not be used to present the two complexes simultaneously and distinguishably. In this paper, we describe the ‘‘inversion-fusion’’ method we have developed for processing sequential 2D MR datasets of patients with implanted Craggs AUSs [2]. This facilitates the simultaneous 3D visualization of the implant and the lower urinary tract. We then examine whether this 3D method could bring out information additional to 2D imaging. Finally, we discuss the feasibility of using this relatively simple, generic methodology with 3D facilities available on advanced commercial imaging modalities to improve 3D data reconstruction.

Materials and methods Data acquisition With approval from the Local Ethics Committee, nine patients suffering persistent post-prostatectomy incontinence received the Craggs AUS implantation after giving informed consent. Excluding one patient with extensive surgical metalwork in his pelvis, all other eight underwent MR scans about 2 months after implantation. They were scanned on a 1.5 T Siemens Vision system (Erlangen, Germany) using a phase-array body coil. Images were acquired using T2 weighted 2D Tru-FISP sequence (repetition time (TR): 6.32 ms; echo time (TE): 3 ms; field of view (FOV): 230 mm; rectangular field of view (RFOV): 50%; flip angle (FA): 70 ˚; time of aquisition (TA): 12–21 s) in the coronal, transverse and sagittal planes. In this study, the term ‘‘series’’ is being used to describe a set of anatomically consecutive slices acquired during one MR scan. Because of the difficulty comprehending global relationships between all AUS parts and relevant urinary structures using individual series acquired from the first MR patient, several anatomically consecutive series were acquired from each of the remaining seven patients (aged from 59 years to 76 years). This allowed the entire AUS and the neighbouring anatomy to be sampled by these series at regular intervals. A total of 86 series were recorded from the 7 patients (4 to 20 from each). Each series consisted of 7 to 12 contiguous slices, all with a slice thickness of 5 mm (minimum achievable by the scanner at the time). A serial 3D dataset was then formed by one series or by combining two to six consecutive series, consisting of 7 to 42 slices (Table 1). In five patients, some serial datasets were acquired in pairs, offset by 50% of the slice thickness (Figure 2a,b), and each of the pairs was later used for creating an interlaced 3D dataset (Figure 2c). The pixel matrix in each slice was 2566256, with a pixel size of 0.898460.8984 (mm). The voxel matrix (a 3D dataset) was therefore 2566256 for x and y dimensions and from 7 to 42 for z dimension. The resultant voxel size is 0.898460.898465.0000 (mm) for serial datasets and 0.898460.898462.5 (mm) for interlaced datasets.

Data processing The datasets were transferred to an MGI 3D Workstation (Medical Graphics & Imaging Group, The British Journal of Radiology, June 2006

3D MRI of male urethra and artificial sphincter Table 1. 3D results from 35 original 3D datasets and 8 interlaced 3D datasets Patient age (years)

Sequences

Slices per series

Series per 3D dataset

Slice thickness

Slice interval

Orientation

3D image quality

Notes on reconstruction quality

59

B006-026 B027-047 B006-047mix

7 7 7

3 3 6

5 5 5

5 5 2.5*

Coronal Coronal Coronal

Adequate Adequate Good

Big slice interval Big slice interval Mild motion artefact, not affecting ROI

76

D069-077 D078-085 D128-135 D145-154 D015-041mix D101-127mix D005-014 D042-050 D050-058 D092-100 D137-145

9 9 9 9 9 9 9 9 9 9 9

1 1 1 1 3 3 1 1 1 1 1

5 5 5 5 5 5 5 5 5 5 5

5 5 5 5 5 5 5 5 5 5 5

Coronal Coronal Coronal Coronal Coronal Coronal Sagittal Sagittal Sagittal Sagittal Sagittal

Adequate Adequate Adequate Adequate Good Good Adequate Poor Poor Poor Poor

Big Big Big Big

H005-033mix H034-051mix H074-108mix H109-143mix H005-034int-mix

7 7 7 7 7

4 4 5 5 8

5 5 5 5 2.5

5 5 5 5 2.5*

Coronal Coronal Coronal Coronal Coronal

Good Good Good Good Good

H074-109int-mix

7

10

2.5

2.5*

Coronal

Good

H052-058 H144-150

7 7

1 1

5 5

5 5

Sagittal Sagittal

Adequate Adequate

M005-033 M034-051 M057-094 M095-122 M005-034int

7 7 7 7 7

4 4 4 4 8

5 5 5 5 2.5

5 5 5 5 2.5*



M057-095int

7

8

2.5

2.5*

Coronal

Good

S011-031mix S032-052mix S072-092mix S093-113mix S011-032int-mix

7 7 7 7 7

3 3 3 3 5

5 5 5 5 2.5

5 5 5 5 2.5*



S072-093int-mix

7

5

2.5

2.5*

Coronal

Good

S053-059 S114-120

7 7

1 1

5 5

5 5

Sagittal Sagittal

Adequate Adequate

Mild motion artefact, not affecting ROI Mild motion artefact, not affecting ROI Big slice interval Big slice interval

V005-017 V018-029 V030-041mix

12 12 12

1 1 1

5 5 5

5 5 5

Coronal Sagittal Transverse

Good Adequate Poor

Big slice interval Big slice interval

Y005-029mix Y030-053mix Y005-030int-mix Y055-101mix

12 12 12 12

2 2 4 3

5 5 5 5

5 5 2.5* 5

Coronal Coronal Coronal Transverse

Good Adequate Adequate Adequate

Moderate motion artefact Moderate motion artefact Big slice interval

55

60

69

64

67

slice slice slice slice

interval interval interval interval

Big slice interval Too big slice interval Too big slice interval Too big slice interval Too big slice interval

Mild motion artefact, not affecting ROI Mild motion artefact, not affecting ROI Big slice interval Big slice interval

Mild motion artefact, not affecting ROI Mild motion artefact, not affecting ROI

*Interlaced datasets. ROI, region of interest.

University College London, www.medphys.ucl.ac.uk/ mgi/workstat.htm), and converted into MGI format using DispImage program [11]. Interlaced 3D datasets were generated by interweaving and then renumbering paired slices from paired serial datasets (Figure 2c). The British Journal of Radiology, June 2006

There were two complexes of interest: the urinary complex consisting of the bladder, urethra and penis, and the AUS complex consisting of the cuff, the balloons and the tubes connecting them. On T2 weighted images, the urinary complex mainly showed low intensity except for the urine within the bladder and urinary tract, which 457

J Deng, M A Hall-Craggs, M D Craggs et al Figure 2. Schematic drawings of data acquisition and post-processing. Pixels (voxels) of high, intermediate and low intensities on original (positive) images are numbered 1, 2, and 3, respectively. (a,b) Two series are acquired, each with a 5 mm slice interval. The second series is scanned with a 2.5 mm offset from the first one. (c) After interweaving the slices of the two series, a (positive) dataset (Slices 0–5) is created, with the slice interval halved to 2.5 mm and the number of slices doubled. A negative dataset (Slices 09–59) is also created, changing pixels of low-intensity (numbered 3) into high-intensity. (d) Two thresholds appropriate for displaying originally high intensity pixels in Slices 0–5 and negatively high intensity pixels in Slices 09–59, correspondingly, remove intermediate and low intensity pixels. Background pixels in both positive and negative datasets are not displayed. The remaining pixels (voxels) are areas of interest, which can then be fused into a single 3D image (not shown here), restoring the original spatial relationship between pixels (voxels) 1 s and 3 s.

(a)

(b)

Figure 3. One of serial slices from a coronal scan. (a) In the original (positive) sequence, the bladder, balloon and cuff of the artificial sphincter are of high-intensity, which can be segmented by simply applying a threshold to remove low-intensity signals, and then rendered in 3D surface display (Figure 4a). However, this will also remove the corpora cavernosa from being displayed because their intensities are below the threshold. (b) After inversing the grey scale, the corpora become high-intensity structures which can be displayed by simply applying another threshold (Figure 4b). In the 2D images, it is difficult to tell whether the cuff is fully sealed around the corpora cavernosa urethrae.

was high signal. The AUS complex showed high intensity. As mentioned earlier, simple thresholding could not be used to present the two complexes together clearly. 458

To address this problem, a negative dataset was created by inverting the intensity of the native, positive dataset (Figure 2c,d and Figure 3). The low-intensity structures (urethra and penis) became high-intensity The British Journal of Radiology, June 2006

3D MRI of male urethra and artificial sphincter

(a)

(b)

Figure 4. 3D reconstruction with surface displays. (a) Structures above corresponding thresholds in the positive and negative sequences are separately visualized. In order not to obscure the smaller structures (in grey), the positive sequence is placed closer to the readers, so the bladder is placed in front of the lower abdomen. The threshold in the negative sequence is set to intermediate low so that the separations of the three spongy structures are displayed as solid while their inner regions as (artificially) hollow. (b) The structures from both positive and negative sequences are fused together to reveal the relationship between the cuff and the urethra; in this case, the cuff is fully sealed (also see movies on our website for a better 3D perception).

ones in the negative dataset. The positive and negative datasets were then combined into a new, positive+negative dataset (Figure 2c).

Figure 4a). Finally, by putting the two objects into the same co-ordinate system, a fused image is created for assessing their spatial relationships (Figure 4b and Figure 5).

3D reconstruction This was also carried out on the MGI Workstation. The system provides various options for the reconstruction and manipulation of 3D datasets from serial tomographic imaging modalities [12–16]. 3D surface display and multiplanar reformatting were used in this study to present the structures of interest for analysis. 3D surface display for structures in the overlapping datasets was achieved by four main steps. First, two blocks containing structures of interest in a positive+negative sequence were defined under the guidance of multiplanar reformatting. This avoided segmenting most of the fat in the abdomen, skeletal muscles in the pelvic floor and thighs (in a negative dataset which had more or less the same intensity as that of the urethra) and the image background area (which became the highest intensity area in a negative dataset). Second, two thresholds appropriate for rendering the originally high-intensity structures (the bladder, the balloon and the cuff) and the inversely high-intensity structures (the corpora cavernosa) were applied to the two blocks separately. This resulted in the structures of interest being roughly visualized (Figure 2d). Third, some ‘‘image surgery’’ was performed to remove remaining unwanted parts that were still above thresholds and therefore might obscure the viewing of structures of interest. Having done these, the two complexes could be displayed for inspection separately (Figure 2d and The British Journal of Radiology, June 2006

Data analysis The level of analysis was at the dataset level (Table 1) and between various scan settings (Table 2). The quality of the final reconstructed 3D images in terms of depicting the spatial relationship between the structures of interest was scored by two observers (JD, RR) as poor, adequate and good. ‘‘Poor’’ means that in a single fused 3D object, the spatial relationship between the structures of interest cannot be visualized, and subsequently, the AUS functioning position cannot be diagnostically assessed. ‘‘Adequate’’ means that in a single fused 3D object, the spatial relationship of the AUS cuff to the corpora cavernosa urethra was visualized. Some other parts of the urinary tract or of the AUS were not clearly visualized, but this did not affect our judgement of the completeness of the AUS cuff wrapping the urethra, the most important diagnostic issue to be assessed. ‘‘Good’’ means that in addition to the cuff-urethral relationship, the entire AUS complex and the entire urinary complex were well visualized.

Results A total of 35 datasets containing the region of interest for 2D and 3D analysis were obtained directly or combined from 86 series acquired. Eight interlaced 459


Figure 5. A 3D surface display showing that the cuff is not fully wrapped around the corporus cavernosum urethrae. Again, it is an image after fusing the cuff in the positive and the penis in the negative sequences. The threshold in the negative sequence is set very low so that the entire spongy structures are displayed as solid (compare with Figure 4). Also see a movie on our website.

artefact was noticeable in all eight datasets, but this degraded the bladder more than the cuff (Figure 6). The spatial relationships of all the structures of interest were best assessed when the images from both positive and negative datasets were fused. Urinary structures and the AUS were clearly demonstrated using 3D surface display, and this was particularly well seen when viewed as a 3D movie (available at: www.medphys. ucl.ac.uk/mgi/jdeng/ follow the links ‘‘Dynamic Face/ Body Parts’’ on the top panel, then ‘‘Erecting Penis’’ on the left panel). The 3D images have depicted various cuff conditions between the well-implanted, fully wrapped (Figure 4) to the failed, fully open (Figure 5), and from un-inflated to fully inflated. 2D views of the region of interest can be interactively obtained using multiplanar reformatting under the guidance of 3D objects (Figure 6), offering a means for detailed examination of structures that may not be well visualized by simple 3D reconstruction. This is particularly useful for following up structures of heterogeneous intensity such as the tubing between the balloons and cuffs and for following the thin urine residues along the urethra that were not acquired in single original imaging planes. All these have been found difficult when only viewing original 2D slices.

Discussion 2D versus 3D imaging

datasets were further created from the 35 datasets (Table 1). Depending on data quality and suitability, it took about 10–30 min to process one sequence and attain fused 3D images. The longer time was caused by processing few, but graphically more difficult, sagittal and transverse datasets using manual segmentation of regions of non-interest (i.e. unwanted abdominal fat or other tissues with complex greyscale composition, see Discussion on Image processing). The ‘‘inversion-fusion’’ method itself, together with series-combination and sliceinterlacing, only took a few seconds. The 3D results were listed in Table 1. The comparison of the 3D images acquired from coronal, sagittal and transverse base data is shown in Table 2. The coronal datasets offered the best 3D results with 65.2% good and 34.8% adequate images. None of the sagittal and transverse datasets produced good 3D images, but six of 10 sagittal datasets were adequate. There were only two transverse datasets, precluding assessment. Interlaced datasets provided the best 3D results with good 3D images obtained in seven (87.5%) of eight datasets (all derived from coronal datasets). Motion

This initial study has shown that, by using our ‘‘inversion-fusion’’ method, 3D images simultaneously displaying both AUS and urinary complexes can be obtained from serial MR datasets. Although subject to further large-scale studies, visualizing the datasets as fused 3D objects and through reformatted 2D images appears to have at least three advantages over viewing them as original individual 2D slices. First, images reconstructed by 3D surface rendering can offer more straightforward, panoramic views of all four main parts of the AUS and their relationships to each other and to the urinary tract. This is important for post-operative assessment as the four parts are implanted in three very different anatomical places. It is possible for an experienced radiologist to form such a 3D impression from individual slices, but it requires greater training while still subject to human errors. Second, multiplanar images reformatted from a 3D dataset in any desirable orientation allow tracking of the entire course of the tubing connection between the AUS parts. This has proved to be very difficult when tracking through the original 2D slices. It is crucial to check this

Table 2. Initial comparison of the number and percentage of useful 3D series/datasets between various scan settings and postprocessings 3D image quality

Coronal series/datasets Poor

Adequate Good

Original No. (%) 0 (0.00) 8 (34.8) Interlaced No. (%) 0 (0.00) 1 (12.5)

460

Sagittal series/datasets Subtotal Poor

15 (55.2) 23 7 (87.5) 8

Transverse series/datasets

Adequate Good

4 (40.0) 5 (50.0)

Subtotal Poor

0 (0.0) 10

Adequate Good

1 (50.0) 1 (50.0)

Subtotal

0 (0.0) 2


3D MRI of male urethra and artificial sphincter

(a)

(b)

Figure 6. Multiplanar reformatting of the same 3D dataset in Figure 4a: Three orthogonal images are obtained by crosssectioning the bladder, showing the motion artefact (arrowheads) in the interlaced data, which appears more severe in the upper side of the bladder. (b) Another three orthogonal images are obtained by cross-sectioning the cuff, showing the wellsealed cuff, although not so apparent as in 3D images. Note the depiction of the tubing (arrows) connecting the cuff and the balloons. Its entire course may not be visualized by 3D surface display due to its heterogeneous signal intensity, but can be traced by interactively reformatting sequential 2D images.

connection when blockage and/or entanglement are suspected. Third, the full 3D visualization of the AUS complex has paved the way for its volumetric quantification. This will be helpful for evaluation of possible AUS leakage by comparing the given amount of saline pumped in to the increment in volume before and after the pumping. It is impossible to work out such a change mentally unless there is serious leak. More generically speaking, this study has shown that clinically useful 3D images can be attained by combined use of commercially available modality and acquisition methods (MR interlaced acquisition with coronal sequences) with relatively simple post-processing and display methods (intensity inversion and image fusion). Admittedly, 3D techniques have not been sophisticatedly developed such that they can render texture details. In this study, these features, such as the intricate structures of the AUS were seen better on the 2D slices (either from original cross-sectional data or from multiplanar reformatting).

Imaging modality Ultrasound examination has a narrow FOV for superficial structures, and soft-tissue deformation results from the direct contact between transducer and the skin [15]. There is also insufficient definition of the AUS structures, as seen in our unpublished 3D ultrasound study of some of the patients. These prevent ultrasound from being our first choice for structural assessment. However, the latest real-time 3D colour Doppler ultrasound is likely to play a role in evaluation of penile haemodynamics under The British Journal of Radiology, June 2006

different AUS hydraulic and other physiological circumstances. It may also enable real-time checking of leakage during pumping [17–19].

Imaging resolution The in-plane image quality obtained from this MR study is much higher than that obtained from our unpublished 3D ultrasound study. This is because, compared with ultrasound, MRI has higher intrinsic tissue contrast and no restriction on imaging window [20, 21]. However, the out-of-plane (z-dimension) spatial resolution is far from ideal. For instance, while the voxel sizes in the x and y dimensions were equally small (about 1 mm), the size in the z-dimension was 5 mm. To address this unbalanced match, interlaced acquisitions were performed in the last five patients and this halved the z voxel size to 2.5 mm (Figure 2b). Thinner slices and true 3D data sets can now be acquired in our centre after this study. Although this will reduce the degree of anisotropy, it is unlikely to be eliminated. In addition, no matter how thin a slice may become available from a newer scanner, radiologists will immediately expect to see tissue details at an even smaller scale. Consequently, interlaced acquisition may still be helpful in future studies.

Imaging orientation In this study, coronal datasets provided the best 3D results (Table 2). This is because the slices are obtained perpendicular to the long axis of the penile urethra and 461


the cuff, offering multiple serial slices through the structures of interest. In contrast a sagittal or transverse series acquires only three to four slices through the cuff in the penis, simply not offering sufficient data for 3D reconstruction. Though, recently available thinner slice acquisition may change this.

Motion artefacts There are several sources of movement and misregistration artefacts that cause image degradation: respiration, peristalsis, urine and blood flow, gross body and penile movement. Because the sequence acquisition time was up to 21 s, periodic motion was likely to cause artefacts. However, obvious motion artefact was mainly due to misregistration between the added image series. This is mainly because it is difficult to breath-hold at exactly the same point of respiration for each image series. Fortunately, the impact of motion artefact appeared minimal in the region of the cuff (Figure 6), primarily because it was extraperitoneal and stabilized by the urethra and local ligaments. In the future, faster acquisition and motion tracking [22] may minimize motion artefacts.

Image processing With a few computer mouse clicks, the interlacing and inverting techniques have improved resolution and speeded up segmentation of the region of interest, respectively. Nonetheless, cumbersome manual segmentation is still needed for removal of regions of noninterest but of a wide-range of mixed signal intensities. Clinicians are frequently caught in a dilemma in 3D segmentation of graphically complicated structures. On one hand, simple algorithms such as thresholding are often unable to depict all the structures of interest that have heterogeneous intensities (even when they are made up of the same biological tissues), or have variable intensity gradients between different tissue borders. This often results in incomprehensible or even misleading 3D images. On the other hand, complex algorithms often demand extra expertise and are time-consuming. This study has tried to strike a balance between the simplicity of the image processing and the quality of the resulting images. Some of the generic tricks developed in this study such as ‘‘inversion-fusion’’, can be easily adapted by commercial systems to speed up and simplify 3D reconstruction. In fact this has been implemented in the latest MGI 3D Workstation for commercial release. In addition, automatic contouring with operator interactive correction may allow rapid segmentation with sufficient accuracy. Automatic contouring may also be used for fast removal of extraneous, but homogeneous, tissues, such as high-intensity fat from the abdominal wall (Figure 6), which would otherwise obscure the region of interest in 3D images. Hence, the entire post-processing and display time are expected to be reduced to about 10 min. A final point to make is that for visual reality and clarity, the ‘‘inversion-fusion’’ method was only used for 3D surface rendering in this study, but in principle, it can also benefit volume rendering. 462

Conclusion This MR study has demonstrated the feasibility of simultaneous visualization of all relevant parts of the AUS and the lower urinary tract in three dimensions using these novel image processing methods. This has potential for reducing difficulty in comprehending the spatial position, geometry and function (and malfunction) of implanted sphincters and the impact on the local anatomy. The inversion-fusion approach can be developed as a general application for simplifying and improving 3D visualization of many other complex morphological and pathological conditions.

Acknowledgments We thank the MR staff of Middlesex Hospital for data acquisition, D Plummer and D McDonald of Medical Physics Department for developing DispImage and other image conversion programmes. We are also grateful to A Todd-Pokropek for advice on MR data processing and C H Rodeck of Obstetric Department for research planning.

References 1. Venn SN, Greenwell TJ, Mundy AR. The long-term outcome of artificial urinary sphincters. J Urol 2000;164:702–6. 2. Craggs MD, Mundy AR, Dwyer PL, Susser J, Knight SL. Long-term clinical outcome of a new AUS with conditional occlusion for genuine stress incontinence: optimal cuff pressures for continence. BJU Int 2002;90:18. 3. McFarlane IP, Foley SJ, Shah PJ. Long-term outcome of permanent urethral stents in the treatment of detrusorsphincter dyssynergia. Br J Urol 1996;78:729–32. 4. Maki DD, et al. Injected periurethral collagen for postprostatectomy urinary incontinence: MR and CT appearance. Abdom Imaging 2000;25:658–62. 5. Mostwin JL, Genadry R, Saunders R, Yang A. Stress incontinence observed with real time sonography and dynamic fastscan magnetic resonance imaging--insights into pathophysiology. Scand J Urol Nephrol Suppl 2001;94–9. 6. Klein HM, Kirschner-Hermanns R, Lagunilla J, Gunther RW. Assessment of incontinence with intraurethral US: preliminary results. Radiology 1993;187:141–3. 7. Umek W, et al. The urethra during pelvic floor contraction: observations on three-dimensional ultrasound. Obstet Gynecol 2002;100:796. 8. Stenzl A, et al. 3-dimensional computerized tomography and virtual reality of the lower urinary tract in women. Int Urogynecol J Pelvic Floor Dysfunct 1999;10:248–53. 9. Myers RP, et al. Puboperineales: muscular boundaries of the male urogenital hiatus in 3D from magnetic resonance imaging. J Urol 2000;164:1412–5. 10. Ueno S, et al. Three-dimensional display of the pelvic structure of anorectal malformations based on CT and MR images. J Pediatr Surg 1995;30:682–6. 11. Plummer DL. A display and analysis tool for medical images. Rivista di Neuroradiologia 1992;5:489–95. 12. Linney AD, et al. Three-dimensional visualization of data on human anatomy: diagnosis and surgical planning. J Audiov Media Med 1993;16:4–10. 13. Brookes JA, Deng J, Wilkinson ID, Lees WR. Threedimensional imaging of the postmortem fetus by MRI: early experience. Fetal Diagn Ther 1999;14:166–71.


3D MRI of male urethra and artificial sphincter 14. Deng J, Brookes JA, Gardener JE, Rodeck CH, Lees WR. Three-dimensional magnetic resonance imaging of the postmortem fetal heart. Fetal Diagn Ther 1996;11: 417–21. 15. Deng J, et al. Novel technique for three-dimensional visualisation and quantification of deformable, moving soft-tissue body parts. Lancet 2000;356:127–31. 16. Deng J, et al. Dynamic three-dimensional colour Doppler ultrasound of human fetal intracardiac flow. Ultrasound Obstet Gynecol 2002;20:131–6. 17. Deng J. Terminology of three-dimensional and fourdimensional ultrasound imaging of the fetal heart and other moving body parts. Ultrasound Obstet Gynecol 2003;22:336–44. 18. Deng J. Minimally compressed imaging of deformable moving body parts using dynamic 3D ultrasonography and colour Doppler - with facial and genital parts as exemplars.


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Heel bone densitometry: device specific thresholds for the assessment of osteoporosis 1

E MCCAULEY,

MSc,

1

A MACKIE,

PhD,

3

D ELLIOTT,

RGN, ONC

and 2A CHUCK,

FRCP

1

Regional Medical Physics Department and 2Department of Rheumatology, University Hospital of North Durham, North Road, Durham DH1 5TW and 3Durham and Chester-le-Street Primary Care Trust, John Snow House, Durham University Science Park, County Durham DH1 3YG, UK

ABSTRACT. A diagnosis of osteoporosis is facilitated by bone mineral density (BMD) measurement of the lumbar spine and hip using dual energy X-ray absorptiometry (DXA), interpreted in accordance with criteria published by the World Health Organization (WHO). The use of peripheral DXA is growing in primary care and guidance on its use has recently been published by the National Osteoporosis Society (NOS), recommending a triage approach using thresholds specific to each type of peripheral device. However, no data currently exist for the Norland Apollo heel densitometer (Cooper Surgical, Trumbull, USA). 215 women between 50 years and 75 years of age (mean age 64.6 years) referred for hip and spine BMD measurements also had a heel BMD measurement. Device specific upper and lower thresholds were calculated for the Norland Apollo heel densitometer to give a 90% sensitivity and 90% specificity for osteoporosis at the hip or spine. Patients with a heel T-score of above –1.2 are very likely to have normal bone density on axial densitometry, whilst patients with heel T-score of below –2.2 are very likely to have osteoporosis at the hip or spine. Patients whose measurements lie between the thresholds should be referred for axial DXA.

A diagnosis of osteoporosis or osteopenia can be made using dual energy X-ray absorptiometry (DXA) of the lumbar spine or femur. The defining criterion for the diagnosis is the value of the T-score, which is the number of standard deviations the measured bone mineral density (BMD) is above or below the young adult reference BMD. According to the World Health Organization (WHO) criteria [1], the patient is osteoporotic if the T-score # –2.5 and osteopenic if the T-score .–2.5 but # –1. Axial densitometers, which are designed to measure BMD at the hip and/or spine, are large, expensive and require a permanent location, ordinarily in a secondary care setting. Significant research into the development of alternative methodologies has resulted in a plethora of techniques and alternative measurement sites [2]. The measurement of peripheral bone density is growing outside secondary care centres, principally because of the much lower cost of these devices compared with axial densitometers. They are also very easy to use, requiring minimal set up and positioning, and are relatively portable. Heel densitometers produce results in terms of BMD (g cm22) and a heel T-score. Manufacturers’ data accompanying devices often quote the WHO criteria indicating that a heel T-score of # 22.5 would be indicative of osteoporosis. It is well established, however, that it is inappropriate to interpret any peripheral T-score as one would for axial densitometry [3, 4] and WHO T-score criteria applied to heel DXA would significantly underestimate the prevalence of osteoporosis [3, 5]. 464

Received 30 August 2005 Revised 14 October 2005 Accepted 31 October 2005 DOI: 10.1259/bjr/47352663 ’ 2006 The British Institute of Radiology

In addition, it is well known that measurements of BMD vary from scanner to scanner and between scanner types; particularly between manufacturers because of the variation in acquisition and analysis techniques as well as the absence of a standard calibration technique [6–10]. The International Society for Clinical Densitometry (ISCD) state that although peripheral DXA measurements can theoretically be used to identify patients unlikely to have osteoporosis, and also identify patients who should be treated, this cannot be applied in clinical practice until device-specific cut-off points are established [11]. The National Osteoporosis Society (NOS) has recently produced a revised position statement concerning the recommended use of peripheral densitometry [12]. NOS recommend a triage approach using T-score thresholds, which are specific to each type of peripheral device. These thresholds are defined so that patients with osteoporosis at the hip or spine are identified with 90% sensitivity and 90% specificity. Patients with a peripheral T-score below the lower threshold are very likely to have osteoporosis at the hip or spine, whereas patients with a peripheral T-score above the upper threshold are unlikely to have osteoporosis. Those patients who have a heel T-score that falls between the two thresholds require a hip and spine BMD examination for a definitive diagnosis. The NOS statement gives device specific threshold data for five peripheral devices, two of which are heel densitometers. Data in this form are not yet published for the Norland Apollo heel densitometer (Cooper Surgical, Trumbull, USA). The British Journal of Radiology, June 2006

Heel bone densitometry

In 2001, the Durham and Chester-le-Street Primary Care Trust (PCT) purchased a Norland Apollo DXA scanner for the measurement of heel bone density within the primary care setting. The device was to be operated by the osteoporosis prevention specialist nurse within a protocol agreed between primary and secondary care. The aim of this study was to establish thresholds for Apollo heel BMD T-scores for use in risk stratifying patients within a primary care management protocol.

Method Ethical committee approval was obtained to measure heel BMD with the Apollo in women attending for axial bone density measurements; all fulfilled at least one accepted referral criterion [13]. 76 women aged between 50 years and 75 years (mean age 64.0 years) gave informed written consent. Axial BMD was measured for the lumbar spine, L2–4, and for the non-dominant femur using a Lunar DPX-IQ (General Electric/Lunar Corp., Madison, WI) densitometer. Heel BMD, of the dominant side, was measured with a Norland Apollo portable densitometer. Measurement of the dominant side was acceptable since published data demonstrates no significant difference between BMD measurements of the heel between the right and left sides [5, 14]. Patients were then categorised as either not osteoporotic (T-score . –2.5) or osteoporotic (T-score # –2.5), dependent upon their lowest axial T-score, i.e. total femur or spine (L2–4). 23 of these patients were found to have normal axial T-scores, 34 were osteopenic and 19 patients were osteoporotic. Threshold values were initially calculated using this preliminary data so that the Apollo DXA scanner could be used in clinical practice as quickly as possible, although insufficient patients had been examined for the data to achieve the statistical accuracy recommended by NOS, i.e. 95% confidence that the true sensitivity and specificity do not fall below 80%. The initial lower threshold was a T-score of –2.2. The initial upper threshold was 0.1, although this value was obviously very inaccurate since there was a large difference in the T-scores of the patients at the cut off point and those patients immediately above and below them, as is demonstrated in Figure 1. Patients were seen in a primary care setting by an osteoporosis prevention specialist nurse, who performed heel DXA measurements on patients considered at risk and referred those patients for axial densitometry with heel T-scores below the upper threshold value of 0.1. Heel and axial T-scores were recorded for every patient referred in this way. Data for a further 139 patients between the ages of 50 years and 75 years were recorded, mean age 64.9 years.

Figure 1. Scatter plot of heel T-scores for the initial group of patients; 90% of the 57 not osteoporotic patients lie above the lower threshold of 22.2, whilst 90% of the 19 patients who are osteoporotic lie below the higher threshold of 0.1. The limited data points in the osteoporotic group do not allow an accurate cut off value to be selected.

Results In total, heel and axial DXA T-scores were measured for 215 women, mean age 64.6 years. Of these, 71 patients were found to be osteoporotic on axial densitometry and 144 patients were not osteoporotic. Figure 2 shows a scatter plot of heel T-scores for patients in the two diagnostic categories. The line The British Journal of Radiology, June 2006

Figure 2. Scatter plot of all heel T-scores for patients in the two diagnostic categories; 90% of patients who have a Tscore above the lower threshold of 22.2 do not have osteoporosis, whilst 90% of patients who have a T-score below the upper T-score threshold of 21.2 are osteoporotic. Patients with heel T-score between the two threshold values should be referred for axial densitometry.

465

E McCauley, A Mackie, D Elliott and A Chuck

through the not osteoporotic group shows the lower threshold, a T-score of 22.2; 90% of patients who have a T-score above this threshold do not have osteoporosis. The line through the osteoporotic group shows the upper threshold, a T-score of 21.2; 90% of patients who have a T-score below this threshold are osteoporotic.

Discussion The information provided by the manufacturers of the Norland Apollo heel densitometer states that a heel Tscore of # 22.5 is indicative of osteoporosis. This assumption that the WHO criterion for axial DXA can be applied to heel DXA has been previously demonstrated to be incorrect. Only 15 of the 71 patients (21.1%) who were found to be osteoporotic on axial densitometry had heel T-scores of # 22.5. Data were initially acquired so that a working threshold value could be quickly calculated and used within the primary care setting. These original 76 patients had their hip, spine and heel DXA measurements all performed on the same day. Subsequently, patients were then referred for axial DXA on the basis of established referral criteria or on the value of their heel DXA measurement. These latter results will therefore introduce some bias in the data, as patients who were found to have high T-scores at the heel, above the higher threshold value, were not referred for axial DXA, despite the presence of other risk factors. These patients, with a very low risk of axial osteoporosis, were therefore excluded from the study as it was the intention of the initial threshold values to avoid unnecessary referrals for axial DXA. Since patients with high heel T-scores were excluded, this bias will have the effect of lowering the average T-score measured in the latter group of patients. These patients represent 65% of the total patient group. This measurement bias prevents these data from being incorporated into the NOS published data for device specific thresholds. Also, NOS consider only women between the ages of 55 years and 70 years, whereas these data have been acquired for patients between 50 years and 75 years of age. However, the acquisition of truly unbiased data within a more narrow age range and with enough patients for statistical accuracy, i.e. 95% confidence that the true sensitivity and specificity do not fall below 80%, would take more than 10 months with this department’s current workload. This estimation assumes that only half of all patients referred are within the required age range (based on 2004 data) and less than one third of patients in this age group are likely to be osteoporotic. These biased data therefore, may be considered an approximation of the true thresholds for the Norland Apollo heel densitometer. The true values are likely be slightly higher, due to the exclusion of patients with high T-scores, therefore caution is advised with heel T-scores just above the upper threshold of 21.2. These patients should be considered to have an equivocal result and be referred for axial DXA to prevent misdiagnosis. Patients with heel T-scores just above the lower threshold (22.2) will be referred for axial densitometry rather than be given a diagnosis of osteoporosis from the heel DXA alone. In this case, no patients should be misdiagnosed, 466

but the number of referrals to axial DXA will therefore be higher than if the threshold values were unbiased. The lower threshold of 22.2 was unchanged from that calculated using the initial group of 57 patients. The higher threshold was significantly lowered to 21.2 when a total of 70 osteoporotic patients were considered, compared with the initial value of 0.1 which was calculated from only 19 patients. The difference in these figures is not surprising, and is simply due to the variability of T-scores within small numbers of patients.

Conclusions Thresholds have been calculated for a Norland Apollo heel densitometer; patients with a heel T-score of above 21.2 are very likely to have normal bone density on axial densitometry, whilst patients with heel T-score of below 22.2 are very likely to have osteoporosis at the hip or spine. Patients with heel T-scores that lie between these two thresholds should be referred for axial densitometry. Users of a Norland Apollo heel densitometer may wish to implement these threshold values, with the caveat that they are an approximation but are the best data currently available for this scanner.

References 1. World Health Organization. Assessment of fracture risk and its application to screening for postmenopausal osteoporosis: Report of WHO Study Group. Technical report series 843. Geneva, Switzerland, WHO 1994. 2. Placide J, Martens MG. Comparing screening methods for osteoporosis. Curr Womens Health Rep 2003;3:207–10. 3. Fordham JN, Chinn DJ, Kumar N. Identification of women with reduced bone density at the lumbar spine and femoral neck using BMD at the os calcis. Osteoporos Int 2000;11:797–802. 4. Miller PD, Njeh CF, Jankowski LG, Lenchik L. International Society for Clinical Densitometry Position Development Panel and Scientific Advisory Committee. What are the standards by which bone mass measurement at peripheral skeletal sites should be used in the diagnosis of osteoporosis? J Clin Densitom 2002;5Suppl.:S39–45. 5. Michalska D, Zikan V, Stepan J, Weichetova M, Kubova V, Krenkova J, et al. X-ray densitometry and ultrasonography of the heel bone – sensitivity and comparison with densitometry of the axial skeleton. Cas Lek Cesk 2000;139:231–6. 6. Genant HK, Grampp S, Glueer CC, Faulkner KG, Jergas M, Engelke K, et al. Universal standardisation for dual x-ray absorptiometry: patient and phantom cross-calibration results. J Bone Miner Res 1994;9:1503–14. 7. Pearson J, Dequeker J, Reeve J, Felsenberg D, Henley M, Bright J, et al. Dual x-ray absorptiometry of the proximal femur: normal European values standardized with the European spine phantom. J Bone Miner Res 1995;10:315–24. 8. Formica CA. Standardization of BMD measurements (editorial). Osteoporos Int 1998;8:1–3. 9. Grigorian M, Shepherd JA, Cheng XG, Njeh CF, Toschke JO, Genant HK. Does osteoporosis classification using heel BMD agree across manufacturers? Osteoporos Int 2002;13:613–7. 10. Faulkner KG, Roberts LA, McClung MR. Discrepancies in normative data between Lunar and Hologic DXA systems. Osteoporosis Int 1996;6:432–6. 11. Leib ES, Lewiecki EM, Binkley N, Hamdy RC. Official positions of the International Society for Clinical Densitometry. J Clin Densitom 2004;7:1–6.


Heel bone densitometry 12. Blake G, Chinn D, Steel S, Patel R, Panayiotou E, Thorpe J, et al. The Revised NOS Position Statement on peripheral X-ray Absorptiometry: A listing of device specific T-score thresholds for the clinical interpretation of pDXA examinations. NOS, 2005.


13. Royal College of Physicians. Osteoporosis: Clinical guidelines for prevention and treatment. London: RCP, 1999;7:66. 14. Williams ED, Daymond TJ. Evaluation of calcaneus bone densitometry against hip and spine for diagnosis of osteoporosis. Br J Radiol 2003;76:123–8.

467


The impact of a short course of study on the performance of radiographers when highlighting fractures on trauma radiographs: ‘‘The Red Dot System’’ S J MACKAY,

TDCR, MSc, PhD

Salford University, Frederick Road Campus, Salford, Manchester M6 6PU, UK ABSTRACT. This study was set up to determine the effect of a short course of study on radiographers’ ability to identify fractures in a trauma context (Red Dot system). The performance of 133 radiographers attending Red Dot courses delivered from April 1999 to September 2003 was evaluated prospectively using a selected sample of radiographs. Assessment of fracture identification occurred before, immediately after and 6 months after the course. The same film set was used and consisted of 30 axial and appendicular cases, 18 with fractures and 12 normal or normal variants. Following a test for normality of the data, a Wilcoxon Signed-Rank Test was selected and paired tests were done between each assessment for sensitivity and specificity. There were significant differences in sensitivity between all three assessments (p#0.05), the pre-course/postcourse and the post-course/6 month comparison being highly significant (p#0.01). Specificity showed significant differences between the pre-course/post-course (p#0.01) and the post-course/6 month follow-up scores (p#0.05), but no difference was found between the 6 months/pre-course scores. The participants improved their ability to identify fractures and this appears to be as a result of the course. This improvement was not demonstrable after 6 months, although only 30% of participants took part in this follow-up. Radiographers can improve their image interpretation skills from a short course of study, but probably need continuing professional development to maintain these skills.

The Health Professions Council standards of proficiency for radiographers [1] are designed to ensure the safe and effective practice of radiographers in the UK. They state that part of the role of the radiographer is to be able to ‘‘distinguish between normal and abnormal appearances evident on images’’ [1]. The Quality Assurance Agency benchmark statements [2], which also describe the role of the radiographer, but from an educational perspective, are more precise on this matter and state that radiographers should be able to ‘‘recognize and respond appropriately to abnormal, aberrant and pathological appearances on radiographic images’’ [2]. In an accident and emergency setting, this practice is termed a Red Dot system [3] and radiographers use these skills to alert the casualty officer to fractures. It is now over 30 years since Swinburne [4] suggested that radiographers could be used to distinguish between normal and abnormal radiographs, and such trauma flagging systems were reported in 2000 to be in operation in over 85% of accident and emergency departments in the UK (excluding Northern Ireland) [5]. Several authors have provided evidence of the potential of radiographers to perform this role with training [5, 6] and without [7]. However, these evaluations have limited external validity. This is mainly due to methodological limitation. These studies used either small numbers (n57) in one department [6], a small non-random national sample (n522) [5], or radiographers untrained in triaging 468

Received 8 March 2005 Revised 7 October 2005 Accepted 31 October 2005 DOI: 10.1259/bjr/53513558 ’ 2006 The British Institute of Radiology

casualty radiographs [7]. Many studies have also purposefully selected senior or experienced radiographers as subjects and cannot therefore provide evidence on the performance of the range of radiographers performing Red Dot or similar systems in the UK today. A short course of study in image interpretation to enable radiographers to highlight abnormal trauma radiographs has been delivered at the author’s university since 1999 and has attracted radiographers of varying experience from throughout the UK. Their performance has been evaluated in an attempt to add to the evidence base in radiographer’s performance in Red Dot systems. One study [4] selected senior or experienced radiographers, the assumption presumably being that their experience of image interpretation in their role as a radiographer will have an effect on their film viewing performance. This study will also investigate the relationship between a radiographer’s years of experience and performance in a Red Dot scheme.

Method Overview The performance of radiographers attending one of seven Red Dot courses delivered from April 1999 to September 2003 was evaluated prospectively using a selected sample of radiographs. They were assessed The British Journal of Radiology, June 2006

Impact of a short course on radiographers’ ability to detect fractures

before the start of the course, immediately after the course and then after 6 months of clinical practice. The same film series was used for each evaluation.

The course This was a 2 day course with short keynote lectures and small group tutorials delivered by radiologists and reporting radiographers. Topics covered were the identification of fractures and some non-traumatic pathologies of the axial and appendicular skeleton along with normal and abnormal variants. Both adult and paediatric patients were included so as to represent the range of clinical work that might present to an accident and emergency department.

The sample The 133 radiographers from across the UK self-selected for the courses. Their clinical experience ranged from newly qualified radiographers to those with 36 years of experience (mean512.1 years, SD510.4). 36 (27.0%) of the sample had no experience of working within a Red Dot system before and, of the 97 (73%) who had, the maximum number of years was 20 years (mean53.1 years, SD53.4). Consent was obtained from the participants for the use of their test scores, which form part of the course feedback, in this research and it was explained that the data would be anonymized so no individuals could be identified.

The film assessment The sample of radiographs were selected by a consultant radiologist and a reporting radiographer who aimed to make them representative of the range and quality of films that might pass through an accident and emergency department at a district general hospital. There were 30 cases of which 18 were pathological and 12 normal or normal variant. The pathological films had been reported upon by a consultant radiologist in practice and the consultant and reporting radiographer who selected the film were also agreed on the diagnosis. 10 radiographs were of the upper and nine of the lower limbs, seven spine and pelvis, and four skull and face. Three were paediatric cases. The assessment was undertaken at the university under controlled conditions. Radiographs were viewed in one session and each radiographer was allowed to view a case for 90 s before recording one of two answers, either normal (which included normal variants) or abnormal. For the 6 month follow-up assessment, radiographers were invited back to the university to complete the assessment. Those who were unable to do this were sent a CD containing the digitized film set along with instructions regarding the conditions in which to view the radiographs, but this could not be controlled.

Analysis True and false, and positive and negative scores were calculated to determine sensitivity, specificity, The British Journal of Radiology, June 2006

and positive and negative predictive values. Accuracy values were not used, following the advice of Maisey and Hutton [8] who describe this as a global parameter, ‘‘…which is often misleading as a measure of test performance and is rarely of any value’’ [8]. The sensitivity and specificity data obtained was inspected for normality using a Kolmogorov-Smirnov statistic [9]. This showed that the distributions were significantly different from the normal distribution for the pre-course and post-course (p,0.01) sensitivity data and the pre-course (p,0.05) and post-course (p,0.01) specificity data. No significant differences were found for the 6 month data for either variable. Therefore data were treated as non-parametric. Comparison was required of pairs of scores in relation to the course, e.g. pre/post-course; so the non-parametric Wilcoxon Signed-Rank Test [9] was selected for analysis. All analyses were performed using the Statistical Package for the Social Sciences (SPSS v11.5; SPSS, Chicago, IL) and a 0.05 significance level. The number of years subjects had been qualified and the number of years they had spent working within a Red Dot scheme were correlated (Spearman’s Rho, p# 0.05) with sensitivity and specificity values to examine the effect of experience on image interpretation performance.

Errors False negative errors constitute missed fractures and could be considered as the more important type of detection error, which could lead to significant morbidity and/or legal liability. Therefore, the most commonly occurring cases of false negatives were highlighted and described. This will help to inform course curricula and enable radiographers to focus their professional development.

Results Response rate 133 (100%) radiographers completed the pre-course assessment and 132 (99.2%) the post-course assessment. 39 (29.3%) radiographers completed the 6 month follow up with 10 of these (25.6% of the responders at 6 months) doing so using the digitized film set.

Sensitivity and specificity The results are presented in Tables 1 and 2, and box plots (using SPSS v11.5) are presented in Figures 1 and 2. The positive and negative predictive values were not analysed using inferential statistics and are presented for comparative purposes only. The median scores for sensitivity show that the radiographers achieved a score of 78.9% initially. This improved to 88.2% following the course, but then fell away to less than the pre-course score after 6 months. All differences were statistically significant. 469

S J Mackay

post-course and 6 month follow-up. Visual inspection of the box plot in Figure 2 demonstrate that the radiographers did worse immediately after the course, with more scores below the median. They then improved slightly at 6 months but were not significantly different from the pre-course score.

Table 1. Breakdown of raw scores Median

Sensitivity Pre-course Post-course 6 months

Interquartile range

25

75

78.9 88.2 76.5

70.6 82.4 64.7

88.2 94.1 88.2

76.9 76.9 76.9

69.2 62.0 69.2

85.0 85.0 86.0

86.0 84.2 84.4

81.0 79.0 79.0

93.0 89.1 92.0

69.6 77.8 65.7

62.2 70.3 56.0

79.0 88.0 78.0

Specificity

6 month follow-up

Pre-course Post-course 6 months

Further analysis was undertaken as less than one third of the subjects completed the 6 month follow-up. This investigated whether this group of subjects (the completers) had scored differently at other assessment times to those who did not complete the assessment. This would help to clarify whether the completers were representative of the group as a whole. A Wilcoxon test found no statistically significant differences between the completers and non-completers for number of years practicing radiography or the number of years in a Red Dot scheme. The differences in the pre- and post-test scores for these groups are displayed in Table 3.

Positive predictive value Pre-course Post-course 6 months Negative predictive value Pre-course Post-course 6 months

The median scores for specificity were the same, but the distributions varied and showed a highly statistically significant difference between the pre-course and postcourse tests, and significant differences between the

Effect of experience No significant correlations were found between the number of years since qualification or length of time in a

Table 2. Wilcoxon Signed-Rank Test results for sensitivity and specificity Pre-course/post-course

Post-course/6 months

Pre-course/6 months

Sensitivity

Median change Median difference Z p,

9.3 5.9 –7.049 0.01

–11.7 –11.8 –4.190 0.01

–2.4 0 –1.984 0.05

Specificity

Median change Median difference Z p,

0 –7.7 –4.135 0.01

0 7.7 –1.983 0.05

0 0 –0.587 NS

NS, not significant.

Figure 1. Sensitivity changes over time. Box plot where the box length is the interquartile range with the median identified as a thick black line traversing the box. Outliers (ringed) are cases with values between 1.5 and 3 box lengths from the upper or lower edge of the box. From SPSS v11.5. 470


Impact of a short course on radiographers’ ability to detect fractures

Figure 2. Specificity changes over time. Box plot where the box length is the interquartile range with the median identified as a thick black line traversing the box. Outliers (ringed) are cases with values between 1.5 and 3 box lengths from the upper or lower edge of the box. From SPSS v11.5.

Red Dot scheme and sensitivity or specificity scores precourse, post-course or 6 months after the course.

False negative errors The 10 most common false negative errors before and immediately after the course were ranked and collated (11 cases appear in column 1 as there were 2 cases of 10.9% tied for 10th place). These key errors are where further learning should be focused. These can be found in Table 4.

Discussion The ability of the radiographers to identify the abnormal trauma radiographs, sensitivity, increased immediately after the course, suggesting that they had learned from the course. This improvement declined over 6 months and ended up being worse than before they started the course. The ability of the radiographers to identify normality on trauma radiographs, specificity, decreasd after the course and then improved slightly, but was then no different from the original scores. This ‘‘post-course dip’’ in score was also identified by another author [5] and appears to be a phenomenon inherent in the image interpretation learning process. This may be due in part Table 3. Wilcoxon Signed-Rank Test results completers and non-completers Completers pre-course/ non-completers pre-course

Completers post-course/ non-completers post-course

Sensitivity

Z p,

–1.733 NS

–2.646 0.01

Specificity

Z p,

–2.120 0.05

–1.169 NS

NS, not significant.


to the course, the main bulk of which shows students how to recognize fractures. The students may then have an immediate tendency to look for the abnormal rather than the normal. A similar study [5] followed up subjects for 6–10 weeks and found a decrease in sensitivity which, although not statistically significant, would support the trend found in this work suggesting that improvements in performance were short lived. This suggests that radiographers need to continue to acquire the knowledge and skills necessary for interpreting images following any course of study. Based on these data, they should be topping up their knowledge and skills before a period of 6 months has elapsed. In light of the impending mandatory continuing professional development (CPD) requirements of the Health Professions Council, radiographers will need to consider undertaking a range of CPD activities to consolidate and enhance their image interpretation skills. Table 4. Most commonly occurring false negative errors Case description

Skull fracture Fracture of head of 5th metatarsal with Os Peroneum Greenstick fracture of tibia Slipped upper femoral epyphysis Sacral fracture Fracture radius Fracture radial head Subtle fracture base of 1st metacarpal Fracture humeral head Fracture lunate Fracture base of 1st distal phalanx foot Fracture zygoma

Proportion of subjects making false negative errors per case (%) Pre-course

Post-course

50.0 40.2

63.0 17.4

39.1 38.0

32.6 10.9

27.2 22.8 18.5 17.4

20.7 7.6 12.0 0

16.3 10.9 10.9

10.9 25.0 0

0

16.3

471

S J Mackay Table 5. Comparison of pre-training sensitivity and specificity levels with other authors

Loughran [10] McConnell [5] Hargreaves [6] Mackay (this study)

Sensitivity (%)

Specificity (%)

Year published

Measure

62.7 92 76.2 78.9

97.9 64 96.4 76.9

1996 2000 2003 2006

mean median mean median

It is acknowledged that the small sample that completed the 6 month assessment may not have been representative of the group as a whole. However, comparisons between the completers and non-completers suggest that on the measures of length of experience since qualification and working in a Red Dot scheme they were not significantly different. In addition, the differences in performance between these two groups were on the post-course sensitivity scores in which the completers did better. This would suggest that the scores from this study at 6 months may even have exaggerated the performance of the whole group of radiographers. The specificity scores between these groups were not significantly different after the course and are therefore likely to be representative. Comparison with other results [5, 6, 10] shows a range of sensitivity and specificity levels for radiographers (Table 5). Close statistical comparison is difficult as the measures of central tendency used are different. However, looking across the values it would appear that there is a need for radiographers to improve their performance in image interpretation within Red Dot protocols. No published performance levels were found for radiographers working in a Red Dot scheme. However, post-graduate reporting radiographers are expected to achieve 90–95% sensitivity and specificity levels [11]. If we apply this standard of performance to Red Dot radiographers, they are clearly falling short of this level. One notable trend when comparing the data in this study with two others [5, 6] is that subjects in one [5] tended to over report (high sensitivity and low specificity) and in the other [6] tended to under report (low sensitivity and high specificity). In the current study, relative to these other two, the subjects maintained a more stable performance over both measures. The areas where radiographers missed fractures have been highlighted in this study and could be added to those from another study [6]. These data could help to inform ‘‘Red Dot’’ course developers and evaluators to take a more evidence based approach to radiographer education. Experience did not appear to be a factor in the performance of this group of radiographers. Further research is needed to explore the phenomenon of ‘‘experience’’ in radiography and try to tease out the factors that might affect the performance of radiographers in image interpretation.

472

Conclusions The radiographers did appear to learn from the short course and improved their levels of performance in image interpretation. These improvements were short lived and it is recommended that regular updating is undertaken by radiographers working within a Red Dot scheme. There appeared to be a reduction in performance immediately following the image interpretation course, which caused an increase in false positives. Radiographers should be aware of this and attempt to counteract the effect. Programme evaluators should factor in this effect when planning any assessments of image interpretation performance after a short course of study. The scores achieved in this evaluation fall well short of the 90–95% sensitivity and specificity levels used in radiographer reporting. Radiographers would benefit from continuing to improve their image interpretation skills throughout their working lives.

References 1. Health Professions Council. Standards of Proficiency: Radiographers. London, UK: HPC; 2003. 2. The Quality Assurance Agency for Higher Education. Benchmark Statement for Radiographers. QAA; 2001. 3. Berman L, de Lacey G, Twomey E, Twomey B, Welch T, Eban R. Reducing errors in the accident and emergency department :a simple method using radiographers. Br Med J 1985;290:421–2. 4. Swinburn K. Pattern recognition for radiographers. Lancet 1971;1:589–90. 5. McConnell JR, Webster AJ. Improving radiographer highlighting of trauma films in the Accident and Emergency department with a short course of study – an evaluation. Br J Radiol 2000;73:608–12. 6. Hargreaves J, Mackay S. The accuracy of the red dot system: can it improve with training? Radiography 2003;9:283–9. 7. Renwick I, Butt W, Steele B. How well can radiographers triage x-ray films in accident and emergency departments? Br Med J 1991;302:568–9. 8. Maisey MN, Hutton J. Guidelines for the evaluation of radiological technologies. Report No.: 13874. London, UK: British Institute of Radiology, 1995. 9. Pallant J. SPSS survival manual (SPSS v10). 1st edn. Buckingham, UK: O.U. Press; 2001. 10. Loughran CF, Raynor J, Quine M, Mulley A. The red dot system: how good is it? In: Radiology UK; May 1996. 11. Salford University. MSc Advanced Practice (Clinical Radiographic Reporting) Programme Handbook. 2004.



Portal vein embolisation prior to hepatic resection for colorectal liver metastases and the effects of periprocedure chemotherapy 1

I K BEAL, MBBS, MRCP, FRCR, 1S ANTHONY, MBBS, MRCP, FRCR, 1A PAPADOPOULOU, MBBS, FRCR, 2 R HUTCHINS, MBBS, MS, FRCS, 2G FUSAI, FRCS, 3R BEGENT, MBBS, MRCS, LRCP, 1N DAVIES, MBBS, FRCS, 1 J TIBBALLS, MRCP, FRCR, FRANZCR and 2B DAVIDSON, MBChB, MD, FRCS

FRCR,

Departments of 1Radiology 2HPB Surgery and 3Medical Oncology, Royal Free Hospital NHS Trust and Royal Free and University College School of Medicine, Pond Street, London NW3 2QG, UK ABSTRACT. Portal vein embolisation (PVE) is an effective method of increasing future liver remnant (FLR) but may stimulate tumour growth. The effect of periprocedure chemotherapy has not been established. 15 consecutive patients underwent PVE prior to hepatic resection for colorectal liver metastases with a FLR ,30% of tumour-free liver (TFL). Liver and tumour volumes pre-PVE and 6 weeks post-PVE were calculated by CT or MRI volumetry and correlated with the periprocedure chemotherapy regimen. PVE increased the FLR from 18¡5% of TFL to 27¡8% post-PVE (p,0.01). Post-PVE chemotherapy did not prevent hypertrophy of the FLR but the volume increase with chemotherapy (median 89 ml, range 7–149 ml) was significantly reduced (median 135 ml, range 110–254 ml without chemotherapy) (p50.016). Tumour volume (TV) decreased in those receiving post-PVE chemotherapy (median TV decrease 8 ml, range 277 ml to +450 ml) and increased without chemotherapy (median TV increase 39 ml, range 258 ml to +239 ml). Of the 15 patients, eight underwent resection; four were not resected due to disease progression and three due to insufficient hypertrophy of the FLR. PVE increased the FLR by an average of 9% allowing resection in 50% of patients. Periprocedure chemotherapy did not prevent but did reduce hypertrophy. A trend towards tumour regression was observed.

Liver resection surgery provides the main possibility of cure in patients with colorectal liver metastases. 5-year survival ranges from 25% to 40% [1–4]. One of the contraindications to hepatic resection is a small future liver remnant (FLR), which usually arises in patients with congenitally small left lateral segments (II and III) who require an extended right hepatectomy. The small residual liver volume increases the risk of post-operative liver failure. A FLR of 25–30% of non-tumour liver volume has been recommended in patients with normal liver function who are to undergo hepatic resection [3, 5– 8], although resections have been carried out successfully with a FLR of ,20% [9]. Selective portal vein embolisation (PVE) can produce atrophy of the segments affected by the cancer and compensatory hypertrophy of the contralateral segments [6–8]. The technique was first developed in patients with hilar cholangiocarcinoma where extended right hepatectomy is often required [10]. Although residual liver volumes can be increased, there remains some controversy as to whether this results in improved post-operative outcome [11]. There are now several reports in the literature of PVE applied to patients with colorectal metastases [6–8, 12–14]. One of the potential difficulties with using PVE in metastatic colorectal cancer is that tumour growth may be promoted by PVE. Kokudo and colleagues [15] studied patients undergoing PVE prior to hepatic resection for colorectal liver metastases and found that PVE increased tumour growth in the embolised The British Journal of Radiology, June 2006

Received 26 April 2005 Revised 2 October 2005 Accepted 14 November 2005 DOI: 10.1259/bjr/29855825 ’ 2006 The British Institute of Radiology

segments, possibly increasing the incidence of recurrence following liver resection. This also risks tumours becoming unresectable during the interval between PVE and surgery. The growth of liver metastases in the non-embolised segments following PVE has also been shown to be greater than the hypertrophy rate of the normal liver parenchyma [16]. However, in a study of the long term survival following PVE, which included 41 patients with colorectal cancer (CRC) liver metastases, there was no evidence to suggest that patients whose surgery had been made possible by PVE were associated with a poorer long term survival [7]. PVE has also been applied to patients with hepatocellular carcinoma and has been associated with improved operative outcomes as well as long term survival [17, 18]. Chemotherapy has been shown to increase the survival of patients with unresectable liver metastases [1]. Downstaging chemotherapy using Oxaloplatinum and 5-Fluorouracil (5FU) allows some initially unresectable CRC liver metastases to be successfully resected [19–21]. The effect of chemotherapy administered after PVE on the FLR hypertrophy and on the tumour growth in embolised segments, in patients with colorectal liver metastases, has not been investigated. The aim of this study was to assess, first, whether chemotherapy prevents tumour progression between PVE and surgery and, second, whether chemotherapy has a detrimental effect on FLR hypertrophy. 473

I K Beal, S Anthony, A Papadopoulou et al

Patients and methods Between September 1999 and December 2002, a single surgical firm in a specialist tertiary referral centre performed right or extended right hepatectomy for colorectal liver metastases in 57 patients. Over this time period, 15 patients who required a right or extended right hepatectomy to resect their metastases were found to have a calculated FLR of less than 30% in an otherwise normal liver (n514), or less than 40% (n51) in a diseased liver (imaging findings consistent with steatosis). These patients underwent right PVE. Nine patients had synchronous and six metachronous colorectal liver metastases. There were six females and nine males, mean age 65 years (range 52– 74 years). The stage of the original tumour was Dukes’ B in 4 and Dukes’ C in 11. The mean number of metastases was 4 (range 1–11). Three of the 15 patients had undergone resection for metastases in the left lobe of liver prior to PVE. Three patients had minor extension of tumour into segment IV, but no patient had discrete metastases in nonembolised liver.

Chemotherapy Following initial diagnosis and bowel cancer resection, adjuvant chemotherapy had been carried out in the referring hospital in 11 of the 15 patients (Dukes’ C n59, Dukes’ B n52). In all cases this comprised a standard 5FU and folinic acid based protocol. In five of these patients, this had been completed between 6 months and 18 months prior to PVE. In six patients (synchronous metastases n55, metachronous metastases n51) chemotherapy was continued until 1 month prior to PVE (5FU/FA (n52) or 5FU/Oxaliplatin (n54)). In the remaining four patients no chemotherapy was administered following bowel cancer resection. Following PVE, eight patients who had already received chemotherapy underwent a further course of 5FU/FA (n51), 5FU/ FA+oxaliplatin (n55) or 5FU/FA+Irinotecan (n52). Two patients who had not previously had chemotherapy were given post-PVE chemotherapy with 5FU/FA +oxaliplatinum (n51), 5FU/FA alone (n51). Chemotherapy post-PVE was started at 2 weeks postembolisation and continued for 6 weeks. Three patients who had received chemotherapy post-bowel resection or prior to embolisation did not receive post-PVE chemotherapy. Two patients remained chemo-naive. For the purpose of analysis, the patients were grouped into post-PVE chemotherapy (n510) and no chemotherapy post-PVE (n55).

Volumetric analysis Liver volume measurements of the whole liver (WL), segments I-III or I-IV (FLR) and tumour (TV) were performed by CT (n52) (GE Medical Systems High speed system, Milwaukee, WI) or MRI (n513) using Philips Intera 1 T system (Philips, Best, Netherlands). Total functioning liver (TFL) was measured by subtracting TV from WL (Philips Easyvision workstation [release 4.3]) and GE Medical Systems Advantage Windows 2.0 474

3D Analysis). The technique for measuring volumes by CT has been described elsewhere [5, 7]. Briefly, axial 10 mm sections through the liver were obtained in a single breath-hold. Using a workstation, the areas of TV, WL and FLR were calculated by multiplying the area of each liver image by the slice thickness. The percentage fraction of FLR was calculated: FLR/TFL6100. The measurements performed by MRI involved the same process using the scanner software and a similar technique as described by Caldwell et al [22]. An experienced Hepatobiliary Radiologist performed all measurements. The liver volume measurements were repeated approximately 6 weeks post-PVE.

Portal vein embolisation (PVE) This was performed by a percutaneous transhepatic approach under general anaesthetic (n52) or local anaesthesia and intravenous sedoanalgesia (n513). An ultrasound-guided percutaneous portal vein puncture was performed using a Neff set (William Cook Europe A/S, Bjaeverskov, Denmark). A left-sided approach was preferred (n59) but when not possible due to the small size of the left lobe a right-sided approach was used (n56). 5 Fr or 6 Fr sheaths were introduced and the branches of the right portal vein were embolised sequentially using 4 Fr cobra or sidewinder shaped catheters. Embolisation was performed with alcohol (n53) and with/without 500 mm polyvinylalcohol (PVA) particles (William Cook Europe A/S, Bjaeverskov, Denmark) (n52/1, respectively) or enbucrilate tissue adhesive (Histoacryl; B Braun Medical AG, Emmenbrucke, Switzerland) and Lipiodol (Guerbet, Roissy, France) at an enbucrilate:Lipiodol ratio of 1:3– 1:5 (n512). Prior to PVE and daily for 2 days post-PVE, full blood count, clotting profile and liver function tests were measured.

Statistics Paired and two sample equal variance t-tests were used for statistical analysis. Data are expressed as means (¡ 1 SD), median values with range and 95% confidence intervals (CIs). p,0.05 was considered significant.

Results Procedure of portal vein embolisation (PVE) Embolisation of the right portal vein was technically successful in all 15 patients with no immediate complications. The first three patients were embolised with alcohol, which produced severe right upper quadrant pain. Subsequent embolisations were performed using enbucrilate. Post-embolisation, most patients had mild discomfort and low-grade pyrexia, which settled within 24 h. In one patient the 6-week follow-up MRI demonstrated complete portal vein thrombosis. This had not been recognized at the time of the procedure and his clinical course post-PVE was unremarkable. A CT scan showed evidence of enbucrilate:Lipiodol within the The British Journal of Radiology, June 2006

Portal vein embolisation and the effects of periprocedure chemotherapy Table 1. Effect of portal vein embolisation (PVE) on liver volumes

Mean WL (ml) Median WL (range) (ml) Mean TV (ml) Median TV (range) (ml) Mean FLR (ml) Median FLR (range) (ml) Mean FLR/TFL%

Table 2. The effect of chemotherapy on liver hypertrophy Patient Chemo regimen

Pre-PVE

Post-PVE

p

1912 1739 (1222–3114)

1761 1631 (1160–3398)

0.8

235 123 (29–1147)

262 139 (10–1597)

0.71

270 281 (110–558)

400 404 (245–653)

,0.01

18(¡5)

27(¡8)

,0.01

WL, whole liver volume; TV, tumour volume; FLR, future liver remnant volume; TFL, tumour free liver volume.

proximal left portal vein branch. This patient was not excluded from analysis.

Effect of portal vein embolisation (PVE) on future liver remnant (FLR) volume Following PVE, there was no significant change in whole liver (WL) volume (median volume pre-PVE 1739 ml (range 1222–3114 ml) to a median volume post-PVE 1631 ml (range 1160–3398 ml)). Volumetric analysis was carried out prior to PVE and was repeated at a median of 6 weeks post-PVE (range 5– 12 weeks). Prior to PVE, the future liver remnant (FLR) volume ranged from 110 ml to 558 ml (median 281 ml), which represented 18(¡5)% of the tumour free liver (TFL). The volume was significantly increased following PVE to a FLR range of 245–653 ml (median 404 ml) which represented 27(¡8)% of the TFL (p50.0089). The mean increase in parenchymal volume was 123 ml, which represented a mean increase in FLR/TFL ratio of 9% (Table 1). The one patient who had no significant increase in FLR post-PVE was found to have complete portal vein thrombosis.

Effect of portal vein embolisation (PVE) on tumour volume (TV)

Pre-PVE Pre-PVE FLR/TFL% FLR vol. ml

Post-PVE Post-PVE FLR/TFL% FLR vol. ml

16.5 20 15 22 23 24 12 30 17.5 15

297 237 182 423 303 330 258 558 344 254

25 31 25 30 31 31 19.5 44 22 15

395 320 256 533 418 413 407 653 398 261

16 19 22 16 7.5

293 245 281 235 110

26 33 39 21 17

404 421 535 345 245

Post-PVE chemo

1 2 3 4 5 6 7 8 9 10 No post-PVE chemo

11 12 13 14 15

Chemo regimen, chemotherapy regimen; PVE, portal vein embolisation; FLR, future liver remnant; TFL, tumour free liver.

median5135 ml, range5110–254 ml) (p50.016). The difference between the means was 70 ml (CIs526–115 ml). The effects of chemotherapy on FLR hypertrophy are illustrated by Figures 1 and 2.

Effect of chemotherapy on tumour volume (TV) The effect of chemotherapy on TV following PVE is shown in Table 3. There was a median TV decrease of 8 ml (range –77 ml to +450 ml) in those receiving post-PVE chemotherapy (n510) and a median TV increase of 39 ml (range –58 ml to +239 ml) in those who did not receive post-PVE chemotherapy (p50.476). The difference between the means was 58 ml (CIs5–82–198 ml). Of the three patients who had tumour extending into segment IV, two showed no significant change in TV

The patient group had a median TV of 123 ml (range 29–1147 ml) prior to PVE and 139 ml following PVE (range 10–1597 ml). This difference was not statistically significant (p50.71). Of the 15 patients 8 had TV increase and 7 had a TV reduction.

Effect of chemotherapy on liver hypertrophy (future liver remnant (FLR) volume) The effect of chemotherapy on the hypertrophy of the FLR following PVE is shown in Table 2. There was a significant increase in FLR following PVE in both groups. Patients who received chemotherapy post-PVE (n510) had less hypertrophy (increase in FLR volume median589 ml, range 7–149 ml) than those who received no chemotherapy post-PVE (n55) (increase in FLR volume The British Journal of Radiology, June 2006

Figure 1. Change in FLR volume after embolisation in all patients. FLR, future liver remnant volume; Pre, volume prior to PVE; Post, volume after PVE. 475


(AST) in all patients (p50.04) and bilirubin in 12 patients (p50.0007). An increase in international normalized ratio (INR) was seen in eight patients (p50.019) and a decrease in platelets noted in 13 patients (p50.002). A transient increase in white cell count (WCC) was noted in 10 patients (p50.009).

Surgery following portal vein embolisation (PVE)

Figure 2. Change in FLR volume after embolisation, according to chemotherapy administered. FLR, future liver remnant volume; Pre, volume prior to PVE; Post, volume after PVE; Post-PVE Chemo, chemotherapy administered after PVE; No post-PVE Chemo, no chemotherapy administered after PVE.

post-PVE (no post-PVE chemotherapy n51, post-PVE chemotherapy n51) and one had an overall increase in TV but no significant change of the TV in segment IV (no post-PVE chemotherapy). The effects of chemotherapy on TV are illustrated by Figure 3. Figures 4 and 5 demonstrate typical MRI findings prePVE and post-PVE in a patient receiving post-PVE chemotherapy with hypertrophy of the FLR and tumour volume regression.

Haematology and biochemistry Following PVE there were significant but transient increases in the levels of aspartate aminotransferase Table 3. Chemotherapy regimen and tumour volume following PVE Patient

Chemo regimen

Pre-PVE TV ml

Post-PVE TV ml

236 64 232 203 46 29 123 179 1147 29

228 32 155 242 54 10 116 139 1597 34

727 90 66 266 120

966 32 77 486 159

Of the 15 patients who underwent PVE who would otherwise have been deemed unresectable due to their small-anticipated FLR, eight proceeded to a right or extended right hepatectomy (53%). Four patients had disease progression on their post-PVE imaging at 6 weeks, which precluded potentially curative resection (extrahepatic disease progression, n53 and intrahepatic and extrahepatic disease progression, n51). Of these four, two had received post-PVE chemotherapy and two had not received chemotherapy following embolisation. Three patients had insufficient hypertrophy of the FLR. Of the eight who had surgery, there were no operative or post-operative mortalities.

Follow up Five of the eight patients undergoing surgery remain well and free of local or distant disease at a median of 18 months post-liver resection (range 14–26 months). Two patients developed distant metastatic disease and the third has undergone abdominoperineal resection for local recurrence of a rectal cancer. The three patients who had insufficient hypertrophy of the FLR, including the first patient with complete portal vein thrombosis, underwent palliative chemotherapy. The second of the three patients had an initial FLR of just 7.5% of the TFL. Despite a 123% increase in FLR this was still only 17% of TFL at 6 weeks post-PVE. The third patient had an extremely large tumour prior to embolisation, measuring 1147 ml. The percentage increase in FLR in this patient was just 16%. All three patients were included in the analysis.

Post-PVE chemo

1 2 3 4 5 6 7 8 9 10 No post-PVE chemo

11 12 13 14 15

Chemo regimen, chemotherapy regimen; PVE, portal vein embolisation; TV, tumour volume.

476

Discussion PVE is one of the major recent developments to facilitate resection in patients who would otherwise be contraindicated for liver surgery due to a small-anticipated residual liver volume. Patients in this study all had an anticipated FLR of 30% or less of tumour-free liver and were therefore deemed at risk of post-operative liver failure. Alcohol was initially used for embolisation, but due to severe alcohol-induced pain requiring general anaesthetic, subsequent embolisations were performed using enbucrilate. This was well tolerated and produced results comparable with that in other series. A literature review has suggested no clear advantage to any specific embolic substance [5]. One patient in the series developed complete portal vein thrombosis. In retrospect, a number of factors may have contributed to this complication. Due to the small left lobe The British Journal of Radiology, June 2006

Portal vein embolisation and the effects of periprocedure chemotherapy

Figure 3. Change in tumour volume after embolisation, according to chemotherapy administered. Pre, volume prior to PVE; Post, volume after PVE; Post-PVE Chemo, chemotherapy administered after PVE; No post-PVE Chemo, no chemotherapy administered after PVE.

volume, the approach was right-sided and, on completion of embolisation, the catheter was withdrawn so no postprocedure venogram was obtained. This does not cause the complication, but prevented it being recognized at the time of the procedure. Embolisation of the left portal vein may occur if too large a volume of enbucrilate:Lipiodol is injected, i.e. due to operator error. This is not thought to have occurred in this case. Anatomical variants such as portal vein trifurcation, present in this case, may predispose to left portal vein embolisation. The ability of PVE to increase the FLR was confirmed in the present study with a mean percentage increase in FLR/ TFL ratio of 9%. This figure correlates well with the previous reports of 8–13% [5, 7, 8]. 50% of the patients undergoing PVE were able to proceed with potentially curative resection. In the series reported by Elias and colleagues, 88% of patients proceeded to resection [7]. This difference in resection rate may relate to the centre’s selection of patients for PVE. In the present series the mean pre-PVE FLR/TFL ratio was 18%. This is lower than the 19–32% reported by other centres [5, 7]. Four of the patients in the present series had a FLR/TFL ratio of 15% or less prior to PVE and only one of these underwent resection. If a minimum FLR of 30% is to be achieved and the percentage increase with PVE is in the order of 10%, then patients with an FLR of less than 20% may not benefit from PVE. The minimal remnant volume required for safe resection, however, remains controversial and may be

Figure 4. MRI demonstrates axial section through the liver prior to portal vein embolisation.


Figure 5. Post-embolisation: MRI demonstrates future liver remnant volume (FLR) hypertrophy and tumour volume regression.

increased by any underlying parenchymal abnormality, such as steatosis. One series concluded that the hypertrophy of FLR induced by PVE had no beneficial effect on the post-operative outcome in patients with normal liver [11]. However, in this series, patients were undergoing a standard right hepatectomy with a satisfactory preembolisation FLR/TFL volume ratio of 33¡10%. The same study concluded that in patients with chronic liver disease, pre-operative PVE significantly decreased postoperative complications. Another study has confirmed an association between a small FLR and an increase in post-operative complications [9]. Those patients with a standardized FLR of 20% or less had an increase in postoperative morbidity, but not mortality. A FLR of 40% has been recommended in patients who have previously undergone high dose chemotherapy [12]. A time interval between PVE and hepatic resection is necessary to allow parenchymal hypertrophy. Although the FLR growth is maximal within the first 2 weeks [5, 8] it continues for 4–5 weeks following PVE [5, 6, 8]. For this reason, in the anticipation of maximum hypertrophy, this study allowed 6 weeks between PVE and repeat volume measurement. The delay of 12 weeks before repeat volume measurement in one patient was due to external circumstances and subsequent surgical resection was successful. This is the first study to examine the use of chemotherapy in patients with colorectal liver metastases undergoing pre-operative PVE. Chemotherapy has usually been discontinued prior to PVE [7, 12] to encourage hypertrophy of the remnant liver parenchyma. In our study, chemotherapy was discontinued at least 1 month prior to PVE and it has been assumed that effects of this chemotherapy on post-PVE liver regeneration and oncogenesis would be minimal or insignificant. Chemotherapy in these circumstances is unproven. The particular chemotherapy regimen was administered according to individual oncologist practice. There was no significant difference in disease stage between the two groups. Due to the small study size, only the effects of chemotherapy post-PVE have been analysed. In the present study FLR hypertrophy occurred whether or not chemotherapy was given post-PVE. The hypertrophy was, however, significantly less in those patients who received post-PVE chemotherapy suggesting that although chemotherapy is not contraindicated, it 477


should be considered carefully in those patients who require a large compensatory hypertrophy following the procedure (FLR of ,20%). There is evidence to suggest PVE may stimulate tumour growth in both the embolised and non-embolised lobes of the liver. We have only assessed tumours in the embolised segments. Elias et al [16] reported increased growth of liver metastases compared with normal liver parenchyma following PVE in five patients with tumour in the non-embolised lobe. Kokudo et al [15] assessed tumour growth in the embolised lobe and found oncogenesis was promoted by PVE, with the tumour doubling time being reduced from 92 days to 76 days. In the present study, patients undergoing chemotherapy had no overall increase in tumour volume, which would suggest that chemotherapy did inhibit the tumour growth associated with PVE. Tumour volume regression was not demonstrated by the use of chemotherapy but this would seem unlikely over the short period between pre-PVE staging and re-assessment at 6 weeks post-PVE. The confidence interval for the data is wide, suggesting that further numbers would be helpful. This is the first study on the effects of peri-PVE chemotherapy and the relationship with FLR hypertrophy and tumour volume. Further studies are required to define the optimal chemotherapy regimen for patients with colorectal liver metastases who are undergoing PVE.

References 1. Scheithauer W, Rosen H, Kornek GV, Sebesta C, Depisch D. Randomised comparison of combination chemotherapy plus supportive care with supportive care alone in patients with metastatic colorectal cancer. BMJ 1993;306:752–5. 2. Fong Y, Salo J. Surgical therapy of hepatic colorectal metastasis. Semin Oncol 1999;26:514–23. 3. Fusai G, Davidson BR. Management of colorectal liver metastases. Colorectal Dis 2003;5:2–23. 4. Jaeck D, Bachellier P, Guiguet M, Boudjema K, Vaillant JC, Balladur P, et al. Long-term survival following resection of colorectal hepatic metastases. Association Francaise de Chirurgie. Br J Surg 1997;84:977–80. 5. Abdalla EK, Hicks ME, Vauthey JN. Portal vein embolization: rationale, technique and future prospects. Br J Surg 2001;88:165–75. 6. de Baere T, Roche A, Vavasseur D, Therasse E, Indushekar S, Elias D, et al. Portal vein embolization: utility for inducing left hepatic lobe hypertrophy before surgery. Radiology 1993;188:73–7. 7. Elias D, Ouellet JF, de Baere T, Lasser P, Roche A. Preoperative selective portal vein embolization before hepatectomy for liver metastases: long-term results and impact on survival. Surgery 2002;131:294–9. 8. Madoff DC, Hicks ME, Abdalla EK, Morris JS, Vauthey JN. Portal vein embolization with polyvinyl alcohol particles and coils in preparation for major liver resection for

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hepatobiliary malignancy: safety and effectiveness--study in 26 patients. Radiology 2003;227:251–60. Abdalla EK, Barnett CC, Doherty D, Curley SA, Vauthey JN. Extended hepatectomy in patients with hepatobiliary malignancies with and without preoperative portal vein embolization. Arch Surg 2002;137:675–81. Makuuchi M, Thai BL, Takayasu K, Takayama T, Kosuge T, Gunven P, et al. Preoperative portal embolization to increase safety of major hepatectomy for hilar bile duct carcinoma: a preliminary report. Surgery 1990;107:521–7. Farges O, Belghiti J, Kianmanesh R, Regimbeau JM, Santoro R, Vilgrain V, et al. Portal vein embolization before right hepatectomy: prospective clinical trial. Ann Surg 2003;237:208–17. Azoulay D, Castaing D, Smail A, Adam R, Cailliez V, Laurent A, et al. Resection of nonresectable liver metastases from colorectal cancer after percutaneous portal vein embolization. Ann Surg 2000;231:480–6. de Baere T, Roche A, Elias D, Lasser P, Lagrange C, Bousson V. Preoperative portal vein embolization for extension of hepatectomy indications. Hepatology 1996;24:1386–91. Imamura H, Shimada R, Kubota M, Matsuyama Y, Nakayama A, Miyagawa S, et al. Preoperative portal vein embolization: an audit of 84 patients. Hepatology 1999;29:1099–105. Kokudo N, Tada K, Seki M, Ohta H, Azekura K, Ueno M, et al. Proliferative activity of intrahepatic colorectal metastases after preoperative hemihepatic portal vein embolization. Hepatology 2001;34:267–72. Elias D, de Baere T, Roche A, Leclere J, Leclere J, Lasser P. During liver regeneration following right portal embolization the growth rate of liver metastases is more rapid than that of the liver parenchyma. Br J Surg 1999;86:784–8. Azoulay D, Castaing D, Krissat J, Smail A, Hargreaves GM, Lemoine A, et al. Percutaneous portal vein embolization increases the feasibility and safety of major liver resection for hepatocellular carcinoma in injured liver. Ann Surg 2000;232:665–72. Tanaka H, Hirohashi K, Kubo S, Shuto T, Higaki I, Kinoshita H. Preoperative portal vein embolization improves prognosis after right hepatectomy for hepatocellular carcinoma in patients with impaired hepatic function. Br J Surg 2000;87:879–82. Adam R, Avisar E, Ariche A, Giachetti S, Azoulay D, Castaing D, et al. Five-year survival following hepatic resection after neoadjuvant therapy for nonresectable colorectal [liver] metastases. Ann Surg Oncol 2001;8:347–53. Bismuth H, Adam R, Levi F, Farabos C, Waechter F, Castaing D, et al. Resection of nonresectable liver metastases from colorectal cancer after neoadjuvant chemotherapy. Ann Surg 1996;224:509–20. Elias D, Lasser P, Rougier P, Ducreux M, Bognel C, Roche A. Frequency, technical aspects, results, and indications of major hepatectomy after prolonged intra-arterial hepatic chemotherapy for initially unresectable hepatic tumors. J Am Coll Surg 1995;180:213–9. Caldwell SH, de Lange EE, Gaffey MJ, Sue M, Boyd JC, Dickson RC, et al. Accuracy and significance of pretransplant liver volume measured by magnetic resonance imaging. Liver Transpl Surg 1996;2:438–42.



Radiation benefit and risk at the assessment stage of the UK Breast Screening Programme 1

J LAW,

PhD

and 2K FAULKNER,

PhD

1

Edinburgh University Department of Medical Physics, Chancellor’s Building, Little France Crescent, Edinburgh EH16 4SB and 2Quality Assurance Reference Centre, 9 Kingfisher Way, Silverlink Business Park, Wallsend NE28 9ND, UK ABSTRACT. The balance between benefit and radiation risk in a breast cancer screening programme has received much attention at the initial screening stage. This paper extends that attention to first stage assessment, i.e. the first stage at which women are recalled for further investigation because of suspected lesions or other suspect film features, and prior to any biopsy. Numbers of films, including magnification films, taken at this stage, have been established in two UK regions by different methods. Average total mean glandular dose has been calculated using published data of dose per film from initial screening and a multiplying factor to allow for magnification film dose. It is concluded that the benefit/risk ratio is considerably higher at first stage assessment than at initial screening by a factor of between 4 and 9, because of the very much higher cancer detection rate in this well-defined sub-group of women. Qualitatively, this conclusion is unaffected by the quite wide variation between screening centres in the numbers of films taken at first stage assessment.

Considerable attention has been given to benefit and risk in the initial screening stage of a breast cancer screening programme. This attention has sometimes centred on estimations of the numbers of cancers detected and induced as an indicator of the benefit/risk balance. Numbers of cancers detected and induced are not in themselves a direct measure of benefit and risk. Not all cancers detected by screening will be cured (though earlier detection usually gives a better prognosis), and not all of those induced will be fatal. A fuller discussion of these and related points has been given by us elsewhere [1]. Subsequent assessment stages, after the initial screening visit, have received very much less attention of this kind. This may be because it is easy to see that qualitatively the benefit/risk-balance will be better at assessment than at initial screening, but this point appears not to have been examined in quantitative detail. In the UK Breast Screening Programme, screening rounds occur at 3-yearly intervals. Within each round, after the initial screening visit and examination, a proportion of the screened women are recalled for assessment because of suspicious or doubtful features on their films. This proportion is typically between 5% and 10% of those screened [2]. Women recalled in this way may have further mammograms taken in ‘‘contact’’ or conventional geometry and some will also have magnification films. Films of either of these kinds may be full field, i.e. cover the whole breast, or be collimated to the region of interest. This is sometimes referred to as first stage assessment, and that term will be used in this paper. Most of the women requiring such assessment are eventually given reassurance that all is well, while a minority (say 10–20%) of these are recalled for further The British Journal of Radiology, June 2006

Received 7 April 2005 Revised 20 July 2005 Accepted 31 August 2005 DOI: 10.1259/bjr/33577478 ’ 2006 The British Institute of Radiology

assessment (second stage) which involves biopsy or needle samples [3]. This paper will concentrate principally on the benefit/ risk balance at the first of these two stages, but some aspects of that balance at the second stage will also be discussed. The ratio of cancers detected to estimates of cancers induced at the radiation dose levels involved is not the same as the benefit/risk ratio, but the two ratios have been shown to be related, with the benefit/risk ratio being probably in the broad region of half to two thirds of the detection/induction ratio [4, 5]. It is the detection/ induction ratio which will be mainly considered in this paper, because the numbers required can be estimated with relative ease and reliability.

Numbers of films taken at first stage assessment In order to estimate the numbers of cancers detected and induced at first stage assessment it is necessary to know the average numbers of films of each kind that are taken (i.e. conventional, magnification, coned or full field), the mean dose per film, and the cancer detection rate among women undergoing this assessment stage. Data for the second and third of these three can be obtained from published sources and will be discussed later. This section will describe attempts to find out numbers of films taken. Initially, information was obtained from nine screening centres in the North East, Yorkshire and the Humber regions of England by means of a questionnaire sent to superintendent radiographers. The superintendents 479

J Law and K Faulkner Table 1. Average numbers of films taken per woman attending at first stage assessment – England (North East, Yorkshire and The Humber) (9 screening centres) Plain contact films

0.8–1.9

Magnification films Full field

Coned

1.1–2.1

0.7–1.1

These are ranges of values, derived from a questionnaire sent to nine screening centres.

were asked to provide estimates of the numbers of films of each kind taken per woman attending in their own centre. In doing this, they were asked to use their own judgement regarding current normal practice rather than keeping detailed records for a prospective period, and to provide estimates of both the minimum numbers of films normally taken on all or most women and a maximum number. It was emphasised that this maximum should represent the upper end of the normal range and not the maximum ever known or the worst possible case, since the latter would tend to distort the general picture. The results of this enquiry are given in Table 1. All the centres concerned, except one, now take two views (normally craniocaudal (CC) and lateral oblique (LO)) at the initial visit on each screening round, and this will tend to reduce the numbers of films required at assessment. In the absence of sufficient detailed information, it is assumed for calculation purposes that all films taken in conventional contact geometry are full-field, though this is almost certainly not the case. All the films will have been taken on one breast only, unless a suspicious feature was seen in each breast; in the latter case, the consequent greater dose is balanced by the corresponding increased probability of a cancer being identified. In a separate exercise, data on film numbers were obtained from four Scottish Screening Centres, who extracted numbers from their records retrospectively. These were therefore actual numbers for a defined period chosen essentially at random, e.g. the previous month, and did not depend on memory or judgement. The results obtained in this way are shown in Table 2. This region has not yet adopted two-view screening except for the first screening round. Consequently at first stage assessment films are taken of the CC view on each breast unless these have already been taken at the basic screening visit. Occasionally the mediolateral (ML) view is substituted for the CC view. When two-view screening on all rounds is adopted, all these films will be taken at Table 2. Average numbers of films taken per woman attending at first stage assessment – Scotland (four Screening Centres) Centre

A B C D Mean*

Plain contact films: Full field

Coned

0.34 2.93 1.10 0.73 1.22

1.01 0 0 0.26 0.29

Magnification films (all coned)

Results Mean glandular breast dose at first stage assessment Mean breast dose in the UK Breast Screening Programme has been estimated at a little under 2.5 mGy [6] per woman for the LO view and 2.0 mGy for the CC view, giving a mean dose of 2.25 mGy per view to the assessed breast. A more recent estimate [7] suggests very slightly lower values, but the earlier higher values are retained in this paper. Future estimates of mean breast dose may fluctuate around these values. The ratio of magnification film dose to contact geometry film dose has very recently been estimated to be 2.2¡0.15, giving magnification film doses of 5.5 mGy (LO) and 4.4 mGy (CC) [8]. Since magnification films are equally likely to be taken on either view, the mean magnification film dose is 5.0 mGy. Coned films are assumed, for this calculation, to cover approximately one third of the full breast area, on average [9]. In practice, this proportion varies with sizes of individual breasts and with exact sizes of defining cones on different X-ray units, but one third is considered to be a reasonable average value. In general, for a suspicious feature or features in one breast only, only one breast is imaged at assessment. All Table 3. Average numbers of films taken per woman attending at first stage assessment: results combined from two UK regions given equal weight, based on two-view initial screening Region

0.36 0.30 1.09 0.23 0.70

*Weighted mean, related to size of centre.

480

the basic screening visit, and therefore their numbers are excluded from the following account. All other films are categorised as plain films with full field exposed, plain films coned down, and magnification films, all of which are coned down in this region. Average numbers of films of each kind in each of four centres are given in Table 2. It is clear that practice varies between these four centres in some respects. For example, Centre C does three times as many magnification films as Centres A and B; this is a general policy in that centre, not related to the preference of any one radiologist. In some centres, films taken with a small compression paddle are routinely coned down to about the area of the paddle; in other centres the beam may not be coned in this way, to allow the area of interest to be viewed in relation to the architecture of the breast as a whole. Centre B follows this latter policy so that all their plain films are assumed to be effectively full field for dose estimation purposes. In this region, film numbers are broadly similar to those in Region A, but with many fewer magnification films. Average numbers of films from both these regions are combined in Table 3.

A B Mean*

Plain contact films

Magnification films

Full field

Coned

Full field

Coned

0.8–1.9 1.22 1.6

0 0.29 0.15

1.1–2.1 0 1.0

0.7–1.1 0.70 0.9

*Taking upper end of ranges for Region A.


Radiation benefit and risk Table 4. Mean glandular dose at first stage assessment (i.e. number of films 6 dose per film). For further calculation these numbers are divided by 2 because only one breast is irradiated (i.e. effective mean glandular dose to a single breast 5.1 mGy) Film type

Full contact Coned contact Full magnification Coned magnification Total (single breast)

No. of films

Mean glandular dose (mGy)

1.6 0.15 1.0 0.9

2.25 0.75 5.0 1.7

Per film

Per woman

3.6 0.11 5.0 1.5 10.2

estimates of risk factor for breast cancer induction are based on assumptions of equal radiation exposure to both breasts. Therefore, if the usual breast cancer induction risk factors [1, 8] are taken, the effective breast dose (and hence the effective risk) is half the actual dose to the one breast (Table 4). The mean glandular breast dose is calculated in this way in Table 4, using film numbers from Table 3. Because ranges of film numbers were given in Table 1, the upper ends of those ranges have been used in this calculation.

Cancer detection rate at first stage assessment Typical cancer detection rates are around 6 per 1000 women who attend for basic screening, with an agedependant range mostly within 5–7 per 1000 [2]. The proportion of women recalled for assessment is usually between 5% and 10% [2], and all cancers detected are detected within this sub-group. Thus, for a 10% assessment recall rate, the detection rate at assessment is 10 times what it is at initial screening, and for a 5% recall rate it will be 20 times that figure.

Discussion and conclusions Benefit and risk are difficult to define, but one simple possibility is to regard lives saved from death by breast cancer as benefit and lives lost to radiation induced breast cancer as risk. Cancers detected and induced are not the same as these, but there are reasonable grounds

for thinking that the benefit/risk ratio is around half to two-thirds of the detection/induction ratio [4, 5]. It has already been shown that compared with initial screening, the cancer detection rate has increased 10 times in the sub-group recalled for assessment, if the recall rate is 10%, or 20 times if the recall rate is only 5%. The effective mean glandular dose per woman at first stage assessment has been shown in Table 4 to be up to about 5.1 mGy as against 4.5 mGy at initial screening, an increase of 15%. If we consider only the 5.1 mGy received at assessment, the cancer detection/induction ratio increases by a factor of 10/1.1558.7 compared with its value at initial screening; if the doses at each stage are combined to give 9.6 mGy, this factor becomes 10/ 2.1554.7. In both cases these factors are doubled if the assessment recall rate is 5% rather than 10%. Table 5 shows ratios of cancers detected/induced at various ages for the UK Breast Screening Programme. Column two shows values previously published [1], based on the most recent survey of mean glandular breast doses, and breast cancer induction risk factors from NRPB [1, 10]. These values refer to two-view screening. Column three shows values for first stage assessment, assuming a 10% recall rate; if the recall rate were 5%, these values would be twice as large. Values in column three are greater than those in column two by the factor of 8.7 as outlined in the previous paragraph. Values in column four result from combining the mean glandular dose received at the initial screening visit with that received at first stage assessment, and are obtained by multiplying the values in column two by the factor 4.7 from the previous paragraph. Again, the resulting detection/induction ratios would be doubled if the assessment recall rate were 5% rather than 10%. Uncertainties in the numbers quoted in column two of Table 5 have been discussed in the paper from which these figures have been taken. They will be at least a factor of two in each direction due to the uncertainty in the underlying values for cancer induction per unit dose. Uncertainties in the numbers in columns three and four will be further increased because of the variability of protocol and practice between screening centres at the assessment stage. These considerations do not, however, affect the main conclusion that the benefit/risk balance is much more favourable at first-stage assessment than at initial screening. Table 5. Cancers detected/induced at first stage assessment: two-view screening. Assumes 10% of women are assessed (these numbers would be doubled if the assessment rate were only 5%) Age range (years)

Basic screening*

1st stage assessment (a)

50–54 55–59 60–64 65–70

Figure 1. Flow chart of assessment process. The British Journal of Radiology, June 2006

89 120 182 300

780 1100 1600 2600

(b)

420 560 850 1400

*From Law and Faulkner [1]. (a) Taking effective mean glandular dose received at 1st stage assessment only (5.1 mGy). (b) Taking dose received at initial screening visit (4.5 mGy) plus dose received at 1st stage assessment (5.1 mGy), i.e. 9.6 mGy.

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The numerical risk factors for numbers of cancers induced per unit dose, which underlie all these ratios, may change at some time in the future. Any such change will not affect the relationship between ratios given in column two and those of columns three and four in Table 5, and thus will not affect our main qualitative conclusion. Thus, whatever the relation between these figures and benefit/risk ratios, assuming that this relation is similar at all screening stages, the benefit/risk ratio is greater at first stage assessment than at initial screening, despite the greater dose level, because of the greatly increased cancer detection rate. At later stages of assessment, when biopsy or needle-aspiration and stereoscopic X-ray pairs may occur, the dose will be greater still, but again the cancer detection rate will rise, typically to the region of 30–50% as against 6 per 1000 at initial screening, or 6– 12% at first stage assessment, so that again the benefit/ risk ratio will increase compared with that at earlier screening stages. At the initial screening stage, the female population is assumed to be symptom-free and therefore ‘‘healthy’’. For this reason they are not called patients. Thus, at first stage assessment, the status of an individual may change from being one of a selected group of the general population to a patient, since there are now specific grounds to suspect that disease may be present, which must be confirmed or excluded. For patients, justification of an examination disregards prior exposure, for whatever reason. It is therefore right that at this stage particular regard should be paid to ensuring that the balance of benefit and risk is favourable. At second stage assessment there is a relatively high probability that disease is present. Many other patients will routinely receive X-ray examinations to confirm or exclude the presence of a disease, and often where the probability of disease being present is less than it is at this stage of breast screening. Justification of X-ray exposure at second stage assessment should thus certainly be made as it is for other patients. At first stage assessment, this point is perhaps less clear-cut, but the recalls have been made on suspicion of disease being present at a probability in the region of 5–10%. Many hospital X-ray examinations will be routinely performed at a lower probability of disease or fracture than this, and the women concerned cannot be regarded as ‘‘healthy’’ in the way they are at initial screening. The collective dose to the population of screened women is affected by the recall to assessment rate. Women who attend for first stage assessment receive an additional mean glandular dose of 10.2 mGy per woman (Table 4). If this were to be averaged across the screened population, then it would lead to an increase in the mean glandular dose to a typical woman of 1 mGy to 5.5 mGy assuming a recall rate of 10%. If the recall rate was 5%, then the mean glandular dose to a typical woman would be 5.0 mGy. However, this is not the approach we have

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chosen to adopt, as justification of an examination disregards prior exposure. Regardless of these considerations, once it appears that breast cancer may be present it is essential that all necessary X-ray exposures are made to confirm or reject that possibility. The dose levels reported in this paper for first stage assessment are such that justification of the exposure on the balance of benefit and risk should be straightforward. Thus, there should be no problem in terms of benefit/risk balance in taking whatever films are necessary to this end. As in any X-ray investigation that does not imply that films should be taken without good cause. Similar arguments should also apply at second stage assessment. The worst possible outcome of a screening visit is a false-negative or a false-positive report.

Acknowledgments We wish to thank the Superintendent Radiographers in Breast Screening Centres throughout the North East, Yorkshire and the Humber Region and Scotland for answering our questions and supplying information on screening procedures and numbers of films taken at assessment. This research was partially supported by the European Commission’s Radiation Protection Research Programme (DIMOND III and SENTINEL Projects).

References 1. Law J, Faulkner K. Cancers detected and induced, and associated risk and benefit, in a breast screening programme. Br J Radiol 2001;74:1121–7. 2. National Health Service Breast Screening Programme Reviews, 1994-5 to 2003-4. Statistical Bulletin, London: Department of Health. 3. NHS BSP Annual Review 2004 Changing Lives. Sheffield NHS BSP, 2004. 4. Young KC, Faulkner K, Wall BF, Muirhead C. Review of Radiation risk in Breast Screening. NHS BSP Report 54; Sheffield, NHS BSP, 2003. 5. Law J, Faulkner K. Concerning the relationship between benefit and radiation risk, and cancers detected and induced, in a breast screening programme. Br J Radiol 2002;75:678–84. 6. Young KC. Radiation doses in the UK trial of breast screening in women aged 40-48 years. Br J Radiol 2002;75:362–70. 7. Young KC, Burch A, Oduko JM. Radiation doses received in the UK Breast Screening Programme in 2001 and 2002. Br J Radiol 2005;78:207–18. 8. Law J. Breast dose from magnification films in mammography. Br J Radiol 2005;78:816–20. 9. Faulkner K, Bennison K. An assessment of digital stereotaxis in the NHSBSP. Radiat Prot Dosim 2005;117:327–9. 10. Stokell PJ, Robb JD. SPIDER-1 Software for evaluating the detriment associated with radiation exposure. NRPB-SR261. Chilton: NRPB, 1994.



Radiation risks for the radiologist performing transjugular intrahepatic portosystemic shunt (TIPS) N HIDAJAT,

MD,

P WUST,

MD,

M KREUSCHNER,

MD,

R FELIX,

MD

¨ DER, and R-J SCHRO

MD

Department of Radiology, Charite´ Campus Virchow-Klinikum, University Medicine Berlin, Augustenburger Platz 1, 13353 Berlin, Germany ABSTRACT. The aim of this study is to evaluate the radiation dose to the interventional radiologist in transjugular intrahepatic portosystemic shunt (TIPS) concerning the risk of cancer and deterministic radiation effects and the relation to recommended dose limits. In 18 TIPS interventions radiation doses were measured with thermoluminescence dosemeters (TLD) fixed at the eyebrow, thyroid and hand of the radiologist without special lead shielding of these body parts and at the chest, abdomen and testes under the lead apron. The doses of the eye lens, thyroid gland and hand were assumed to be equal to the corresponding surface doses. The dose at the abdomen under the lead apron was used as an estimation of the ovarian dose. Effective dose equivalent was estimated by Webster’s method. The estimated effective dose equivalent was 0.087 mSv and the effective dose 0.110 mSv. The risk of fatal cancer was of 1026 and the risk of severe genetic defect of 1027 for one single intervention. The maximum permissible number of TIPS interventions was 181, otherwise the dose limit for effective dose would be exceeded. When the radiologist performed more than 372 TIPS procedures per year for many years, the dose to the lens of the eye could exceed the threshold for cataract. If the interventionist performs a large number of TIPS procedures in a year, the risk of fatal cancer and developing cataracts becomes relatively high.

A major source of radiation doses to medical personnel are fluoroscopic procedures [1]. Interventional angiographic treatments are typically associated with higher radiation exposure than diagnostic angiography. The transjugular intrahepatic portosystemic shunt (TIPS) has been shown as effective in the treatment of complications of portal hypertension such as variceal bleeding and refractory ascites [2, 3]. Furthermore, it was shown that TIPS is more effective in preventing variceal re-bleeding than the competitive endoscopic method [2]. Nevertheless, TIPS has been known as an intervention with the highest radiation exposure to patients undergoing abdominal angiographic procedures [4, 5]. In association with the high radiation exposure to the patient undergoing TIPS, it can be assumed that the interventional radiologist is exposed to high levels of scattered radiation. The International Commission on Radiological Protection [6] and the Food and Drug Administration (FDA) [7] has estimated the radiation risk of cancer and deterministic effects, such as cataract of the eye lens, and recommended dose limits. The aim of this study is to evaluate the radiation dose to the interventional radiologist in TIPS concerning the risk of cancer, deterministic radiation effects and the relation to the recommended dose limits.

Methods Radiation doses were measured in 18 TIPS procedures (18 patients, age 54.8¡10.0 years; 15 male and 3 female) with liver cirrhosis and refractory ascites as The British Journal of Radiology, June 2006


complications of the portal hypertension. The intervention was performed with angiographic equipment (Multidiagnost; Philips, Eindhoven, The Netherlands) with under table tube and with pulsed fluoroscopy mode with 12 images per second. Thermoluminescence dosemeters (TLD; Harshaw, Cleveland, OH) with lithium fluoride crystals (3.2 mm63.2 mm60.9 mm) doped with Mg (LiF:Mg) were used for dosimetry. The minimum detectable level was 0.01 mSv. As the total time between regeneration, exposure and next regeneration of the detectors was very short (3 days maximum), no correction for natural radiation (0.05 mSv month21) had to be made. The fading rate of 5% in 12 months was also negligible. Taking into account the individual responsivity of the detectors to the energy of X-ray radiation and the filter that was used (2.6 mm Al), the energy correction factor was found to range from 1.00 for a tube voltage of 60 kV to 1.05 at 100 kV. Reproducibility was found to be 1–2% in doses more than 10 times greater than the detection limit. To measure the radiation exposure to the radiologist (NH), one TLD was attached on both sides of the eyebrows and the thyroid gland over the lead apron, on the back of the hand at the base of the middle finger, on the chest at the breastbone, the abdomen at the navel and at the testes under the lead apron. The radiologist used a 0.35 mm lead equivalent lead apron. A lead glass with 1.0 mm lead equivalent was placed permanently in front of the patient chest between the image intensifier and the radiologist. Thyroid collar and lead glasses were not worn. 483

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To establish a portosystemic shunt, a puncture needle was advanced transjugularly in a catheter through the inferior vena cava into the right hepatic vein. An intrahepatic branch of the portal vein was punctured and self-expandable bare nitinol stents (Angiomed, Karlsruhe, Germany) with a diameter of 10 mm and a length of 4 cm or 6 cm were implanted. The end of the stent was located within the hepatic vein. The stents and the hepatic vein were dilated to 10 mm after deployment. Pressure measurements were performed for the portal vein and right atrium in order to ensure that the portosystemic pressure gradient was reduced to 10– 15 mm Hg. When TIPS placement was completed, the guide wire and angiographic sheath were removed. Doses of the eye lens and thyroid gland were assumed to be equal to the corresponding measured surface doses. The dose at the abdomen under the apron was used as an estimation of the dose to the ovaries. The effective dose equivalent was estimated by Webster’s method [8]. Webster developed an empirical method for estimating effective dose equivalent according to the ICRP 26 [9]. This method is based on organ dose measurements and effective dose calculations of Faulkner and Harrison [10] and uses the 1977 ICRP tissue weighting factors [9] to the under apron dose (Hu) and the over collar dose (Ho) with the equation effective dose equivalent51.56Hu+0.046Ho.

Results Fluoroscopy time was 77.8¡66.3 min, the dose–area product (DAP) 446.0¡279.7 Gy cm2 (mean¡standard deviation). The surface doses are shown in Table 1. One surface dose at the thorax and under the apron was at the minimum detectable level. All other surface doses were above this level. The doses at the eyebrows and the thyroid gland were higher on the right side than on the left side. The left hand was exposed more strongly than the right hand and showed the highest dose among all body parts monitored by TLD. Using Webster’s method [8], the effective dose equivalent [9] was calculated by using the values of the over collar and under apron dosemeters. For radiologists who do not wear a thyroid collar, the effective dose equivalent calculated with Webster’s method is 21% below the effective dose [6] calculated with organ dose Table 1. Radiation (surface) doses of the interventional radiologist as mean¡standard deviation. (In brackets dosemeter position) Organ/body part

Right (mSv)

Left (mSv)

Eye lens (eyebrow) Thyroid gland (over apron at collar) Chest (under apron at breast bone) Abdomen (under apron at waist) Testes (under apron at symphysis) Hand (back of the hand)

0.403¡0.328 0.589¡0.721

0.229¡0.120 0.506¡0.225

484

0.041¡0.025

Table 2. Comparison of the doses of the interventional radiologist with the recommended dose limits and the maximum permissible numer of TIPS interventions per year Organ/body part

Dose (mSv)

Recommended dose Maximum limit [6] (mSv) permissible number/year (n)

Eye lensa Testes Handb Whole body

0.403 0.057 0.918 0.110c

150 50 500 20

a

Right side. Left side. c Effective dose derived from effective dose equivalent (EDE). EDE estimated according to Webster [8]. b

tables and depth dose charts [11]. The effective dose would be 0.110 mSv and the maximum permissible number of TIPS would be 181 per year (Table 2). To meet the recommended dose limit [6], the maximum permissible number of TIPS interventions was estimated to be 181, otherwise the dose limit for the effective dose would be exceeded (Table 2). A pregnant radiologist could perform a maximum of 23 interventions after the declaration of pregnancy during the remainder of the pregnancy (Table 3). The risk of fatal cancer was approximately 1026 and the risk of severe genetic defect 1027 for a single intervention (Table 4). If a radiologist performed more than 248 TIPS per year for many years, the radiologist could reach the threshold for radiation-induced detectable opacity of his eye lens. For a workload of more than 372 times a year and for many years, he could possibly develop a cataract (Table 5).

Discussion In angiography, the patient is exposed by direct radiation and the radiologist by scatter radiation. However, the radiologist may work with and be exposed to radiation every day for many years. In a previous, study with a very limited number of patients undergoing transarterial hepatic chemoembolisation, we showed that the dose to the eye lens is a quantity that may limit the permissible number of interventions before a dose limit would be exceeded [12]. This study shows that it may be deduced from one to two TIPS interventions per day, the dose limit for the eye lens can be exceeded. It also reveals that the effective dose can be a limiting quantity for the number of the interventions, as one TIPS per day can cause the cumulative effective dose to be above the limit Table 3. Comparison of the dose to the abdomen of the interventional radiologist with the recommended dose limit for the conceptus of a pregnant woman and the maximum permissible number of TIPS interventions after the declaration of pregnancy during the remainder of the pregnancy

0.043¡0.032 0.057¡0.041 0.549¡0.313

0.918¡0.516

372 877 544 181

Conceptus a

Dose (mSv) Recommended dose limit (mSv)

Maximum permissible number (n)

0.043a

23

1

Dose to abdomen under apron


Radiation risk for radiologist in TIPS Table 4. Risk of some stochastic damage for the radiologist, calculated according to ICRP 60 [6] Stochastic damage

Risk coefficient

Dose (mSv)*

Risk*

Fatal cancer Genetic Defect

5%/Sv 1%/Sv 1%/Sv

0.110¡0.080a 0.057¡0.041b 0.043¡0.032c

5.5061026¡4.0061026 5.7061027¡4.1061027 4.3061027¡3.2061027

a

Effective dose derived from effective dose equivalent (EDE). EDE determined according to Webster [8]. Dose to testes. c Dose at abdomen used as dose to ovaries. *Dose and risk presented as mean¡standard deviation. b

after 1 year. In the case of pregnancy, one TIPS per day can cause the dose limit recommended by the ICRP for the unborn child be exceeded after 24 days. The risks of fatal cancer derived from effective dose and of severe genetic defect are extremely small for the radiologist after one single TIPS. However, for high workloads, e.g. 500 per year and over 10 years, these stochastic risks can increase to a magnitude of 1023 to 1022. The radiologist should be aware of the radiation exposure to the eye lens and the associated possibility of cataract when he performs interventions like TIPS one to two times per day over many years. Sterility will not occur as a deterministic radiation effect to personnel after TIPS. The dose differences between the right and left side of the eye lens, thyroid gland and hands can be explained by the position of the radiologist in relation to the X-ray tube. While standing at the head of the patient during the TIPS intervention, the tube is mostly located at the liver side, i.e. right side of the patient. Thus, the distance of the right lens and side, of the thyroid gland to the tube is shorter than on the left side. The left hand was used to guide and rotate the catheter and placed at the neck, while the right hand was placed at the proximal end of the catheter to inject contrast media into the catheter. Thus, the left hand was closer to the tube than the right hand. Therefore, in the interests of the radiologist, dose measurements in interventions should be performed at both sides when it is not clear which side is irradiated more strongly. Not only can there be dose differences between the sides, but the dose distribution at the same hand may not be uniform. Whitby and Martin [13] have found that for most interventional radiological procedures the bases of Table 5. Comparison between the measured doses of the radiologist and threshold for radiation-induced deterministic damage [6, 7] Deterministic damage

Eye lensa Detectable opacity Visual impairment (cataract) Testicles Temporary sterility Permanent sterility Ovaries Sterility

Threshold (mSv year21)

.100 .150

400 2000 .200

Mean dose (mSv)

0.403 (n.248) (n.372) 0.057 (n.7017) (n.35087) 0.043b (n.4651)

a

Right side. Dose at abdomen used as dose to ovaries.

b


the ring and little finger receive the highest dose. In this study we measured the dose at the base of the middle finger, so that the maximum dose to the hand could be higher than we have found. Organs or body parts located near the surface, such as the eye lens, may be measured on the surface. For the estimation of the ovarian dose we used the surface dose to the abdomen. In this case, it can be argued that the lead apron shields the individual from radiation, preferentially at lower energies. The parts of the energy spectrum which penetrates the lead apron are the higher energies and these should not be absorbed significantly by the abdomen on their way to the ovaries. Thus, the surface dose under the apron should be a reasonable estimation of the ovarian dose. We did not measure the dose to the leg. A study by Whitby and Martin [14] showed that in TIPS, the leg dose could be as much as two to three times greater than that to the hands. In contrast, Meier et al [5] have found that the dose to the hand was the highest in TIPS (about 0.7 mSv at the hand and 0.2 mSv at the leg). Regarding radiation dose to the radiologist, there are two particular aspects of TIPS. First, it is associated with a long fluoroscopy time [4]. This is because the catheterization of the hepatic vein, puncture and catheterization of the portal vein must be guided by fluoroscopy and is often difficult. Normally, TIPS is performed by an experienced interventionist. Second, many patients needing a TIPS have large amount of ascites, which automatically leads to higher voltage and tube current. In big departments with many different radiological procedures, the competences of the physicians are divided because no one can be really specialized in all of these procedures. Those who are competent for TIPS and perform about 100 of these interventions per year accumulate more radiation dose and may have less reserve for different interventions than others who, for example, perform similar numbers of angioplasties of the lower extremity. To protect personnel from stochastic and deterministic risks, measures of radiation protection should be used when interventions such as TIPS are performed frequently. The distance to the X-ray source should be kept as great as possible as the dose is reduced with the square of the distance. Lead apron, thyroid collar and lead glasses may greatly reduce the operator radiation exposure in cardiac interventions (to 0.8%) [15]. Thyroid collars could reduce the effective dose by approximately a factor of two [11]. Permanent lead shielding, e.g. a lead plate as we have used, can be placed between the patient and the radiologist to reduce the amount of scatter radiation reaching the trunk and head of the radiologist. 485

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Lead gloves have been found to be flexible and could lead to dose reduction of about 20% [16]. For a number of years we have used the pulsed fluoroscopy mode with 12 images per second and cannot see significant reduction of the image quality during the intervention. Zweers et al [17] described significant reduction of the estimated staff effective dose in TIPS using dedicated fluoroscopy exposure settings. Last but not least, a recently described MR guided TIPS with use of a hybrid radiography/MR system should facilitate the puncture of the portal vein and be associated with less radiation exposure than a conventional TIPS procedure [18].

References 1. National Council on Radiation Protection and Measurements. Implementation of the principle of as low as reasonably achievable (ALARA) for medical and dental personnel. NCRP Report no. 107. Bethesda, MD: NRCP, 1990. 2. Ro¨ssle M, Deibert MP, Haag K, et al. Randomised trial of transjugular-intrahepatic-portosystemic shunt versus endoscopy plus propranolol for prevention of variceal rebleeding. Lancet 1997;349:1043–9. 3. Lind CD, Malisch TW, Chong WK, Richards WO, Pinson CW, Meranze SG, et al. Incidence of shunt occlusion or stenosis following transjugular intrahepatic portosystemic shunt placement. Gastroenterology 1994;106:1277–83. 4. McParland BJ. A study of patient radiation doses in interventional radiological procedures. Br J Radiol 1998;71:175–85. 5. Meier N, Lenzen H, Sudhoff A, Fiebich M, Ko¨tter L. X-ray exposures of radiologists and staff in interventional radiology. Radiologe 1995;35:152–5. 6. ICRP. 1990 Recommendations of the International Commission on Radiological Protection. ICRP Publication 60, Ann ICRP, 1991. 7. Food and Drug Administration. Avoidance of serious x-ray induced skin injuries to patients during fluoroscopicallyguided procedures. Food and Drug Administration 1994; September 9: 1–6.

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8. Webster EW. EDE for exposure with protective aprons. Health Phys 1989;56:568–9. 9. ICRP. Recommendations of the International Commission on Radiological Protection. ICRP Publication 26, Ann ICRP, 1977. 10. Faulkner K, Harrison RM. Estimation of effective dose equivalent to staff in diagnostic radiology. Phys Med Biol 1988;33:83–91. 11. Niklason LT, Marx V, Chan HP. Interventional radiologists: occupational radiation dose and risks. Radiology 1993;187: 729–33. 12. Hidajat N, Vogl T, Biamino G, Wust P, Panzer W, Zankl M, et al. Radiation exposure in interventional radiology as exemplified by the chemoembolization of hepatocellular carcinoma and laser angioplasty of the pelvic arteries. Fortschr Ro¨ntgenstr 1996;164:249–56. 13. Whitby M, Martin CJ. A study of the distribution of dose across the hands of interventional radiologists and cardiologists. Br J Radiol 2005;78:219–29. 14. Whitby M, Martin CJ. Radiation doses to the legs of radiologists performing interventional procedures: are they a cause for concern? Br J Radiol 2003;76:321–7. 15. Kuon E, Schmitt M, Dahm JB. Significant reduction of radiation exposure to operator and staff during cardiac interventions by analysis of radiation leakage and improved lead shielding. Am J Cardiol 2002; 89:44–9. 16. Damilakis J, Koukourakis M, Hatjidakis A, Karabekios S, Gourtsoyiannis N. Radiation exposure to the hands of operators during angiographic procedures. Eur J Radiol 1995;21:72–5. 17. Zweers D, Geleijns J, Aarts NJ, Hardam LJ, Lameris JS, Schultz FW, et al. Patient and staff radiation dose in fluoroscopy-guided TIPS procedures and dose reduction, using dedicated fluoroscopy exposure settings. Br J Radiol 1998;71:1333–4. 18. Kee ST, Ganguly A, Daniel BL, Wen Z, Butts K, Shimikawa A, et al. MR-guided transjugular intrahepatic portosystemic shunt creation with use of a hybrid radiography/MR system. J Vasc Interv Radiol 2005;16:227–34.



Organ doses from prostate radiotherapy and associated concomitant exposures 1

R M HARRISON, PhD, 1M WILKINSON, DCR(T), 1A SHEMILT, BSc, 1D J RAWLINGS, MPhil, 1M MOORE and 2 A R LECOMBER, PhD 1

Regional Medical Physics Department, Newcastle General Hospital, Newcastle upon Tyne NE4 6BE and 2Regional Medical Physics Department, Royal Victoria Infirmary, Newcastle upon Tyne NE1 4PP, UK ABSTRACT. In addition to the therapeutic exposure, a course of radiotherapy will involve the additional (concomitant) irradiation of the patient using CT, simulator or portal imaging systems, for localization of the target volume and subsequent verification of treatment delivery. The number of concomitant exposures is likely to increase as the developing technical capabilities for conformal, image-guided radiotherapy make target and critical organ definition an increasingly important aspect of radiotherapy. Estimation of doses and risks to critical organs in the body from all sources is thus necessary to provide the basis for adequate justification of the exposures as required by ICRP. In this paper, doses to selected organs and tissues for which ICRP have identified fatal cancer probabilities have been measured using a realistic anthropomorphic phantom loaded with thermoluminescent dosemeters and irradiated using a treatment protocol for radical radiotherapy of the prostate. Independently, doses to the same organs and tissues have been measured from concomitant CT and portal imaging exposures given for localization and verification purposes. Although negligible in comparison with the target dose, realistic numbers of concomitant exposures give a small but significant contribution to the total dose to most organs and tissues outside the target volume. Generally, this is in the range 5–10% of the total organ dose, but can be as high as 20% for bone surfaces. These data may be used to estimate concomitant doses from any combination of CT and portal imaging and may help in the justification process, especially when additional verification exposures may be required during treatment.

The entire process of radiotherapy involves exposure of the patient to various sources of ionizing radiation. In addition to the therapeutic exposure itself, localization and verification (concomitant) exposures may be carried out before and during the treatment using CT scanning, conventional simulation and megavoltage portal imaging. Estimation of doses to critical organs in the body from all sources is necessary in order to assess total stochastic (second cancer) risks and thus provide adequate justification of the exposures as required by ICRP [1] and the corresponding UK legislation (The Ionising Radiation (Medical Exposures) Regulations, 2000 (IR(ME)R). [2]). This is increasingly important because of the concomitant exposures associated with the development of complex conformal and imageguided radiotherapy. Numerous workers have reported estimates of second cancer incidence following radiotherapy [3]. This is likely to remain an important issue as, following improvements in cancer treatment, more patients may survive longer than the latent period for expression of the second cancer. In addition, several new developments in radiotherapy may lead to higher doses to organs and tissues throughout the body. Intensity-modulated radiotherapy (IMRT) may lead to substantial increases in whole body The British Journal of Radiology, June 2006

Received 8 July 2005 Revised 25 August 2005 Accepted 23 September 2005 DOI: 10.1259/bjr/16187818 ’ 2006 The British Institute of Radiology

doses, because of the longer beam-on times and the consequent increased exposure to leakage radiation. The use of an increased number of fields results in a larger volume of tissue exposed to relatively low doses, although the target volume may be smaller. Compared with conventional radiotherapy, Hall and Wuu have estimated that IMRT may almost double the incidence of second cancers [4]. New imaging techniques are also contributing to verification of treatment field positions during the course of radiotherapy. Although megavoltage portal imaging is a long-established technique, image-guided radiotherapy (IGRT), involving CT studies at many fractions [5], cone beam CT using megavoltage and kilovoltage X-ray energies [6, 7] and integrated systems such as tomotherapy [8], are also emerging as viable techniques for improving the accuracy of dose delivery to the target volume and may increase organ doses and subsequent risks. In contrast, the potential for reduction of second cancers by using proton radiotherapy has been investigated by Miralbell et al [9]. Previous estimates of second cancer risk have necessarily made simplifying assumptions about doses and risks. Followill et al [10] estimated a single whole body dose at a point distant from the treatment field and applied a risk factor for the general population. Mutic 487

R M Harrison, M Wilkinson, A Shemilt et al

and Low [11] measured doses outside the target volume using thermoluminescent dosimetry (TLD) and a block phantom. Verellen et al [12] used personal dosemeters to measure the personal dose equivalent Hp(10) for six patient treatments, related this to the effective dose and used a nominal risk factor for the general population. More recently, Kry et al [13] have used NCRP risk factors [14] to compare several IMRT treatments of the prostate. The purpose of these papers was to compare different irradiation techniques (e.g. conventional radiotherapy versus IMRT), thus placing the emphasis on relative, rather than absolute estimates of risk. In other fields, diverse exposures and irradiation patterns are usually combined using the quantity effective dose. However, there are several problems associated with using this concept in the radiotherapy context [15]. ICRP risk and tissue weighting factors refer to the general population, whereas the age distribution of radiotherapy patients will be skewed towards the higher ages, implying lower risk estimates. Conversely, tissue weighting factors refer to low doses associated with occupational exposure and a dose and dose rate effectiveness factor (DDREF) of 2 has been applied. It is not clear what modifications to the DDREF should be made for the effects of radiotherapy fractionation schedules and cell kill on induced cancer probability at high doses. Thus, the estimation of second cancer induction probability from a combination of weighted organ doses is not trivial and no attempt has been made in this paper to combine individual organ doses, which include a radiotherapy component, to form an overall measure of risk. Nevertheless, as a first step, organ doses measured in several critical organs in a realistic anthropomorphic phantom may be useful for justification purposes and will help to place concomitant doses in perspective, particularly for developing and justifying techniques where more extensive imaging is used to improve the accuracy of dose delivery. This paper describes the experimental simulation of radical prostate radiotherapy together with associated CT scanning and portal imaging, with doses measured using lithium fluoride TLD chips loaded in a male anthropomorphic phantom. In anticipation of eventually calculating second cancer induction probabilities, doses were measured in organs (apart from the skin) to which ICRP has ascribed tissue weighting factors [1]. Skin was omitted because of the uncertainties arising from the dose gradient at the surface due to build up effects at megavoltage energies, the identification of the effective point of measurement of a TLD chip under these conditions and the large number of surface measurements which would in any case be necessary in order to form an unambiguous estimate of total skin dose. Furthermore, the skin has the lowest fatal cancer probability coefficient of all the main organs identified by ICRP and is not as critical as other tissues when considering second cancers following radiotherapy. External beam radiotherapy of the prostate was chosen for study because it is the most common cancer in UK men and advances in treatment have shown an increase in 10 year survival from 20% in 1971 to 50% in 2001 [16]. Brenner et al [17], in a comparison of second cancers in patients treated for prostate carcinoma by radiotherapy 488

compared with surgery, have shown that there is a small but significant increase in solid tumours from 5 years following radiotherapy, principally showing as carcinomas of the bladder, rectum and lung, and sarcomas in or near the treatment field. A further study has also shown a significant increase in second rectal cancers [18]. Measurement of organ doses from CT scanning and portal imaging allows doses to be calculated for any number and combination of these imaging techniques, thus facilitating the estimation of the organ dose contribution from image-intensive IGRT.

Materials and methods Radiotherapy The position of the prostate and surrounding critical organs such as the bladder and rectum were established within a male RANDO phantom (The Phantom Laboratory Incorporated, Salem, NY) and a treatment plan for isocentric external beam prostate radiotherapy developed according to protocols in place in this centre, using a Helax-TMS treatment planning system (Nucletron B.V., Veenendaal, The Netherlands) and CT localization scans of the phantom obtained with a Siemens Emotion Duo CT scanner (Siemens Medical Solutions, Erlangen, Germany). The simulated prostate treatment consists of two phases. In phase 1, 64 Gy is delivered in 32 fractions to the planning target volume (PTV), followed in phase 2 by 10 Gy in 5 fractions to a reduced target volume. The prostate, bladder and rectum were identified on slabs 32, 33 and 34 of the phantom, following consideration of a series of patient CT scans, and a three-field plan for both phases developed as shown in Figure 1. Field sizes and gantry angles for both treatment phases are given in Table 1. Anteroposterior (AP) and lateral digitally reconstructed radiographs were also generated (Figure 2). The phantom was loaded with Harshaw TLD-100 (LiF:Mg:Ti) chips (Thermo Electron Corporation, Solon, OH), positioned within each organ of interest at the centre of each phantom slab and secured with tissue equivalent plugs. Each chip had a unique identifier and calibration factor. Anatomy texts were used to check the position of organs and tissues, and sampling points were selected to give several measurements for each critical organ. Table 2 shows the numbers of TLD chips contributing to the dose measurement for each organ or tissue. Treatment was delivered using a Siemens PRIMUS H1 linear accelerator (Siemens Medical Solutions, Erlangen, Germany) operating at 15 MV. For the simulation of this treatment, the phantom was irradiated to 0.1 of the total target dose (i.e. 7.4 Gy) and the subsequent measured photon doses scaled to give the doses which would have been received for a full clinical treatment.

Neutron dose contribution At 15 MV, photoneutron production in the linear accelerator head will contribute to the organ and tissue The British Journal of Radiology, June 2006

Organ doses from prostate radiotherapy

(a)

(b)

Figure 1. Plan for prostate treatment simulation on slab 34 of a RANDO phantom (a) phase 1 treatment (64 Gy) and (b) phase 2 treatment to a smaller target volume (10 Gy).

doses. Measurement of individual organ doses due to neutrons was not attempted, but instead calculated from previously reported measurements of neutron fluence. Expressions for direct, thermal and scattered neutrons from medical linear accelerators, derived from Monte Carlo calculations by McCall [19, 20] were used to calculate the neutron fluence at the centre point of each slice of the RANDO phantom. They are reproduced in Equation (1) where the three terms give direct, scattered and thermal neutron components, respectively. The total neutron fluence per photon dose at the isocentre (W), is given by: W~ aQ=4pd 2 zð5:4aQ=S Þzð1:26Q=S Þ

ð1Þ

where a is the transmission factor for neutrons in head shielding (taken as 0.85); Q is the neutron source strength per unit photon dose at the isocentre for a Siemens PRIMUS accelerator (taken as 0.2161012 neutrons Gy21) [21]; d is the distance from the source; and S is the treatment room surface area (180 m2). Conversion of W to equivalent dose per Gy at the isocentre used the fluence-to-dose conversion factors given by NCRP [14], assuming a mean neutron energy of 0.5 MeV for an accelerator running at 15 MV [22, 23] and taking values of radiation weighting factor wR of 20 and 5 for fast and thermal neutrons, respectively. Although the ionization suffered by lithium fluoride TLD will include contributions from (n,p) and (n,c) reactions, this has been ignored in practice because the wR values ascribed to the

neutron absorbed dose are significantly higher than unity.

CT scanning for localization Using the same TLD loading scheme as for the radiotherapy treatment, the phantom was scanned using the standard departmental protocol for prostate localization, using a Siemens Emotion Duo scanner (Siemens Medical Solutions, Erlangen, Germany). The irradiation consisted of an AP topogram (130 kV; 30 mAs; topogram length 512 mm) followed by an axial scan (130 kV; 105 mAs; slice width 5 mm; pitch 1.5; 33 slices). Ten identical consecutive scans (topogram and axial) of the phantom were carried out in order to ensure that the doses at the TLD measurement positions were well above the minimum detectable value. Measured doses were subsequently scaled to give the doses received for a single localization scan. (Doses from conventional simulation were not measured, since a simulator is not routinely used in this centre for prostate localization).

Portal imaging Using the same TLD loading scheme as for the radiotherapy treatment, the phantom was irradiated by a 15 cm615 cm AP portal image exposure, as used in conjunction with the Siemens portalvision electronic portal imaging system (Siemens Medical Solutions,

Table 1. Treatment parameters for the simulation of prostate radiotherapy

Phase 1 Phase 2

Total target dose (Gy)

AP field size (cm6cm)

Left lateral field size (cm6cm)

Right lateral field size (cm6cm)

64 10

6.866.7 4.964.7

6.766.6 5.564.6

6.766.6 5.564.7


489


TLD calibration and dose measurement Following irradiation, chips were left for a fixed time of 48 h to allow the same fading time for all irradiations and read out in a Harshaw 5500 TLD reader (Thermo Electron Corporation, Solon, OH). Individual chip calibration factors were established based on irradiation of each chip to a known dose using a Cs-137 calibration source. These factors provide a correction for interchip variability within a batch. In addition, batches of chips were irradiated within the range of doses and energies used in the simulations in order to generate a calibration curve. Doses were measured with ionization chambers and electrometers whose calibrations were traceable to national standards.

Figure 2. Digitally reconstructed AP radiograph of the RANDO phantom with target and critical volumes outlined. The prostate is shown in yellow with a red margin for the clinical target volume (CTV). The bladder is shown in orange and the rectum in dark red. The planning target volume (PTV) is demarcated by yellow lines.

Erlangen, Germany) mounted on the same linear accelerator as used for treatment, but running at 6 MV. Typically, a clinical portal image would be acquired using 3 monitor units (MU). For simulation, 200 MU were used in order to ensure a measurable dose to distant organs and the subsequent doses scaled to give the commensurate clinical exposure for a single portal image. A 10 cm610 cm lateral portal image exposure was also simulated in the same way for comparison.

Table 2. Thermoluminescent dosemeter (TLD) loadings for organs and tissues Organ or tissue

Number of TLD chips

RANDO slabs used for each organ or tissue

Gonads (testes) Bone marrow and bone surface Colon (and upper large intestine) Lung Stomach Bladder Liver Oesophagus Thyroid Adrenals Brain Upper large intestine Small intestine Kidney Pancreas Spleen Thymus Rectum Total

4 101

35 5–6, 8–34

490

60

24–31

59 23 8 36 11 12 2 7 29 20 11 7 6 2 3 401

10–20 20–24 33–34 19–25 9–19 9–11 21–22 1, 3, 5 24–27 24–30 21–25 22–23 20–22 14 32–34

Dose calculation Raw data from each chip readout were modified by the individual chip calibration factor and the dose calibration factor for the appropriate photon energy, giving the result as absorbed dose to water at the chip position. Organ doses were then calculated as weighted averages over all the TLD measurement points, using the mass fraction of each organ within each slab [24]. For lung, bone surface and bone marrow, the mass fractions given by Huda [25] were used. Doses to bone marrow and bone surface were calculated from the average of measurements to the positions of these organs within each slab and also, for comparison, from the average dose over the whole slab, in both cases using the same mass fractions. The total doses to these organs were obtained by summing the dose contributions for each slab. All organs, except for bone surfaces, were assumed to be water-equivalent. Doses to bone were derived from doses to water by application of mass stopping power ratios (bone/water) of 0.91 and 0.93 for 6 MV and 15 MV, respectively, and for CT energies a mass energy absorption coefficient ratio (bone/water) of 3.1 was used. Chips which had received doses .2000 mGy were not re-used. Calculated neutron doses for each organ were added to the photon contributions for the simulation of treatment at 15 MV. The current protocol in use in this centre defines the maximum number of concomitant images which can be acquired under a single authorization by the Practitioner, as defined under UK legislation [2]. It allows for the following: (i) three localization episodes on the CT scanner; (ii) three pre-treatment verification episodes, at least one on the CT scanner and one on the treatment unit; and (iii) eight verification episodes during the course of treatment. A localization or verification episode on a CT scanner consists of a maximum of three topograms, two CT scans up to 120 cm in length and three single axial slices. A pre-treatment verification episode on the treatment unit consists of a maximum of four portal images. A verification episode during treatment delivery consists of a pre-treatment verification episode and a localization or verification episode on the CT scanner. Several combinations of CT scans and portal images are possible within this framework and organ doses for two possibilities were calculated: (i) 10 The British Journal of Radiology, June 2006


CT scans and 36 portal images, where portal imaging is the primary method of verification; and (ii) 26 CT scans and 4 portal images, where CT scanning is the primary method of verification.

Results Chip calibration The uncertainty in a single TLD measurement was found to be ¡4.3%, averaged over calibrations at all photon energies. The minimum detectable dose was 1 mSv, taken to be the dose corresponding to three times the mean background signal.

Neutron results Figure 3 shows the neutron dose (in mSv Gy21 at the isocentre) as a function of distance from the isocentre normal to the central axis, i.e. along the long axis of the phantom. This enabled a representative neutron dose to be calculated for each slab. The decrease of dose with distance is simply a consequence of an inverse square term acting on the direct neutron component and involving distance from the X-ray target.

Organ and tissue doses Table 3 gives organ and tissue doses for the complete radiotherapy treatment, single CT localization scan and topogram, and single AP portal image. Organ doses resulting from a lateral 10 cm610 cm portal exposure were less than those arising from the 15 cm615 cm AP exposure and the latter have been adopted as representative of prostate portal images. Doses are given for all organs for which tissue weighting factors have been identified [1], apart from the skin, and include remainder organs. Apart from muscle, bone marrow and bone surface, ranges of doses are also given for organs

containing more than 5 TLD measurement points, or where mean organ doses are .0.1 mSv. Although the dose to the rectum is included within the estimate for the colon, the rectal dose has been calculated separately, because of its proximity to the prostate, and is included in the footnote to Table 3. In fact, part of the rectum and bladder were included in the target volume and received measured doses in excess of 60 Gy and 73 Gy, respectively. The neutron component of organ doses ranges from 19–50 mSv for the whole treatment. Since different approaches are taken when considering high doses compared with the low doses usually associated with radiation protection, some assumptions need to be made when combining or comparing them. Since most organ absorbed doses are ,500 mGy, equivalent dose is used throughout to allow photon and neutron components to be summed and effective dose to be calculated for CT and portal exposures. For consistency, it is assumed that wR51 for the high photon doses to organs close to or within the target volume, although it is accepted that the dose response function for carcinogenesis may differ at low and high doses. Table 4 gives the total organ and tissue doses for a complete course of radiotherapy involving a total of 10 CT scans and 36 portal images. The data are also shown graphically as total doses (Figure 4a) and as percentages of the total doses from radiotherapy, CT scanning and portal imaging for each organ (Figure 4b). Table 5 gives organ doses for 26 CT scans and 4 portal images. Figure 5 shows total doses and doses as percentages of the total organ dose, respectively. A subset of these data have previously been reported in preliminary form [26]. Approximate doses to organs in other positions may be estimated from Figure 6, which shows a plot of mean dose for each slab. Within and around the target volume in slab 34, doses to the chip positions were estimated from the treatment planning system dose calculation algorithm and summed for both phases of treatment. It was assumed that the planned dose distribution was delivered exactly. Figure 7 shows the difference between the delivered and measured doses at 11 points as a fraction of the delivered dose. For reasons given above, effective dose is an inappropriate quantity for use in radiotherapy treatment. However, the concept may be applied, with caution, to diagnostic X-ray examinations such as CT scanning and, in view of the relatively low doses involved, to portal imaging. Applying the tissue weighting factors and remainder organ rules stipulated by ICRP [1], the measured organ doses for a CT localization scan (including topogram), single AP portal image (3 MV 15 cm615 cm at 6 MV) and single lateral portal image (3 MV 10 cm610 cm at 6 MV) gave effective doses of 4.2 mSv, 2.3 mSv and 0.8 mSv, respectively.

Discussion Figure 3. Neutron dose (direct, scattered and thermal) as a function of distance from the isocentre normal to the central axis of the beam. These values have been calculated from equation 1 using a50.85, Q50.2161012 neutrons Gy21 and S 5 (1806104) cm2. The British Journal of Radiology, June 2006

Uncertainties in dose estimation Organ doses have been measured under ideal conditions. There was no movement of the phantom during treatment, unlike the interfraction and intrafraction 491

R M Harrison, M Wilkinson, A Shemilt et al Table 3. Organ and tissue doses for complete radiotherapy treatment, single CT localization scan and single anteroposterior (AP) portal image Organ or tissue

Bladder Colon1 Lung Stomach Liver Oesophagus Thyroid Bone marrow3 Bone surface3 Adrenals Brain Small intestine1 Kidney Muscle Pancreas Spleen Thymus

Equivalent dose (mSv)2 Radiotherapy treatment (74 Gy at 15 MV)

Radiotherapy treatment (74 Gy at 15 MV)

Photon contribution

Neutron contribution

28750 2600 62 95 78 59 45 1350 719 89 29 242 90 100 103 73 31

48 45 40 43 42 40 38 42 27 42 34 45 43 18 43 42 20

Total radiotherapy (dose range in parenthesis)

Single CT localization scan and AP topogram (dose range in parenthesis)

Single 15 cm615 cm AP portal image (6 MV)

28800 (4000–73200) 26501 102 (30–120) 138 (50–170) 120 (40–190) 99 (30–90) 83 (30–90) 13903 7463 132 63 (28–32) 2871 133 (60–160) 118 146 (60–160) 115 (50–140) 51 (30–90)

11.6 (10–12) 3.11 ,0.1 0.2 (0.08–0.5) 0.1 (0.04–0.7) ,0.1 ,0.1 3.33 5.93 0.2 ,0.1 2.21 0.3 (0.2–0.7) 1.5 0.2 (0.2–0.3) 0.1 ,0.1

27.8 (24–61) 2.81 ,0.1 0.11 (0.07–0.19) 0.08 (0.05–0.18) ,0.1 ,0.1 4.503 2.363 0.10 0.20 (0.19–0.22) 0.971 0.14 (0.11–0.28) 0.32 0.12 (0.11–0.16) ,0.1 ,0.1

1

Dose to the colon is a weighted average over the lower and upper large intestine as recommended by ICRP [1]. However, the rectum receives a large dose from prostate radiotherapy because of its proximity to the target volume. Thus the rectal dose has been estimated separately as 24 000 mSv. The ranges of dose for LLI, ULI and SI were (120–1040 mSv), (70–480 mSv) and (110–480 mSv), respectively, for radiotherapy, (1–15 mSv), (0.3–1.8 mSv) and (0.4–13 mSv), respectively, for CT and (0.25– 33 mSv), (0.15–0.43 mSv) and (0.17–19 mSv), respectively, for a single AP portal image. 2 For radiotherapy measurements, doses are quoted to 3 significant figures or 50 mSv for high doses. For the CT results, doses are quoted to 0.1 mSv, since the measured doses were for 10 CT scans. For the portal image doses, results are quoted to 0.01 mSv, since the delivered doses were 67 times the doses for a typical portal image. 3 Dose ranges for these tissues are not given, since they vary over the whole range of doses delivered, depending on the location within the body and the proximity to the target volume.

Table 4. Organ and tissue doses for complete radiotherapy treatment, 10 CT localization scans and 36 portal images Organ or tissue

Total radiotherapy dose1 (mSv)

Bladder 28800 Colon 2650 Lung 102 Stomach 138 Liver 120 Oesophagus 99 Thyroid 83 Bone marrow 1390 Bone surface 746 Adrenals 132 Brain 63 Small intestine 287 Kidney 133 Muscle 118 Pancreas 146 Spleen 115 Thymus 51 1

Dose from 10 CT scans and topograms1 (mSv)

% of total dose due to CT

Dose from 36 AP portal images1 (mSv)

% of total dose due to portal images

Total dose1 (mSv)

116 31.2 0.3 1.7 1.4 0.2 ,0.1 33 59 1.7 ,0.1 21.6 3.1 14.6 2.3 1.4 0.2

0.4 1.1 0.3 1.2 1.1 0.2 0.05 2.1 6.6 1.2 0.0 6.3 2.2 10.1 1.5 1.2 0.4

999 102 1.1 3.8 3.0 3.1 3.3 91 84.8 3.8 7.3 35 5.0 11.6 4.4 3.1 0.5

3.3 3.7 1.1 2.6 2.4 3.0 3.8 6.0 9.5 2.7 10.4 10.2 3.5 8.1 2.9 2.6 0.97

29900 2780 103 144 124 102 86 1520 890 138 70.3 344 141 144 153 120 51.7

For radiotherapy measurements, doses are quoted to 3 significant figures or 50 mSv. For the CT results, doses are quoted to 0.1 mSv or three significant figures, since the measured doses were for 10 CT scans. For the portal image doses, results are quoted to 0.1 mSv or three significant figures, since the delivered doses were 67 times the doses for a typical portal image.

492



Figure 4. Organ doses from radiotherapy (photons and neutrons), 10 CT scans and 36 portal images for each organ or tissue (a) total doses for each organ and (b) % contribution of radiotherapy, CT scanning and portal imaging. Mean bladder, colon and rectal doses were 29.9 Gy, 2.8 Gy and 25 Gy, respectively.

movement uncertainties and changes in shape which might be seen in clinical treatments. In spite of making 401 individual TLD measurements, each critical organ has been sampled fairly coarsely. Nevertheless, the measurement accuracy (4.3% for a single TLD measurement) is adequate, bearing in mind the subsequent use of such data in risk estimation where considerable uncertainties exist. In addition to the uncertainty per TLD measurement, there are several other sources of uncertainty in organ dose estimation. There will be genuine dose variations within most organs because of organ extent, proximity to the target volume and other dose gradients associated with the edges of CT and portal image fields. These are reflected in the ranges given in parenthesis in Table 3. In addition, there are uncertainties associated with the location of organs and the choice of sampling points within them. These are difficult to quantify. Finally, measurements on a single size and shape of anthropomorphic phantom will not represent the interpatient variation seen in practice. The British Journal of Radiology, June 2006

Estimating doses to bone surfaces and red bone marrow is particularly difficult because of the extended and complex distribution of these tissues throughout the body. Although dosemeters were placed close to selected bone locations, the uncertainties in dose estimation for marrow and surfaces are likely to be greater than for more well-defined organs. As pointed out by Golikov and Nikitin [24], the mass fractions for bone given by Huda and Sandison [25] and used here, refer to the mass distribution of the whole skeleton, which may differ from the corresponding distribution of bone surface cells. Doses to bone surfaces and marrow are also some of the higher organ doses measured. This emphasises the potential difficulty of using this technique as a reliable basis for absolute risk assessment and suggests that the technique may be more appropriate for the calculation of relative risks, for example between two different treatment modalities or between treatment and imaging exposures, where dose measurement locations and assumptions about organ position and extent are invariant. For regions in or near the target volume, TLD measurement is likely to be considerably less accurate than calculation using the well-developed algorithms used in current treatment planning systems in combination with accurate ionization chamber measurements of dose distributions in water. Hence, for organs and tissues close to the target volume, use of the treatment planning system is probably the best method for organ dose calculation. However, planning systems are not designed for calculation of doses to distant organs and do not, in any case, include neutron dose estimates. Calculated organ neutron doses are subject to several assumptions inherent in the values given for the terms in Equation (1). Furthermore, it has been assumed that the estimated dose to a slab is appropriate for all organs within the slab, whereas, in practice, this assumption will overestimate the neutron dose to centrally placed organs. Whilst negligible in comparison with doses to the target and its immediate surroundings, the neutron contribution forms a substantial fraction of the total dose to remote organs and tissues, although this would only apply for treatments at energies in excess of approximately 10 MV.

Dose and risk estimates The bladder and rectum, as expected, receive the highest dose, since part of the bladder wall and rectum will be contained within the target volume for this treatment. Considerable variation in bladder and rectal volumes, both within and between patients, have been reported [27] and this will lead to corresponding variations and uncertainties in bladder and rectal doses – and risks – in practice. The observation of second bladder and rectal cancers in patients who have previously received radiotherapy for prostate cancer is consistent with these high doses [17], since it implies that parts of each organ may receive doses near the peak of the dose–response curve. In Figures 4 and 5, the numbers of concomitant exposures have deliberately been set to correspond to the upper limits of what is expected in this centre and it 493

R M Harrison, M Wilkinson, A Shemilt et al Table 5. Organ and tissue doses for complete radiotherapy treatment, 26 CT localization scans and 4 portal images Organ or tissue

Total radiotherapy Dose from 26 CT scans and dose1 (mSv) topograms1 (mSv)

% of total dose due to CT

Dose from 4 AP portal images1 (mSv)

% of total dose due to portal images

Total dose1 (mSv)

Bladder Colon Lung Stomach Liver Oesophagus Thyroid Bone marrow Bone surface Adrenals Brain Small intestine Kidney Muscle Pancreas Spleen Thymus

28800 2650 102 138 120 99 83 1392 746 132 63 287 133 118 146 115 51

1.0 3.0 0.87 3.0 2.9 0.6 0.12 5.7 17 3.2 0.01 16.1 5.6 24 3.9 3.0 0.78

111 11.4 0.1 0.4 0.3 0.3 0.4 18.0 9.4 0.4 0.8 3.9 0.6 1.3 0.5 0.3 0.1

0.36 0.41 0.11 0.29 0.27 0.34 0.42 1.2 1.0 0.3 1.2 1.1 0.38 0.58 0.32 0.29 0.11

29200 2740 103 143 124 100 83 1500 908 137 64 347 142 158 153 119 51.5

1

302 81.1 0.9 4.4 3.6 0.6 0.1 85.6 153 4.4 ,0.1 56.1 8.0 38.2 6.0 3.6 0.4

For radiotherapy measurements, doses are quoted to 3 significant figures or 50 mSv. For the CT results, doses are quoted to 0.1 mSv or three significant figures, since the measured doses were for 10 CT scans. For the portal image doses, results are quoted to 0.1 mSv or three significant figures, since the delivered doses were 67 times the doses for a typical portal image.

is seen that only for bone surfaces do the concomitant exposures exceed 10% of the total organ dose for both cases illustrated, with red bone marrow contributing .10% for the case of treatments in which portal imaging is the verification technique used. To place concomitant doses in perspective, we may consider the excess relative risk (ERR) associated with a population receiving radiotherapy and associated concomitant exposures, compared with a hypothetical population receiving radiotherapy only, assuming that the probability of cancer induction per unit dose is constant. The ERR is given approximately by: ERRT &

ðDT,radio zDT,concom Þ{DT,radio DT,concom ~ DT,radio DT,radio

ð2Þ

where: DT,radio and DT,concom are the doses to organ T from radiotherapy only and concomitant only exposures, respectively. Taking the two examples of concomitant exposures (10 CT scans + 36 portal images and 26 CT scans + 4 portal images), in both cases, ERRs for most organs are ,0.1, with bone surfaces, small intestine and muscle ,0.3. A single CT scan or portal image will give an ERR ,0.01 for all organs. In comparison with work by Waddington and McKenzie [28] the effective dose for a male AP pelvic portal image (15 cm615 cm at 6 MV) is higher (0.78 mSv MU21) compared with their 0.34 mSv MU21). However, no account has been taken here of the fraction of bladder wall volume which lies within the target volume and for which the risk of radiation carcinogenesis may be reduced because of cell kill. This would have the effect of reducing the effective tissue weighting factor for the bladder and hence the effective dose. Moreover, our sampling of the relevant bladder tissue is coarse compared with dose estimations using the dose plotting facilities of a treatment planning system. Waddington 494

and McKenzie assumed that 25% of the bladder was outside the treated volume and a similar assumption would reduce our figure of effective dose to 0.42 mSv MU21. They also assumed that any cancer cells induced in the bladder tissue within the target would have a negligible chance of survival. Whilst simple exponential cell kill would support this assumption, there is conflicting evidence from epidemiological studies, summarized by Hall and Wuu [4]. We find an effective dose of 0.27 mSv MU21 for the lateral portal field, compared with 0.32 mSv MU21 [28] which possibly reflects the smaller field size. In any case, estimates of effective dose will vary with patient size and target volume, so that these results should be considered to be representative rather than definitive. It has previously been argued [15] that for legislative purposes under the UK IR(ME) Regulations, a single justification for a pre-determined combination of radiotherapy and concomitant exposures is practically more appropriate than the justification of individual concomitant exposures, provided that the contributions of the latter are appreciated. These results support this suggestion, since even for a relatively large number of concomitant exposures, the contribution to the overall organ dose, whilst not negligible, is nevertheless small. However, it should be noted that for radiotherapy treatments carrying a good prognosis, especially for younger patients and children, the added risks of concomitant exposures may need to be considered carefully in the justification process. This is especially true for treatments which involve imaging at each fraction, either by CT or portal imaging, such as the developing techniques associated with image-guided radiotherapy. The combination of radiotherapy and concomitant doses to give an unambiguous estimate of risk of second malignancy is not trivial and, because of the inappropriateness of the current definition of effective dose, no attempt has been made in this paper to do so. The British Journal of Radiology, June 2006


Nevertheless, the doses given here allow the total organ and tissue doses from various combinations of radiotherapy, localization and verification exposures to be estimated for this particular treatment.

Conclusions Doses to organs and tissues for which ICRP have identified fatal cancer probabilities have been measured using TLD for a radical radiotherapy treatment of the prostate delivering a total of 74 Gy to the PTV. Independently, doses to the same organs and tissues have been measured from concomitant CT and portal imaging for localization and verification. Combinations of numbers of CT scans and portal images with the radiotherapy treatment itself has allowed total organ doses to be calculated. Realistic numbers of concomitant exposures give a small but significant contribution to the total dose to most organs and tissues. Generally this is in the range 1– 10%, but can be as high a 20% for bone marrow and bone surfaces. At 15 MV, a significant proportion of the doses to organs distant from the target volume is due to photoneutrons produced in the treatment head. These data may be useful for giving a realistic perspective on the contribution of concomitant doses to the overall organ dose burden for this treatment, for protection of the patient and compliance with the requirements of UK Regulations [2]. Whilst doses from similar abdominal radiotherapy treatments may be inferred from these data, further work will be necessary to establish organ and tissue doses for radiotherapy of other parts of the body, both for adults and children. Figure 5. Organ doses from radiotherapy (photons and neutrons), 26 CT scans and 4 portal images for each organ or tissue (a) total doses for each organ and (b) % contribution of radiotherapy, CT scanning and portal imaging. Mean bladder, colon and rectal doses were 29.2 Gy, 2.7 Gy and 24.3 Gy, respectively.

Acknowledgments We would like to thank members of staff of the Regional Medical Physics Department and the Northern Centre for Cancer Treatment who contributed to aspects of this work.

Figure 6. Mean dose in each slice vs distance from isocentre normal to the central axis, as a percentage of the target dose (74 Gy). Each slab of the RANDO phantom is 2.5 cm thick. Error bars represent¡1 SD of the doses measured in each slab. The British Journal of Radiology, June 2006

495


Figure 7. Delivered - measured dose as a fraction of the delivered dose for 11 TLD measurement points within slab 34 of the RANDO phantom. Error bars represent the uncertainty in a single TLD measurement (¡ 4.3%).

References 1. ICRP. 1990 Recommendations of the International Commission on Radiological Protection. Pergammon Press, 1991. Annals of the ICRP; 21 (1–3). 2. IR(ME)R. Ionising Radiation (Medical Exposure) Regulations. HMSO, London, 2000. SI 2000 No. 1059. 3. NRPB. Risks of second cancers in therapeutically irradiated populations. National Radiological Protection Board, Chilton, Didcot, Oxon OX11 0RQ, UK, 2000. Documents of the NRPB; vol. 11 no. 1. 4. Hall JD, Wuu C-S. Radiation-induced second cancers: the impact of 3D-CRT and IMRT. Int J Radiat Oncol Biol Phys 2003;56:83–8. 5. Fung AYC, Grimm S-YL, Wong JR, Uematsu M. Computed tomography localization of radiation treatment delivery versus conventional localization with bony landmarks. J App Clin Med Phys 2003;4:112–9. 6. Lewis DG, Swindell W, Morton EJ, Evans PM, Xiao ZR. A megavoltage CT scanner for radiotherapy verification. Phys Med Biol 1992;37:1985–99. 7. Jaffray DA, Siewerdsen JH, Wong JW, Martinez AA. Flatpanel cone-beam computed tomography for image guided radiation therapy. Int J Radiat Oncol Biol Phys 2002;53: 1337–49. 8. Mackie TR, Holmes T, Swerdloff S, et al. Tomotherapy: a new concept for the delivery of dynamic conformal radiotherapy. Med Phys 1993;20:1709–19. 9. Miralbell R, Lomax A, Cella L, Schneider U. Potential reduction of the incidence of radiation-induced second cancers by using proton beams in the treatment of pediatric tumors. Int J Radiat Oncol Biol Phys 2002;54: 824–9. 10. Followill D, Geis P, Boyer A. Estimates of whole-body dose equivalent produced by beam intensity modulated conformal therapy. Int J Radiat Oncol Biol Phys 1997;38: 667–72. 11. Mutic S, Low DA. Whole-body dose from tomotherapy delivery. Int J Radiat Oncol Biol Phys 1998;42:229–32. 12. Verellen D, Vanhavere F. Risk assessment of radiationinduced malignancies based on whole-body equivalent dose estimates for IMRT treatment in the head and neck region. Radioth Oncol 1999;53:199–203. 13. Kry SF, Salehpour M, Followill DS, et al. The calculated risk of fatal secondary malignancies from intensity-modulated radiation therapy. Int J Radiat Oncol Biol Phys 2005;62: 1195–203. 14. NCRP. Limitation of exposure to ionizing radiation. NCRP Report 116. Bethesda, MD 20814: National Council on Radiation Protection and Measurements, 1993.

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15. Harrison RM. Second cancers following radiotherapy: a suggested common dosimetry framework for therapeutic and concomitant exposures. Br J Radiol 2004;77:986–90. 16. UK CR. Cancer incidence, survival and mortality in the UK and EU. In: Toms JR, editor. CancerStats Monograph 2004. Cancer Research UK, 2004. 17. Brenner DJ, Curtis RE, Hall EJ, Ron E. Second malignancies in prostate carcinoma patients after radiotherapy compared with surgery. Cancer 2000;88:398–406. 18. Baxter NN, Tepper JE, Durham SB, Rothenberger DA, Virnig BA. Increased risk of rectal cancer after prostate radiation: a population-based study. Gastroenterology 2005;128:819–24. 19. McCall RC, McGinley PH, Huffman KE. Room scattered neutrons. Med Phys 1999;26:206–7. 20. McCall RC, Jenkins TM, Shore RA. Transport of accelerator produced neutrons in a concrete room. IEEE Trans Nuc Sci 1979;NS-26:1593–602. 21. Followill DS, Stovall MS, Kry SF, Ibbott GS. Neutron source strength measurements for Varian, Siemens, Elekta and General Electric linear accelerators. J App Clin Med Phys 2003;4:189–94. 22. Palta JR, Hogstrom KR, Tannanonta C. Neutron leakage measurements from a medical linear accelerator. Med Phys 1984;11:498–501. 23. Ongara C, Leon JR, Perez J, Zanini A, Burn K. Monte Carlo simulation and experimental evaluation of photoneutron spectra produced in medical linear accelerators. Proceedings of the 1999 Particle Accelerator Conference. New York: IEEE, 1999. 24. Golikov VY, Nitikin VV. Estimation of the mean organ doses and the effective dose equivalent from RANDO phantom measurements. Health Physics 1989;56:111–5. 25. Huda W, Sandison GA. Estimation of mean organ doses in diagnostic radiology from RANDO phantom measurements. Health Physics 1984;47:463–7. 26. Harrison RM, Wilkinson M, Shemilt A, Rawlings DJ, Moore M, Lecomber AR. Estimating second cancer risk following radiotherapy: organ doses from prostate radiotherapy and concomitant exposures. Biomedizinishe Technik 2005;50(Suppl. vol 1 Part 1):768–9. 27. Lebesque JV, Bruce AM, Kroes AP, Touw A, Shouman RT, van Herk M. Variation in volumes, dose-volume histograms, and estimated normal tissue complication probabilities of rectum and bladder during conformal radiotherapy of T3 prostate cancer. Int J Radiat Oncol Biol Phys 1995;33:1109–19. 28. Waddington SP, McKenzie AL. Assessment of effective dose from concomitant exposures required in verification of the target volume in radiotherapy. Br J Radiol 2004;77: 557–61.



Intensity-modulated radiation therapy in the treatment of gastric cancer: early clinical outcome and dosimetric comparison with conventional techniques 1,2

M T MILANO, MD, PhD, 3M C GAROFALO, MD, 1S J CHMURA, MD, PhD, 1K FARREY, MS, 1C RASH, CMD, R HEIMANN, MD, PhD and 1A B JANI, MD

4

1

Department of Radiation and Cellular Oncology, University of Chicago, Chicago, IL 60637, Department of Radiation Oncology, University of Rochester, Rochester, NY 14642, 3Department of Radiation Oncology, University of Maryland, Baltimore, MD 21201 and 4Division of Radiation Oncology, University of Vermont, Burlington, VT 05401, USA

2

ABSTRACT. The purpose of this study was to assess the efficacy and toxicity of intensitymodulated radiation therapy (IMRT) in the treatment of gastric cancer. Seven patients with gastric cancer were treated with IMRT. Six patients (all Stage III) received postoperative chemoradiotherapy with concurrent 5-fluorouracil and leucovorin. One received planned pre-operative radiation, though did not proceed to surgery. All patients were planned to receive 50.4 Gy in 1.8 Gy fractions. IMRT planning was compared with opposed anterior-posterior: posterior-anterior (AP/PA) and 3-field conventional three-dimensional plans. When compared with either AP/PA or 3-field plans, IMRT significantly reduced the volume exceeding the threshold dose of the liver and at least one kidney. Target coverage with IMRT was excellent, with 98¡1% of the target receiving >100% of the dose. Compared with AP/PA and 3-field plans, IMRT plans had a greater percentage of target receiving the prescribed dose, but also a greater volume receiving .110% of the dose. IMRT was well tolerated; no patients developed acute gastrointestinal toxicity greater than grade 2. All seven experienced grade 2 nausea, three had grade 2 diarrhoea and two had grade 2 oesophagitis. Weight loss ranged from 0–12% (mean 6.1% and median 5.8%). IMRT in the treatment of gastric malignancies reduces the mean and above threshold doses to critical normal tissues. In an initial cohort of seven patients, 50.4 Gy delivered by IMRT is well tolerated and safe.

Gastric carcinoma is an uncommon malignancy in North America, yet it represents the eighth leading cause of cancer death in the USA. In 2004, approximately 21 860 new cases of gastric cancer were estimated to occur in the USA and approximately 11 550 deaths are expected as a result [1]. Due to the lack of a cost-effective screening tool, gastric cancer is often diagnosed at an advanced stage. Surgery is the cornerstone of the treatment for resectable advanced stage gastric cancer. Post-operatively, the locoregional failure rate is approximately 40% and approximately half of those failures represent the only site of failure [2, 3]. In the setting of locally advanced (T3–4 and/or node positive) non-metastatic disease, adjuvant chemoradiotherapy has been established as the standard of care in the USA based upon the recently reported results of the Gastric Surgical Adjuvant Trial Intergroup 0116 trial [4]. This trial demonstrated a statistically significant benefit in relapse-free survival (48% versus 31%, p,0.001) and overall survival (50% versus 41%, p50.01) with the use of adjuvant Address correspondence to: Ashesh B Jani, University of Chicago, Department of Radiation and Cellular Oncology, MC 9006, Chicago, IL 60637, USA.


Received 7 June 2005 Revised 23 August 2005 Accepted 7 October 2005 DOI: 10.1259/bjr/43441736 ’ 2006 The British Institute of Radiology

chemoradiotherapy when compared with surgery alone. Adjuvant therapy also reduced the percentage of failures attributable to local failure (29% vs 19%). Since this landmark trial was reported, the increased use of radiation therapy in the adjuvant treatment of gastric cancer prompted the publication of a consensus report reviewing the details related to radiotherapy delivery technique [5]. According to this consensus report, ‘‘parallel-opposed AP/PA fields are considered the most practical arrangement for the overwhelming majority of post-operative adjuvant radiotherapy cases’’. The toxicity associated with adjuvant chemoradiation using traditional techniques is significant. Treatment volumes in the post-operative setting are necessarily large to account for the patterns of failure established in previous surgical studies [2, 3]. Typical target volumes include the stomach bed (to include surgical clips), a portion of the left hemi-diaphragm, and draining lymphatics at risk. The standard target dose of 45 Gy well exceeds the tolerance of several surrounding critical normal tissues (most notably the kidneys and liver). As a result, conventional treatment volumes are often tailored out of the fear of potential kidney and liver damage. By underdosing portions of the target, local control and survival 497

M T Milano, M C Garofalo, S J Chmura et al

may be compromised. Through inverse planning, intensity-modulated radiation therapy (IMRT) allows for more conformal dose delivery and selective sparing of critical structures such as the kidneys and liver, and may therefore allow for more complete target coverage to full-dose. Locoregional control may be improved through better target coverage and treatment toxicity may be reduced through the use of IMRT. We previously published our experience of using IMRT in the treatment of pancreatic and anal malignancies [6, 7]. Since 2001, our institution has routinely treated gastric cancer patients with post-operative chemotherapy combined with IMRT. Here we report the early clinical outcome of our single institution experience of seven patients treated for gastric cancer with IMRT, with emphasis on toxicity outcome, and the results of a dosimetric comparison with traditional radiotherapy delivery techniques.

Materials and methods Between March 2001 and April 2004, seven consecutive non-metastatic patients with adenocarcinoma of the stomach were treated with IMRT. Six patients received post-operative chemoradiotherapy (CRT) with concurrent 5-fluorouracil and leucovorin, using chemotherapy doses and scheduling as described in the recently reported Intergroup trial [4]. Two of these six patients, treated in early 2001, were not offered additional chemotherapy following radiation based on physician preference. One 87-year-old patient with symptomatic bleeding received planned pre-operative radiation (without concurrent chemotherapy because of comorbidities), since upfront surgery would have been delayed until a pacemaker revision. However, she never proceeded to surgery because of the comorbidities. All patients were planned to receive 50.4 Gy in 1.8 Gy fractions, except for the patient treated pre-operatively, who was prescribed 45 Gy to an initial planned target volume (PTV) with a 5.4 Gy boost to the gross tumour. For the purpose of dosimetric comparisons (described below) this patient was planned similarly to the others (i.e. 50.4 Gy with no boost PTV). The higher dose of 50.4 Gy was chosen in an attempt to improve locoregional control, and because the IMRT dose–volume histograms (DVHs) suggested that it could be safely delivered. All six patients who underwent resection had a D2 resection, including subtotal (n54) or total gastrectomy (n52), resection of perigastric and second echelon lymph nodes as described by the Japanese classification [8, 9].

Simulation and target contouring Patients underwent CT based simulation in the supine position (PQ5000 CT Simulator; Marconi Medical Systems, Cleveland, OH) with 4 mm CT slices. A custom immobilization device (Alpha Cradle; Smithers Medical Product, Inc., North Canton, OH) was used to minimize set-up variability. The pre-operative CT scan was image correlated to the CT simulation scan using the AcQSim VoxelQ software package. The PTV and normal structures (kidneys, liver and spinal cord) were manually contoured onto the CT scan slices following the ICRU 50 recommendations [10]. The clinical target volume (CTV) was contoured on axial CT scan slices. The CTV typically included the original tumour volume, operative bed (as defined by the operative note, pathologic findings, surgical clips and discussion with the surgeon) and the draining lymphatics at risk. The gross tumour volume (GTV) was entered on the single patient treated prior to planned surgery. The radiation dose was prescribed to a PTV, which was generated by expanding the CTV by 1 cm. The PTV design incorporated set-up uncertainty and organ motion [11]. Normal structures were also entered, including the kidneys, liver and spinal cord.

IMRT planning IMRT plans were generated using commercial inverse planning software (CORVUS, versions 3.0-5.0; NOMOS Corp., Sewickley, PA), which produces optimal intensity-modulated profiles using a simulated annealing algorithm. Dynamic multileaf collimators were used to shape the fields. Eight to nine-field coplanar plans were used. Typical angles were 40 ˚ increments starting from 0 ˚ to 320 ˚ (US convention). However, posterior fields that overdosed one kidney were usually removed (hence making an 8-field plan) or angled away from the kidney. The PTV and normal structure dose–volume constraints were iteratively adjusted to ensure optimal target coverage while minimizing dose to the kidneys, liver and spinal cord, thus optimizing the PTV and normal tissue DVHs [12]. Typical input parameters were defined on an individual basis for each patient, and thus varied from patient to patient, depending on the geometry of the PTV relative to the normal structures. Table 1 exemplifies the input parameters for the normal structures in two different patients. Generally, the input parameters for the PTV were ,2% of the PTV receiving ,50.4 Gy with a maximum of 54.0 Gy.

Table 1. Typical input parameters for normal structures for IMRT planning Structure

Right kidney Left kidney Liver Spinal cord

Patient 1

Patient 2

% of tissue volume to receive .dose (Gy)

Maximum dose to structure (Gy)

% of tissue volume to receive .dose (Gy)

Maximum dose to structure (Gy)

,35% ,40% ,80% ,65%

35 48 45 35

,5% . 8 ,5% . 8 ,45% .35 ,30% .35

50 50 50 42

.18 .20 .25 .18

IMRT, intensity-modulated radiation therapy.

498


IMRT in the treatment of gastric cancer

The IMRT plans were optimized to minimize the volume of PTV receiving ,95% of the prescribed dose and the volume receiving . 110% of the prescribed dose. Generally, the prescription dose was normalized to the 90–95% isodose line. The initial plans were considered acceptable with ,2–4% of the PTV receiving ,100% of the prescribed dose, ,5% of the PTV receiving .110% and ,1% receiving .115%. Because the normal tissue DVHs generated by CORVUS ignores regions of overlap with the PTV, the CORVUS IMRT plans were exported into PlanUNC [13] for the purpose of dosimetric comparison with conventional radiation plans.

Table 2. Patient characteristics Number Age (years) T stage T2 T3 N stage N0 N1 N2 Stage IIIa Stage IIIb Not surgically staged Total gastrectomy Subtotal gastrectomy Negative margins Positive margins

7 54.5–86.6 average 62.8 3 4 0 3 4 5 1 1 2 4 4 2

Comparison of three-dimensional AP/PA plans to IMRT plans Three-field (3F) and opposed anterior-posterior: posterior-anterior (AP/PA) (2F) 3DRT plans were generated for comparison with IMRT plans for the last 6 of 7 patients treated with IMRT (for one patient, the CT scan with contoured structures was unable to be recovered from the electronic archives). The 3F plan included a left lateral and AP/PA fields. 3F and 2F plans were generated using PlanUNC [13]. Segmented fields, variable weighting of fields and wedges were used to optimize the plan so as to improve dose homogeneity. With the 3F plan, plans were chosen to minimize right kidney dose without compromising PTV coverage. At a minimum, 2F and 3F plans were acceptable only if .95% of the volume received .98% of the prescribed dose. Normalization was typically set at 99–100% of the prescription dose. Custom blocks were used (1 cm margin around the PTV in each beam’s eye view). All fields were coplanar. 6 MV and 18 MV photons were used with the 3F and 2F plans, while 6 MV photons were used with the IMRT plans. DVHs were obtained for the PTV, kidneys, liver and spinal cord. Acute toxicity was scored using RTOG morbidity scoring criteria [14]. Dosimetric endpoints for the target and critical structures were compared using the two tailed paired t-test.

Results Seven patients with gastric cancer were treated with IMRT. The patient characteristics are outlined in Table 2. All six post-operatively treated patients had pathological stage IIIa-IIIb disease. Tumours were located in the antrum (n54) and lesser curvature (n53).

Dosimetric comparison between conventional and IMRT plans To demonstrate the differences in dose distribution, Figure 1 shows isodose curves on an axial slice for one patient for IMRT, 2F and 3F plans. Additionally, Figure 2 shows the DVH curves for the organs at risk for one representative patient. Table 3 summarizes the mean doses to the PTV, kidneys and liver for 2F, 3F and IMRT plans. Compared with the 2F plan, IMRT significantly The British Journal of Radiology, June 2006

reduced the mean dose to the left kidney. Compared with the 3F plan, IMRT significantly reduced the mean dose to the liver and right kidney, while the reduction in the dose to the left kidney was not significant. Table 4 summarizes the dosimetric endpoint analysis. Table 4a summarizes the volume of the PTV receiving greater than 50.4 Gy and 55.4 Gy. Table 4b summarizes the volume of critical structures receiving greater than the threshold dose [15]. Compared with 2F planning, IMRT significantly reduced the volume of left kidney receiving .20 Gy. Compared with 3F planning, IMRT significantly reduced the volume of right kidney receiving .20 Gy, and the volume of liver receiving .30 Gy. The mean PTV dose as well as the volume receiving .55.4 Gy were greater for the IMRT plans, though compared with 3F, neither parameter was significantly different. The maximum doses for 2F, 3F and IMRT plans were 109%¡3%, 108%¡3% and 115¡2%, respectively, with IMRT being significantly greater than 2F of 3F.

Clinical/toxicity outcome In this small cohort of patients, 3 of 7 are long term survivors (.2 years after diagnosis) and remain without evidence of disease. One died more than a year after diagnosis, after developing a malignant pleural effusion at 9 months. Two died from metastatic disease and rapid deterioration at 7 months and 18 months following diagnosis (both patients did not receive chemotherapy immediately following radiation); it is not known if these two patients had a component of local and/or regional failure. The 87-year-old patient treated with planned preoperative radiation had a stroke preceded by a fall and development of an interventricular haemorrhage (5 months after completing radiotherapy) and died 8 months after diagnosis. All patients completed their planned course of treatment with no planned or unplanned treatment breaks and no reduction or discontinuation of chemotherapy. Chemoradiotherapy with IMRT was well tolerated with no grade > 3 acute toxicity occurring during radiotherapy. Acute gastrointestinal and haematological toxicity is summarized in Table 5. Acute weight loss (up to 1 month post-radiation) ranged from no loss (1 patient) to a maximum of 12.3% weight loss, with a mean 499

M T Milano, M C Garofalo, S J Chmura et al

(a)

(b)

(c) Figure 1. Isodose curves on an axial slice for a representative patient for: (a) 2F (AP/PA) plan, (b) 3F plan, and (c) IMRT plan.

percentage of 6.1%¡4.7% and median of 5.8%. Skin toxicity was grade 0–1 in all patients. Of the four patients who received chemotherapy after radiation, two (50%) had grade 3 haematological (WBC) toxicity, occurring after the last dose of chemotherapy in both patients. No abnormalities were detected on laboratory assessment of kidney function (as compared with pre-treatment/baseline values) for any of the seven patients, either during treatment or on last available clinic follow-up visit; one patient had elevated liver enzymes approximately 4 months after the completion of RT. Although no late toxicity has been seen, the number of long term survivors is too small in this series to permit detailed analysis.

Discussion IMRT in the treatment of gastric cancer has the potential of lowering treatment related toxicity. To date, little has been published on the use of IMRT with gastric cancer. The University of Heidelberg has published two papers comparing the dose distribution of IMRT versus other planning modalities. In the first study, one patient 500

was planned to 45 Gy with the following approaches: 8 field step and shoot IMRT (planned with CORVUS using multiple couch angles), conventional 4-field box (4F), 4F with an off-kidney boost after 16.2 Gy and a noncoplanar 4F plan with a 90 ˚ couch kick and gantry angle, to direct the beam off of the caudal kidneys [16]. IMRT reduced the median dose to the kidneys (particularly the left kidney) and liver; the doses that were exceeded by 30% and by 60% of the volumes (both liver and kidneys) were also reduced with IMRT. In a follow-up paper with 15 patients, IMRT was compared against AP/PA, 4F and serial tomotherapy [17]. IMRT reduced dose to the left kidney (at the expense of greater spinal cord dosing) as compared with 4F, with a slight reduction as compared with AP/PA. Compared with 4F, IMRT more consistently produced high quality plans, as evidenced by the reduced standard deviation of the mean doses and doses that were exceeded by 30% and by 60% of the volumes. The dosing of normal tissues was not significantly different between IMRT and serial tomotherapy, though serial tomotherapy was superior with respect to conformality and homogeneity. Princess Margaret Hospital planned 20 gastric cancer patients to receive 45 Gy with conventional 5F coplanar The British Journal of Radiology, June 2006


(a)

(b)

(c)

(d)

Figure 2. Dose–volume histogram (DVH) curves for the organs at risk for a representative patient. (a) Liver, (b) spinal cord, (c) left kidney, and (d) right kidney. [y-axis: %Volume; x-axis: Dose (Gy); Colour Scheme- Blue: opposed anterior-posterior: posterior-anterior (AP/PA) plan; Red: 3F Plan; Green: IMRT plan].

Table 3. Mean dose to structures (percentage of prescribed dose) Structure

PTV Right kidney Left kidney Liver

2F

102.0¡0.7% 24.8¡14.9% 62.7¡30.2% 41.6¡9.5%

3F

103.5¡1.6% 26.7¡13.0% 47.3¡24.2% 67.9¡17.8%

IMRT

p-values*

105.7¡1.3% 18.9¡9.6% 38.4¡19.1% 44.6¡6.4%

2F vs 3F

2F vs IMRT

2F vs 3F

NS 0.021 NS 0.019

0.0014 NS 0.0049 NS

0.044 NS NS 0.0065

Values are given as mean¡standard deviation. 2F, AP/PA two-field three-dimensional plan; 3F, three-field conformal three dimensional plan; NS, not significant (p.0.05); IMRT, intensity-modulated radiation therapy. *Two-tailed paired t-test.

plans versus 7–9 field IMRT plans, generated by the CADPLAN Helios planning system [18]. Three reviewers examined the plans, with IMRT being preferred in 89%, with subjectively better PTV coverage in 86%, kidney The British Journal of Radiology, June 2006

sparing in 69% liver sparing in 71%, and spinal cord sparing in 74%. An Australian study compared AP/PA with a 3D conformal technique using a mono-isocentric split field 501

M T Milano, M C Garofalo, S J Chmura et al Table 4. Volume of structures receiving greater than the threshold dose for 50.4 Gy treatment (a) Target Volume above given dose (%) Structure

PTV PTV

p-values*

Dose (Gy)

55.4 50.4

2F

3F

IMRT

3F vs IMRT

2F vs IMRT

2F vs 3F

1.3¡2.4 84.1¡11.1

0.7¡1.6 94.9¡4.0

4.4¡3.9 98.1¡1.2

NS NS

NS 0.016

NS 0.022

(b) Organs at Risk Right kidney

20.0

20.1¡18.0

20.9¡16.6

11.6¡8.3

0.027

NS

NS

Left kidney Liver Spinal cord

20.0 30.0 45.0

67.6¡34.7 33.7¡12.6 55.0¡22.3

52.0¡31.9 63.6¡28.5 4.1¡7.3

52.1¡28.9 18.9¡12.9 0

NS 0.010 NS

NS 0.030 0.00039

NS 0.0062 0.00033

Values are given as mean¡standard deviation. 2F, AP/PA two-field three-dimensional plan; 3F, three-field conformal three-dimensional plan; NS, not significant (p.0.05); IMRT, intensity-modulated radiation therapy. *Two-tailed paired t-test.

Table 5. Acute gastrointestinal (GI) and haematological toxicity Grade

0

1

2

.2

Upper GI Lower GI Oesophagitis Haematological*

0 3 4 2

0 1 1 2

7 3 2 3

0 0 0 0

*Toxicity during radiation.

technique [19]. The 3D conformal technique improved mean and threshold dose to the kidneys and spinal cord, but not the liver. Memorial Sloan Kettering recently investigated the use of IMRT in four patients with gastric lymphoma, whose PTV had a high degree of overlap with the kidneys. IMRT tended to improve kidney and liver sparing [20]. The present study confirms the efficacy of IMRT in reducing kidney and liver doses. IMRT, planned with dynamic multileaf collimation, was compared with both the standard 2F AP/PA approach as well as a 3-field technique. As in the Heidelberg study, we used CORVUS planning software. Two differences are that our IMRT planning used dynamic multileaf collimation, and that the IMRT dose matrices were exported into PLUNC, allowing inclusion of normal structure volumes that overlap with PTV. We also prescribed to a higher dose, 50.4 Gy versus 45.0 Gy. Our data demonstrate that IMRT offers better sparing of the right kidney compared with conventional 3F planning, with significantly lower mean dose and volume above threshold dose [15]; and IMRT offers better sparing of the left kidney as compared with 2F planning, with lower mean dose and reduced volume above threshold dose (the latter of which was not significant). IMRT also affords liver sparing. As expected from the use of the additional field in the 3F arrangement, the liver dose increases dramatically with attempts to improve target coverage as compared with 2F planning. IMRT achieves similarly excellent target coverage as compared with 3F planning, while reducing the mean liver dose and volume above threshold dose [15]. In part, these improvements in relation to kidney and 502

liver dose reduction with the use of IMRT may be due to reducing entrance and exit dose to these organs from beam angle selection. To our knowledge, this study is the first to report clinical outcome in gastric cancer patients treated with IMRT. With respect to acute toxicity, patients fared remarkably well (albeit in a very small series), particularly since a higher than standard dose (50.4 Gy vs 45.0 Gy) was administered. There was no grade 3 or greater toxicity during radiation and all patients completed their planned course of chemoradiotherapy. In the Intergroup trial, 33% experienced grade 3 or greater gastrointestinal toxicity and 64% completed the treatment as planned. Grade 3 haematological toxicity was seen in the week following post-radiation chemotherapy in 2/4 patients, a percentage similar to that seen in the Intergroup trial (54% grade 3 haematological toxicity). In summary, IMRT offers improved sparing of normal structures, allowing a dose of 50.4 Gy to be delivered to the PTV. In this small series of patients, treatment was well tolerated with acceptable toxicity, seemingly improved as compared with the Intergroup trial. IMRT in the treatment of gastric cancer warrants further study.

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Changes in applicator positions and dose distribution between high dose rate brachytherapy fractions in cervix carcinoma patients receiving definitive radiotherapy 1

˘ AOG˘LU, MD, 1N TUNÇEL, PhD, 1M G DALMAZ, M GARIPAG ˘ , BSc and 1F G KO ˘ LU, MD ¨ KIZILDAG ¨ SEOG AU

MD,

2

¨ LKESEN, H GU

MD,

1

A TOY,

BSc,

1

Akdeniz University School of Medicine, Departments of 1Radiation Oncology and 2Bioistatistics, Antalya, 07070 Turkey ABSTRACT. This study examines the change of applicator geometry and its effect on rectal/rectum (R) and bladder (B) doses, and obtained radiobiological equivalent doses (RED), between each high dose rate (HDR) brachytherapy (BT) fraction in cervical carcinoma patients. BT using a tandem (T) and two ovoids (O) is included, and any discrepancies in applicator positions among the fractions were calculated. Whether the change of applicator position had an effect on the calculated R and B doses was analysed. Furthermore, the relationship between the size of tumour, the magnitude of displacement and the change in R and B doses was also investigated. Lastly, the changes in R and B RED were noted. The average magnitude of displacement was between 2.0 mm and 16.9 mm, showing time trend. There was no relationship between tumour size and the magnitude of discrepancy of Left O, Right O, T, R, B, and neither change in R and B doses (p.0.05). The mean differences of R and B doses were between 49–78 cGy, and 70–84 cGy, respectively. The magnitude of discrepancy and changes in doses showed no correlation (p.0.05). There were no significant differences in REDs for bladder (p50.8) and rectum (p50.2). In conclusion, there were significant differences in the applicator positions R and B and R and B doses among the fractions, which confirm the necessity of treatment planning in each HDR BT fraction. However, the total calculated R and B REDs did not show a remarkable difference.

Radiotherapy consisting of external radiotherapy and brachytherapy (BT) is the only curative treatment method for locally advanced stage cervical carcinoma [1–4]. The natural steep dose gradient of brachytherapy dose distribution allows a higher dose to the target while not exceeding the tolerance limit of normal rectal (R) and bladder (B) tissues. The optimal integration of brachytherapy and external radiotherapy is the main defining factor of radiotherapy treatment characteristics, such as the total point A dose, total paracentral dose and total treatment time which are independent prognostic factors in the treatment of cervical carcinoma as well as tumour and patient related factors [1, 2, 5–7]. It was claimed that in comparison with low dose rate (LDR) BT, high dose rate (HDR) BT has a physical dose distribution advantage. However, some authors claim that HDR BT has a radiobiological disadvantage [8–15], although performing HDR BT with multiple small fractions could alter this radiobiological disadvantage [5, 11, 16, 17]. Critical organs, namely R and B, are very close to the applicators and high dose region of BT. A small change in the distance from a particular point to the active sources may cause a great dose difference due to the BT Address correspondence to: Dr Melahat Garipag˘aog˘lu, Acibadem Oncology and Neurological Science Hospital, Department of Radiation Oncology, Inonu cad. Okur sok. No:20 Kozyatag˘ı 34742 Istanbul, Turkey.

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Received 28 January 2005 Revised 24 October 2005 Accepted 5 December 2005 DOI: 10.1259/bjr/33762931 ’ 2006 The British Institute of Radiology

having a large dose gradient near the region close to the active sources [14, 18]. Therefore, a precise dose calculation of normal tissue as well as the target is critical for successful treatment in HDR BT. The aim of this study is to detect the change of applicator geometry between HDR BT fractions and its effect on the calculated R and B doses, and the given radiobiological effective doses (RED)s, using tandem (T) and ovoid (O) in cervical carcinomas.

Methods The orthogonal radiographs of patients receiving intracavitary brachytherapy using T and two Os were used for this retrospective study. External radiotherapy was administered with a 1.8 Gy fraction size, totalling 46–50.4 Gy to the whole pelvis. BT was done five times using Ir192 Microselectron HDR machine with 6 Gy fraction size, starting after the completion of at least 40 Gy of external radiotherapy in order to achieve enough tumour regression to perform an optimal application, and BT fractions were used twice a week to reduce the total treatment time. The application was done after the patient had fasted for 12 h. Laxatives were used for the elimination of rectal matters, urinary Foley catheter was used to keep the bladder empty during the whole procedure, and the The British Journal of Radiology, June 2006

Position and dose variations among HDR brachytherapy fractions

balloon of the Foley catheter was filled with 7 cm3 of an opaque solution. Before the BT application, the same procedure of sedation and analgesia was given in all fractions of each patient. An extra analgesic was given when it was needed after the applicator placement. After the completion of the application, one tandem and two ovoids were fixed to each other and a specially designed table used until completion of treatment, to ensure the immobilization of the patient during simulation, treatment and transport. In addition, all the applications were given by the same experienced physician using the same kind of packing technique. Orthogonal X-rays were taken for each fraction. A reference volume definition for R and B dose calculations were carried out according to ICRU-38 recommendations. Treatment planning and optimization were applied in each fraction. Special effort was made not to exceed rectal and bladder tolerance levels. In order to determine the changes of applicator positions between the BT fractions, pelvic bone reference points were used as fixed points, applicators (T, left O, right O), B and R reference points were used as non-fixed points (Figure 1). A Cartesian coordinate centre (for x, y and z axes) was chosen using orthogonal radiographs (anterior-posterior (AP) and lateral) (Figure 1). The position of fixed and non-fixed points were measured according to centre of Cartesian coordinate centre in x, y and z axes. The distances between the fixed and nonfixed points were measured on the x and y axes on AP

radiographs and in z axis on lateral radiograms. The differences in the measured distances were calculated between the first fractions and subsequent fractions. Whether the magnitude of displacement showed a time trend was investigated. Furthermore, the magnitudes of the displacements on the x, y and z axes were used to calculate the resultant vector. The relationship between the initial tumour size and the magnitude of resultant vectors was also examined. For this study, to evaluate whether the change of applicator positions to the bony pelvis has an effect on the calculated R and B doses, both active source positions and calculated treatment time of first fraction were repeated for subsequent fractions of each patient, and inferential R and inferential B dose calculations were made. Then the difference between the ICRU-38 recommended R and B doses in the first fraction and the subsequent inferential R and B doses was determined [19]. The relationship between changes in subsequent inferential R and B doses and initial tumour size obtained from MR images was also evaluated. It was assumed that geometric variations might lead to a change in calculated doses. The relationship between the magnitude of displacement of the applicators and the inferential doses were examined. In an effort to determine whether changes in R and B doses have an effect on R and B REDs, R and B REDs of each fraction were calculated using inferential R and B doses, according to the ‘‘linear quadratic model,’’ then added [20]. The chosen alpha/beta value was 3 for B and R, as in previous studies [21]. The change in total REDs

Figure 1. The fixed points, which are pelvic bone reference points; the non-fixed points, namely left ovoid (LO), right ovoid (RO), tandem stopper (S), bladder (B) and rectal (Rg) reference points are shown on (a) anterior-posterior and (b) lateral diagram. The British Journal of Radiology, June 2006

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of R and B calculated in concordance with the linear quadratic model was investigated.

Statistical analysis Statistical analysis was performed using SPSS 10.0 software, the normality of samples was analysed by a Shapiro-Wilk test. Statistical significance of the displacement and dose changes between fractions was tested using the paired t-test or a Wilcoxon signed-rank test. Spearman’s rho test was used for correlations; p,0.05 is considered to be statistically significant.

Results 13 out of 75 total fractions were excluded from the study because of low quality X-rays or the use of the applications other than tandem and ovoids. 15 distances of each fraction were measured on AP and lateral graphs, with a total of 930 distances recorded. The average magnitude of discrepancy in the x, y and z axes and resultant vector are presented in Table 1. The mean discrepancy of left O, right O, T, B and R reference points between the fractions were between 10.0 mm and 19.4 mm. In the first fraction, the distance of the T from the non-fixed point was 37¡3 mm (mean¡standard deviation), and was 31¡3 mm in the fourth fraction (p50.012). In other words, the stopper of T moved 6 mm superior towards the end of the treatment. No other statistically significant time trend or movement to a certain direction of non-fixed points could be determined (data not shown). There was no significant relationship

between the initial tumour size and the magnitude of resultant vectors of left O (r50.239, p50.132), right O (r50.24, p50.13), T (r50.036, p50.821), R (r50.035, p50.834) and B (r50.309, p50.059). The differences calculated between inferential R and B doses and R and B doses in the first fraction are presented in Table 2. The mean differences for B and R were between 78–149 cGy and 70–84 cGy, respectively. A correlation was seen between initial tumour size and changes in inferential R doses (r50.414, p50.005), but no correlation was seen in the B doses (r5–0.075, p50.621). The magnitude of resultant vectors and changes in doses for R (r50.455, p50.005), and B (r5–0.418, p50.007) showed correlation (Figure 2). The difference between inferential dose and administered dose for B was greater than 60 cGy in 15 (100%) of the patients, and 6 (40%) of these patients had a difference greater than 120 cGy. The difference for the R was greater than 60 cGy in 10 (66%) patients; 6 (40%) of them had a difference greater than 120 cGy. The median of differences between total REDs and total inferential REDs for the R for each patient was 11 Gy, whereas it was 4 Gy for the B as BED values. The differences were not statistically significant.

Discussion Significant geometric variations were seen in all three applicators’, namely left O, right O and T positions, between the HDR brachytherapy fractions. As noted previously in either LDR or HDR brachytherapy, these position differences were more than 1 cm in 60% of the applications [22–29].

Table 1. Absolute discrepancy on x, y, z axes and displacement in resultant vector (mean, mm)

Second fraction (n515) Left ovoid Right ovoid Stopper Bladder Rectum Third fraction (n514) Left ovoid Right ovoid Stopper Bladder Rectum Fourth fraction (n511) Left ovoid Right ovoid Stopper Bladder Rectum Fifth fraction (n55) Left ovoid Right ovoid Stopper Bladder Rectum

Delta x

Delta y

Delta z

Resultant vector

95% CI

SD

10.0 11.4 9.8 8.9 8.4

5.5 5.5 7.6 2.0 5.9

8.3 7.3 5.7 4.9 6.8

16.6 17.2 16.9 12.5 13.9

11.0–22.2 11.2–23.2 11.3–22.5 8.1–16.9 9.9–17.9

10.8 11.8 10.9 8.1 7.4

11.3 12.3 10.1 11.0 8.3

6.2 6.4 6.5 4.2 6.8

7.9 6.5 6.4 8.3 5.9

16.9 17.6 16.7 15.8 15.4

10.9–22.9 11.8–23.4 11.7–21.7 7.8–23.8 11.0–19.8

11.1 11.0 9.2 14.3 8.2

9.5 12.7 7.7 10.9 7.2

6.1 6.3 6.8 3.7 6.1

9.9 10.9 9.9 5.8 7.7

16.7 19.4 15.6 15.1 13.1

12.1–21.3 14.0–24.8 8.6–22.6 5.3–24.9 9.5–16.7

7.5 8.8 11.7 15.4 5.6

14.0 16.9 13.7 6.3 9.2

4.2 5.9 8.2 4.3 8.9

5.7 3.2 4.1 8.5 8.1

16.8 17.1 17.6 10.0 17.3

5.6–28.0 6.9–27.3 7.6–27.6 5.4–14.6 11.3–22.9

12.5 11.4 11.1 4.7 6.1

CI, confidence interval; SD, standard deviation.

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Position and dose variations among HDR brachytherapy fractions Table 2. The difference between inferential rectal and bladder doses, calculated using first fraction active sources positions and treatment times

Second fraction (n515) Bladder Rectum Third fraction (n514) Bladder Rectum Fourth fraction (n511) Bladder Rectum Fifth fraction (n55) Bladder Rectum

Mean dose

95% CI

SD

92 77

37–147 35–119

99 73

149 84

79–219 34–135

126 88

78 83

36–119 15–152

69 113

99 70

38–160 24–116

79 55

CI, confidence interval; SD, standard deviation.

Furthermore, in comparison with T and R applications, the magnitude of differences was lower in the present study, compatible with other reports regarding T and O applications [22–26, 29]. Some authors claim that insufficient fixation of applicators both to each other and to the table is a main potential reason for geometric variation [26, 29]. Although the variations were higher in non-fixed applications than in fixed and the changes in O were higher than in T in non-fixed applications, there was still a significant variation in the present study. Research also showed that another reason for displacement was anaesthesia given among the fractions [27, 28]. In the present study, the same kind of anaesthesia was used in all fractions of each patient. However, there were significant geometric variations in the applicator positions among the fractions. In order to achieve enough tumour response to perform an optimal application, brachytherapy was

started after the completion of 40 Gy external radiotherapy. The stopper of tandem moved approximately 6 mm superior towards the end of the treatment. We speculate that this finding supports continuing tumour shrinkage in a period of brachytherapy fractions [24, 27, 29]. In comparison with the magnitude of changes in B position, the changes in R position were higher. The B was kept empty and the ICRU 38 recommended B reference point is related to the balloon of the Foley catheter, therefore showing that the change in B position was small. The ICRU 38 recommended R reference point is defined according to the position of the applicators, so the changes in applicator positions also affected the changes in ICRU 38 recommended R reference point, as expected [28]. The inferential R and B doses were helpful for predicting R and B doses of subsequent fractions when treatment planning and dose calculation were not performed for remaining fractions. In addition, positional differences were seen, and R and B doses differed among the fractions. (This difference was higher than 60 cGy, which is 10% of prescribed dose in 10 out of 15 patients in B doses and 15 out of 15 patients in R doses, and higher than 120 cGy, which is 20% of prescribed dose in 10 out of 15 patients in B doses and 15 out of 15 patients in R doses). The precise physical dose calculation is more important in HDR than LDR brachytherapy because of its claimed radiobiological disadvantage. Therefore, the prescribed dose should be adjusted considering normal tissue doses. This magnitude of differences in normal tissue dose illustrates the necessity of treatment planning and dose calculation in each fraction of HDR brachytherapy, which is in agreement with other studies [24, 25, 27–29]. Furthermore, there was a relationship between the initial tumour size and the change in R dose, as seen in the literature [24, 28]. However, there was no association between the B dose and the initial tumour size. Presumably, the reason is that the ICRU 38 recommended

Figure 2. Simple scatter plot graphics of relationship between resultant vector and dose change in the rectum (r520.418, p50.007).


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R reference point is defined according to Os location, therefore tumour shrinkage has an effect on R point. On the contrary, ICRU 38 recommended B reference point is dependent on Foley catheter balloon position without using the applicators. There were remarkable relationships between the magnitudes of geometric displacement (as a resultant factor) and dose variation for R and B. Both tumour control and radiation injury increased with increasing radiation doses. B and R are the main dose limiting organs in the pelvis. For this reason, brachytherapy is helpful for increasing central tumour doses while external radiotherapy with midline shielding is helpful for increasing parametrial doses without increasing R and B doses. The hypothesis of using brachytherapy to increase central tumour dose and of using external radiotherapy with midline shielding to increase parametrial doses, without increasing R and B doses has many shortcomings, such as the increase of rectal complications, the difficulty of defining vagina location, which kind of block technique (i.e. stepwedge or rectangle) should be used and what percentage of isodose is best to integrate with external radiotherapy successfully [3, 5, 28, 30–35]. Moreover, according to the results of the current study and previous studies, the calculated doses differ among HDR fractions using T and O [26–29]. For these reasons, the design of the midline shield according to isodose distribution of the first BT fraction could result in an inappropriate isodose distribution, due to changes in subsequent fraction’s isodose distribution. The total inferential R and B REDs were relatively similar. However, one should keep in mind that the acceptable R and B doses were calculated ICRU 38 recommended R and B reference points and based on the points, rather than volume. As shown in other studies, these doses do not represent the real R and B doses [36– 39]. On the other hand, volume based three-dimensional planning for each fraction is not possible in most centres using HDR. The results of the current study suggest that proper treatment planning, dose calculation and optimization according to normal tissue doses for each fraction are necessary in cervical carcinoma patients receiving HDR brachytherapy using T and O.

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Position and dose variations among HDR brachytherapy fractions 24. Hoskin PJ, Cook M, Bouscale D, Cansdale J. Changes in applicator position with fractionated high dose rate gynaecological brachytherapy. Radiother Oncol 1996;40:59–62. 25. Kim RY, Meyer JT, Plott WE, Spencer SA, Meredith RF, Jennelle RL, et al. Major geometric variations between multiple high-dose-rate applications of brachytherapy in cancer of the cervix: frequency and types of variation. Radiology 1995;195:419–22. 26. Kim RY, Meyer JT, Spencer SA, Meredith RF, Jennelle RL, Salter MM. Major geometric variations between intracavitary applications in carcinoma of the cervix: high dose rate vs. low dose rate. Int J Radiat Oncol Biol Phys 1996;35:1035–8. 27. Bahena JH, Martinez A, Yan D, Mele E, Edmunson G, Brown D, et al. Spatial reproducibility of the ring and tandem high-dose rate cervix applicator. Int J Radiat Oncol Biol Phys 1998;41:13–9. 28. Hellebust TP, Dale E, Skjonsberg A, Olsen DR. Inter fraction variations in rectum and bladder volumes and dose distributions during high dose rate brachytherapy treatment of the uterine cervix investigated by repetitive CTexaminations. Radiother Oncol 2001;60:273–80. 29. Datta NR, Kumar S, Das KJ, Pandey CM, Halder S, Ayyagari S. Variations of intracavitary applicator geometry during multiple HDR brachytherapy insertions in carcinoma cervix and its influence on reporting as per ICRU report 38. Radiother Oncol 2001;60:15–24. 30. Huang EY, Lin H, Hsu HC, Wang CJ, Chen HC, Sun LM, et al. High external parametrial dose can increase the probability of radiation proctitis in patients with uterine cervix cancer. Gynecol Oncol 2000;79:406–10. 31. Ting JY, Wolfson AH, Wu X, Fiedler JA, Abdel-Wahab M, Yang CC, et al. Bladder and rectal doses from externalbeam boosts after gynecologic brachytherapy. Radiology 1998;209:825–30.


32. Kim CR, Eaton BA, Stevens KR Jr. Localization of the apex of the vagina: implications for radiation therapy planning. Radiology 1999;212:155–8. 33. Roeske JC, Mundt AJ, Halpern H, Sweeney P, Sutton H, Powers C, et al. Late rectal sequelae following definitive radiation therapy for carcinoma of the uterine cervix: a dosimetric analysis. Int J Radiat Oncol Biol Phys 1997;37:351–8. 34. Chao KS, Williamson JF, Grigsby PW, Perez CA. Uterosacral space involvement in locally advanced carcinoma of the uterine cervix. Int J Radiat Oncol Biol Phys 1998;40:397–403. 35. Wolfson AH, Abdel-Wahab M, Markoe AM, Raub WJ Jr, Diaz D, Desmond JJ, et al. A quantitative assessment of standard vs. customized midline shield construction for invasive cervical carcinoma. Int J Radiat Oncol Biol Phys 1997;37:237–42. 36. Gebara WJ, Weeks KJ, Hahn CA, Montana GS, Anscher MS. Computed axial tomography tandem and ovoids (CATTO) dosimetry: three-dimensional assessment of bladder and rectal doses. Radiat Oncol Investig 1998;6:268–75. 37. Fellner C, Potter R, Knocke TH, Wambersie A. Comparison of radiography- and computed tomography-based treatment planning in cervix cancer in brachytherapy with specific attention to some quality assurance aspects. Radiother Oncol 2001;58:53–62. 38. Serkies K, Badzio A, Jereczek-Fossa B, Tarnawska Z, Nowak R, Szewczyk P, et al. Rectal doses in intracavitary brachytherapy of gynecological malignancies: comparison of two dosimetric methods. Radiother Oncol 2001;58:37–41. 39. Dale E, Hellebust TP, Skjonsberg A, Hogberg T, Olsen DR. Modeling normal tissue complication probability from repetitive computed tomography scans during fractionated high-dose-rate brachytherapy and external beam radiotherapy of the uterine cervix. Int J Radiat Oncol Biol Phys 2000;47:963–71.

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Clinical and cellular ionizing radiation sensitivity in a patient with xeroderma pigmentosum 1

C F ARLETT, PhD, 2P N PLOWMAN, MD, FRCP, FRCR, 2P B ROGERS, MB BS, MRCP, FRCR, 3C N PARRIS, PhD, 3 F ABBASZADEH, MSc, 4M H L GREEN, PhD, 5,6T J MCMILLAN, PhD, 6C BUSH, BSc, 7N FORAY, PhD and 1 A R LEHMANN, PhD 1

Genome Damage & Stability Centre, University of Sussex, Falmer, Brighton BN1 9RQ, Radiotherapy/Clinical Oncology, St Bartholomew’s Hospital, 25 Bartholomew Close, West Smithfield, London EC1A 7BE, 3Division of Biosciences, School of Health Sciences and Social Care, Brunel University, Kingston Lane, Uxbridge, Middlesex UP8 3PH, 4School of Pharmacy and Biomolecular Sciences, University of Brighton, Cockcroft Building, Lewes Road, Brighton BN2 4GJ, 5 Institute of Environmental and Natural Sciences, Faraday Building, Lancaster University, Lancaster LA1 4YA, 6The Institute of Cancer Research, Royal Cancer Hospital, 15 Cotswold Road, Sutton, Surrey SM2 5NG, UK, 7INSERM U647, ID17, European Synchrotron Research Facility, Rue Jules Horowitz, BP220 – 38043 Grenoble, France 2

ABSTRACT. XP14BR is a cell line derived from a xeroderma pigmentosum (XP) patient from complementation group C. The patient was unusual in presenting with an angiosarcoma of the scalp, treated by surgical excision and radiotherapy. Following 38 Gy in 19 fractions with 6 MEV electrons, a severe desquamation and necrosis of the underlying bone ensued, and death followed 4 years later. The cell line was correspondingly hypersensitive to the lethal effects of gamma irradiation. We had previously shown that this sensitivity could be discriminated from that seen in ataxiatelangiectasia (A-T). The cellular response to ultraviolet radiation below 280 nm (UVC) was characteristic of XP cells, indicating the second instance, in our experience, of dual cellular UVC and ionizing radiation hypersensitivity in XP. We then set out to evaluate any defects in repair of ionizing radiation damage and to verify any direct contribution of the XPC gene. The cells were defective in repair of a fraction of double strand breaks, with a pattern reminiscent of A-T. The cell line was immortalized with the vector pSV3neo and the XPC cDNA transfected in to correct the defect. The progeny derived from this transfection showed the presence of the XPC gene product, as measured by immunoblotting. A considerable restoration of normal UVC, but not ionizing radiation, sensitivity was observed amongst the clones. This differential correction of cellular sensitivity is strong evidence for the presence of a defective radiosensitivity gene, distinct from XPC, which is responsible for the clinical hypersensitivity to ionizing radiation. It is important to resolve how widespread ionizing radiation sensitivity is amongst XP patients.

Xeroderma pigmentosum (XP) is a rare autosomal recessive disease [1] characterized by clinical and cellular sensitivity to ultraviolet (UV) radiation. The patients show hypersensitivity to sunlight and extreme susceptibility to sunlight-induced skin cancer [2, 3]. The majority of cases are defective in one of seven genetically distinct nucleotide excision repair genes (complementation groups A–G [3]) which confer increased cellular susceptibility to UV-induced killing and mutation [4–7]. In addition, cells are also hypersensitive to agents which generate bulky lesions in DNA [8]. Approximately 20% of cases are competent in excision repair, but defective in daughter strand repair [9] by virtue of a defect in a Current address for Dr P B Rogers: Royal Berkshire Hospital, London Road, Reading, Berkshire RG1 5AN, UK.

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Received 28 October 2005 Revised 23 December 2005 Accepted 9 January 2006 DOI: 10.1259/bjr/83726649 ’ 2006 The British Institute of Radiology

DNA polymerase (g) [10]. Cells from such patients are minimally sensitive to the lethal effects of UV [11], but are hypermutable [12, 13]. Since ionizing radiation mainly produces single and double DNA strand breaks [14, 15] and types of base damage repaired by the alternative pathway of base excision repair [16, 17], hypersensitivity to ionizing radiation would not be anticipated as a feature in the XP syndrome. However, we have reported cellular hypersensitivity to the lethal effects of gamma radiation in fibroblasts from XP complementation group G, XP3BR [18] derived from a patient who never came to radiotherapy. There are two reports of patients from complementation group C who have been exposed to radiotherapy. DiGiovanna et al [19] recorded the uneventful radiotherapy of such a patient with an The British Journal of Radiology, June 2006

Ionizing radiation hypersensitivity in xeroderma pigmentosum

inoperable spinal cord astrocytoma, no assessment of cellular sensitivity to either UVC or ionizing radiation was provided (fibroblast cell line 5 XP23BE). In the second example [20], the patient died 3 months after a course of combined radiotherapy and chemotherapy for a thalmic glioma. Again, no assessment of cellular sensitivity was reported (fibroblast cell line 5 XP233VA). We detail here radiosensitivity at both the clinical and cellular level in an XP patient assigned to genetic complementation group C. The homozygous mutation in the XPC gene generates a stop codon at codon 718, resulting in a truncated protein missing the C-terminal 233 of the 940 amino acids. To date, this patient is unique in carrying this mutation (see XP mutation database: URL http://xpmutations.org). The cell lines XP23BE and XP233VA acted as reference controls. Earlier published observations reported the clinical hypersensitivity in this patient as one of a series of four sensitive individuals encountered amongst 2000 paediatric radiotherapy patients during a 20 year period at St Bartholomew’s Hospital [21]. Discrimination from ataxia-telangiectasia (A-T) was achieved on the basis of clinical criteria and by demonstrating a normal level of radiation resistant DNA synthesis. We report here details of the pathology consequent to radiotherapy, further characterization of the cell line as XP, DNA repair studies following ionizing radiation and evidence for the existence of a separate radiosensitivity genetic defect.

Materials and methods The patient A clinical description of the patient, a female of Pakistani origin, is provided in Salob et al [22] together with details of the significant reduction to 9% in nucleotide excision repair in fibroblasts (designated 86/ 0029) and assignment to complementation group C. She was considered unusual in presenting with both XP and hypoplastic anaemia. At the age of 14 years, she developed an angiosarcoma of the scalp overlying the right parieto-occipital bone, a feature thought to be unique in XP [23]. Previous skin (epithelial) tumours had been treated with excision or with 5-fluorouracil cream.

Cells Untransformed fibroblast cell lines from reference normal control donors, 1BR.3, 142BR, 149BR, 250BR and 251BR, were established in Brighton using standard procedures [24]. The normal cell line used at Sutton was GF11. Fibroblasts of the patient as used in the original complementation assignment [22] designated 86/0029 (5XP14BR.1), were obtained from Dr P Botcherby (Guy’s Hospital, London, UK). A second biopsy taken from an uninvolved region of skin during radiotherapy at St Bartholomew’s Hospital generated a second cell line, XP14BR.2, in Brighton. Reference XP complementation group C fibroblast cultures from the studies noted above were obtained from Dr K Kraemer, NIH, Bethesda, MD (XP23BE) [19] and Dr A Sarasin, Villejuif, France (XP233VI) [20]. XP15BR is a reference XP cell line, generated in Brighton, from complementation group A (N Jaspers, Rotterdam, personal communication) cell line. Two A-T cell lines AT1BR and AT2BR [25] were used for scaling purposes. All these cells were maintained in culture using standard materials and procedures [11, 24] and were verified to be free of mycoplasma infection. All biopsies were obtained under the approved ethical procedures in place in the relevant institution at the time. Fibroblast cell lines were transformed using the immortalizing vectors pSV3gpt and pSV3neo in Brighton in a programme using standard procedures [26]. Both pre-crisis (not immortal) and post-crisis (immortal) cultures of single cell origin were available for investigation. Reference normal material was generated using this vector from the cell lines 1BR.3 (51BR.3G, from pSV3gpt and 1BR.3neo from pSV3neo), and 142BR (5142BRneo, from pSV3neo) in addition to XP14BR (5XP14BRneo17, from pSV3neo). Immortalized control cell lines MRC5V1 (from Dr P Debenham, former MRC Radiobiology Unit, Harwell, UK) and GM0637 (from NIGMS, Camden, NJ, USA) were also employed. The untransformed fibroblast cell line MRC5 was obtained from C Babbs, Huntingdon Research. Table 1 summarizes the designations of the strains used.

Cell killing Details of sources and dosimetry for UVC and a Cobalt 60 gamma ray facility, together with the design of cell

Table 1. Designations of the cell lines compared in this investigation Cell line

Properties

XP14BR.1 (586/0029), XP14BR.2

Fibroblast lines derived from separate biopsies of the subject of this investigation SV40-transformed derivative of XP14BR.2 Reference normal fibroblast lines Reference SV40-transformed normal fibroblast lines Reference XP fibroblast lines Reference A-T fibroblast lines Derivatives of XP14BRneo17, transfected for the XPC gene. Clones T921A and T943A expressed significant XPC protein, and T941B a low amount

XP14BRneo17 1BR.3, 142BR, 149BR, 250BR, 251BR, GF11, MRC5 1BR.3-G, 1BR.3neo, 142BRneo, MRC5V1, GM0637 XP23BE, XP233VI, XP15BR AT1BR, AT2BR T921A, T943A, T941B, T948B

XP, xeroderma pigmentosum; A-T, ataxia-telangiectasia.


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killing experiments, have been described in detail elsewhere [24, 27]. The response to the cross-linking agent nitrogen mustard (Sigma Chemical Co., Poole, UK) was verified using standard protocols [28].

DNA repair We have used two methods for the evaluation of repair following ionizing radiation. In the first, double strand break (DSB) induction and repair were evaluated at the Institute of Cancer Research, Sutton, using pulsed field gel electrophoresis (PFGE) [29–31], with a repair time of 4 h. A separate series of experiments evaluating DSB induction and repair was performed at Institut GustaveRoussy, Villejuif, France. The experimental procedures are described in Foray et al [32]. A Caesium 137 radiation source was used for the Villejuif experiments. Here, the maximum repair time was 24 h. The control cell line was 149BR. A second method, based upon a process of reduction in the frequency of excess chromosome fragments [33], was also undertaken at Villejuif. The premature chromosome condensation (PCC) technique permits the measurement of chromosome breaks in cells in G0/G1. Cells were irradiated in plateau phase of growth (6 Gy, c-ray

137Cs, 4 ˚C) and were then returned to 37 ˚C for repair. Chromatin condensation was generated by fusion of fibroblasts in plateau phase of growth with synchronized mitotic cells by using PEG 6000, as described previously [33], but with the notable exception that CHO cells were replaced by HeLa cells. After hypotonic treatment, cells were fixed in methanol:acetic acid and stained with Giemsa. At least 30 fusions were analysed by means of light microscopy. PCC data were expressed in chromosome fragments in excess (ECF) (total number PCC fragments minus 46). The reference normal cell line was 149BR.

Transfection with the XPC gene The SV40-transformed XP14BRneo17 cell line was transfected with plasmid containing either the full-length XPC cDNA, generously supplied by P van der Spek (Rotterdam), or a clone truncated at the 5’ end of the open reading frame, which nevertheless can complement the UV sensitivity of an XPC cell line [34]. Each plasmid was cotransfected with pSVgpt using calcium phosphate precipitation and selection was applied for the cotransfecting gpt marker using medium containing mycophenolic acid, aminopterin and xanthine. When colonies became visible to the naked eye, they were picked individually using cloning rings and expanded maintaining the selection for the gpt marker.

XPC protein measurements Pellets of different cells were dissolved directly in Laemmli buffer and 15 mg protein run on 7.5% SDSPAGE gels, followed by immunoblotting. The blots were probed with a polyclonal antibody raised against a peptide containing the C-terminal 18 amino acids of the XPC protein [35] and developed with the ECL kit (Amersham).

Results Radiation pathology

Figure 1. The irradiated scalp portal. The irradiated skin remained depilated and became frankly ulcerated and necrotic in the operation wound site. 512

The angiosarcoma of the scalp was excised and the patient was referred for radiotherapy because of the closeness of the marginal excision. 6 MEV electrons were employed, which have a depth of penetration of approximately 2 cm in soft tissue and less in bone. Due to the characteristics of the electron beam, any further depth would be fast diminishing. After 38 Gy radiation applied in 19 fractions, the child developed severe moist desquamation in the region in the scalp (Figure 1). Ulcerated and underlying necrotic bone at the depth of the ulcer was revealed and this never healed. This reaction is very far outside the normal expected range of response to this clinical radiotherapy dose prescription. 2 years following radiotherapy, the patient suffered a Grand Mal convulsion, and a CT scan of the brain showed extensive cerebral oedema arising from an oedematous focus directly beneath the radiation portal The British Journal of Radiology, June 2006


of 22 months, but the child died from raised intercranial pressure and extending left hemisphere oedema 4 years after the radiotherapy. A diagnosis of ‘‘spreading cerebral necrosis, consequent upon the radiotherapy,’’ was made.

Cell killing

Figure 2. CT brain scan demonstrating widespread hemispheric oedema beneath the irradiated portal. Noteworthy is the widespread extent of the oedema throughout the hemisphere, not withstanding the restricted portal size, the moderate prescription dose (38 Gy) and the low energy electron portal (6 MEV) – that had to penetrate bone first. It was this unrelenting cerebral oedema, emanating from radio necrosis of brain under the radiation portal, that caused death.

(Figure 2). However, the area of the oedema continued to spread through most of the hemisphere over the next few months. Steroids were employed and taken for a period

Figure 3. Ultraviolet radiation below 280 nm (UVC) clonal survival of XP14BR fibroblasts. Key: XP14BR.1 (biopsy 1); XP14BR.2 (biopsy 2, mean ¡ standard error of 4 1BR.3 (mean experiments). Normal fibroblast lines: ¡ standard error of 3 experiments), 142BR, 250BR, 251BR. Additional XP cell lines: XP15BR, XP23BE. Single experiments except where indicated. The British Journal of Radiology, June 2006

The responses of XP14BR and a set of control and reference fibroblast cell lines to the lethal effects of UVC are illustrated in Figure 3. As anticipated XP14BR.2 fibroblasts and the earlier derived replicate culture, XP14BR.1 (86/0029) proved to be hypersensitive. Hypersensitvity was, as expected, also demonstrated with the reference cell lines XP23BE and XP15BR used for scaling purposes. With gamma irradiation, both versions of XP14BR were indistinguishable in their response and proved to be almost as sensitive as the two reference A-T cell lines (AT1BR and AT2BR). Two other cell lines derived from representatives of complementation group C (XP23BE and XP233VI), where no clinical sensitivity had been reported, were normal in their responses (Figure 4) as was XP15BR. With mitomycin C, no distinction in response between XP14BR.2 and the normal cell line 1BR.3 was observed (data not shown). For nitrogen mustard, a comparison between the immortalized XP14BRneo17 and the control lines

Figure 4. Ionizing radiation clonal survival of XP14BR fibroblasts. Key as for Figure 3: Additional XP line; XP233VA; Ataxia telangiectasia fibroblast lines; AT1BR; AT2BR. Means ¡ standard error based on numbers of experiments as follows: XP14BR.1, 1; XP14BR.2, 6; 1BR.3, 58; 142BR, 8; 250BR, 4; 251BR, 13; XP23BE, 2; XP233VA, 2; AT1BR, 5; AT2BR, 3.

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Figure 5. Nitrogen mustard sensitivity of SV40 transformed XP14BR fibroblasts. Key: XP14BRneo17; Normal transMRC5V1; GM0637; formed fibroblast lines; 1BR.3-G. Means ¡ standard error of 2–4 experiments.

MRC5V1, GM0637 and 1BR.3-G (Figure 5) suggested no increase in sensitivity in XP14BRneo17.

DNA Repair DSB were measured using PFGE. Immediately after irradiation there was a slight, but not significant, reduction in the level of induced damage in XP14BR in comparison with the control, GF11, over the dose range of 0–40 Gy. There was no evidence for a difference between the two cell lines after a 4 h period for repair (data not shown). When, however, a time course of repair was generated (Figure 6) DSB repair in XP14BR.2 was more rapid at earlier times than in control cells. A similar pattern has been reported for A-T cells. At later times, however, the residual levels of DSB were greater than in normal cells, but less than in A-T cells. As a second measure of repair, the time course of repair of chromosomal damage, as reflected in the reduction of excess chromosome fragments, revealed a small, but consistent, deficiency in XP14BR.2 (Figure 7).

Transfection with the XPC gene A series of single cell derived clones were established from the transfection of XP14BRneo17 with XPC cDNA. XPC protein was analysed by immunoblotting with an antibody to the C-terminus of the XPC protein. The XP14BR cell line has a mutation which truncates this protein [36], so that it lacks the C-terminus against which 514

Figure 6. DNA double strand break (DSB) repair rate after 30 Gy delivered at 4 ˚C. Repair data are expressed as the fraction of radiation-induced DNA fragments (FAR) migrating in the gel that remained at the indicated repair times. Each point represents the mean ¡ standard error of at least three replicate experiments. Key: XP14BR.2; 149BR, Normal non-transformed fibroblast line.

the antibody was raised. As expected, no XPC protein was detected in the XP14BRneo17 recipient cell line (Figure 8, lane 3). However, the expected 125 kDa band of XPC protein was restored to the level present in MRC5V1 (lane 4) in clones T921A (lane 1) and T943A (lane 2), but not in T941B or T948B (lanes 5 and 6). No protein band was visible in the T941B or T948B transfectants, using a low exposure of the X-ray film. Overexposure of the film revealed a very low level of Cterminal XPC protein in transfectant T941B, but not in T948B (Figure 8, lower panel, lane 5).

Cell killing in transformed and transfected cells In response to UVC, there was a significant enhancement of survival of T921A and T943A (the clones expressing substantial amounts of XPC protein) above XP14BRneo17. However, a similar level of restoration of UV resistance was also seen in clone T941B, which only expressed low levels of XPC protein (Figure 8). The UVC sensitivity recorded for these three clones was intermediate between XP14BRneo17 and GM0637 and MRC5V1 (Figure 9), the concurrent control normal cell lines. The response of XP14BR.2, XP14BRneo17, three normal cell lines, 1BR.3, 142BR and MRC5 and their immortalized derivatives; 1BR.3neo, 142BRneo and MRC5V1 to graded doses of gamma irradiation is illustrated in Figure 10a. The enhancement of survival The British Journal of Radiology, June 2006


Figure 7. Rate of repair of chromosome breaks after 6 Gy delivered at 4 ˚C. Repair data are expressed as percentage of premature chromosome condensation fragments in excess (ECF) at the repair times indicated. Each point represents the mean ¡ standard error of at least three replicate experiments. Key: XP14BR.2; 149BR, Normal nontransformed fibroblast line.

after ionizing radiation as a consequence of transformation was first described by Arlett et al [37]. This is particularly evident in XP14BR, with the consequence that the level of discrimination between it and the reference normals is less than in the corresponding untransformed cell lines. However, unlike the case with UVC, after exposure to gamma radiation no enhancement of survival above the

Figure 8. Immunoblotting for the C-terminal end of the XPC protein. Lanes 1 and 2, T921A, T943A (transfectants); lane 3, XP14BRneo17 (recipient); lane 4, MRC5V1 (control); lanes 5 and 6, T941B and T948B (transfectants). Upper panel, 20 s exposure, lower panel, 5 min exposure.


Figure 9. Ultraviolet radiation below 280 nm (UVC) sensitivity of XP14BRneo17 and its transfectants. Means ¡ standard error of the number of experiments indicated. Key: XP14BRneo17 (non-transfected, 1 expt); T921A (2 expts); T943A (4 expts); T941B (2 expts) (all transfectants); MRC5V1 (1 expt); GM0637 (2 expts).

parental cell line was observed. Indeed, all four clones appeared slightly more sensitive (Figure 10b).

Discussion Our patient was unusual amongst XPs in presenting with hypoplastic anaemia, and was originally thought to be a case of Fanconi’s anaemia [22]. Her angiosarcoma was also considered to be unique [23] in association with XP. The clinical radiosensitivity and consequent pathology which followed treatment of this scapular lesion is peculiar to this individual [21]. At the cellular level, the response to UVC is unexceptional. However, the hypersensitivity to gamma radiation has, in our hands, only been demonstrated in one previous instance using fibroblasts (XP3BR) from an individual from complementation group G [18]. The lack of clinical sensitivity in the two other representatives of complementation group C, XP23BE and XP233VI, which act as reference controls, was reflected in their normal responses in our cellular assays (Figure 4). The cellular sensitivity of XP14BR to ionizing radiation may be contrasted to the lack of sensitivity to cross-linking agents and implies a cellular defect in repair specific to gamma radiation. The study of repair of DSB in XP14BR generates results reminiscent of those obtained with A-T cells; an early, more rapid repair than in normals culminating at 24 h 515

C F Arlett, P N Plowman, P B Rogers et al Figure 10. Ionizing radiation sensitivity of XP14BR, XP14BRneo17, and transfectants. Means ¡ standard error of the number of experiments indicated. A. Effect of SV40 transformation on radiosensitivity. Key: XP14BR (6 expts); XP14BRneo17 (1 expt); 1BR.3 (58 expts); 1BR.3neo (4 expts); 142BR (8 expts); 142BRneo (3 expts); M R C 5 ( 1 e x p t ) ; MRC5V1 (2 expts). B. Comparison of transfectants with X P 1 4 B R n e o 1 7 . K e y : XP14BRneo17 (1 expt); T921A (2 expts); T943A T941B (2 expts); (3 expts); T948B (1 expt).

with less repair [32]. When combined with the outcome of the study of repair of chromosome fragments, these results are consistent with a defect in the repair of a small fraction of DSB as has been found for A-T cells [32, 38]. We have already acknowledged that the cellular radiosensitivity in XP14BR may be discriminated from that manifest in A-T cells by the normal level of radiation resistant DNA synthesis [21]. This alone is suggestive of the possibility of the presence of another non-A-T radiosensitivity gene, but first it was necessary to verify that the XPC mutation [36] itself was not responsible for the hypersensitivity. The correction, by transfection with the XPC gene, of the XP genetic defect was, as expected, successful in restoring cellular resistance to UVC irradiation. In one transfectant, T941B, substantial resistance to UVC irradiation was restored, even though only very low levels of XPC protein were detected in the transfectant. This suggests that XPC protein is not rate-limiting for neuclotide excision repair and that this low level is sufficient to deal with significant levels of UV damage. The psv3neo immortalization of XP14BR itself produced a large enhancement of ionizing radiation resistance. This effect of transformation, reported earlier [37], is specific to ionizing radiation and not UV, and may be generated by the abrogation of functional p53. However, the failure to enhance ionizing radiation resistance in clones, which, in addition to showing an enhancement of their UV response are also positive for the UVC gene product, supports the contention that XP14BR is a double mutant combining defects in both XPC and an unknown radiosensitivity gene. The resolution of this proposed genetic defect and the identification of the gene(s) becomes a priority. The other unusual clinical features exhibited by the patient remain unexplained. There was no evidence for more specific defects in immunity. The present demonstration of both clinical and cellular radiosensitivity in XP14BR raises the question of the frequency of such radiosensitivity in this syndrome. While XP14BR may be unique in being a double mutant, we have reported a previous instance of cellular sensitivity in XP3BR [18]. There was no relevant radiotherapy in this case. In this report we provide data 516

showing normal cellular responses for XP15BR, from complementation group A and XP23BE and XP233VI from complementation group C. The record of the outcome of radiotherapy in the patient who generated the cell line XP23BE is unambiguous in excluding any clinical sensitivity [19]. The clinical information of the outcome of combined chemotherapy and radiotherapy for the second XPC patient is not so clear cut since the child died 3 months after treatment [20]. In the absence of any cellular radiosensitivity there remains the possibility of sensitivity to components of the chemotherapy regimen. As far as we are aware, the patient from whom cell line XP15BR was generated has never come to radiotherapy. In the light of the outcome of radiotherapy in our present case and the observation of cellular radiosensitivity in XP3BR, it becomes important to resolve the extent of radiosensitivity in XP. This is reinforced by the possibility that these patients, because of their high frequency of tumours, are more likely to experience radiation.

Acknowledgments We are grateful to Randy Legerski and Peter van der Spek for the XPC cDNA plasmids, and to Peter van der Spek for the anti-XPC antibody. NF was supported by an INSERM Interface grant, Electricite´ de France, Ligue Nationale Contre le Cancer and ARC fundations. CFA, ARL and MHLG were supported in part by Euratom grant B16-E1042-UK.

References 1. Bootsma D, Kraemer KH, Cleaver JE, Hoeijmakers JHJ. Nucleotide excision repair syndromes: xeroderma pigmentosum, Cockayne Syndrome, and trichothiodystrophy. In: Volgelstein B, Kinzler KW, editors. The genetic basis of human cancer. New York, NY: McGraw-Hill, 1998:245–74. 2. Kraemer KH, Lee MM, Scotto J. Xeroderma pigmentosum. Cutaneous, ocular and neurologic abnormalities in 830 published cases. Arch Dermatol 1987;123:241–50. 3. Arlett CF, Lehmann AR. Xeroderma pigmentosum, Cockayne syndrome and trichothiodystrophy: sun sensitivity, DNA repair defects and skin cancer. In: Eeles R, Ponder B, Easton DF, Horwich EA, editors. Genetic predisposition to cancer, 2nd edn. London: Arnold, 2004:214–31.


Ionizing radiation hypersensitivity in xeroderma pigmentosum 4. Gartler SM. Inborn errors of metabolism at the cell culture level. In: Fishbein M, editor. Second International Conference on Congenital Malformations. New York, NY: International Medical Congress, 1964:94. 5. Cleaver JE. Deficiency in repair replication of DNA in xeroderma pigmentosum. Nature 1968;218:652–6. 6. Maher VM, McCormick JJ. Effect of DNA repair on the cytotoxicity and mutagenicity of UV irradiation and of chemical carcinogens in normal and xeroderma pigmentosum cells. In: Yuhas JM, Tennant RW, Regan JB, editors. Biology of radiation carcinogenesis. New York, NY: Raven Press, 1976:129–45. 7. Arlett CF, Harcourt SA. Variation in response to mutagens amongst normal and repair-defective human cells. In: Lawrence CW, editor. Induced mutagenesis. Molecular mechanisms and their implications for environmental protection. New York, NY: Plenum Press, 1983:249–66. 8. Maher VM, McCormick JJ, Grover P, Sims P. Effect of DNA repair on the cytotoxicity and mutagenicity of polycyclic hydrocarbon derivatives in normal and xeroderma pigmentosum human fibroblasts. Mutation Res 1977;43:117–38. 9. Lehmann AR, Kirk-Bell S, Arlett CF, Paterson MC, Lohman PHM, de Weerd-Kastelein EA, et al. Xeroderma pigmentosum cells with normal levels of excision repair have a defect in DNA synthesis after UV-irradiation. Proc Natl Acad Sci USA 1975;72:219–23. 10. Masutani C, Kusumoto R, Yamada A, Dohmae N, Yokoi M, Yuasa M, et al. The XPV (xeroderma pigmentosum variant) gene encodes human DNA polymerase eta. Nature 1999;399:700–4. 11. Arlett CF, Harcourt SA, Broughton BC. The influence of caffeine on cell survival in excision-proficient and excisiondeficient xeroderma pigmentosum and normal human cell strains following ultraviolet light irradiation. Mutation Res 1975;33:341–6. 12. Maher VM, Ouellette LM, Curren RD, McCormick JJ. Frequency of ultraviolet light-induced mutations is higher in xeroderma pigmentosum variant cells than in normal human cells. Nature 1976;261:593–5. 13. Myhr BC, Turnbull D, DiPaolo JA. Ultraviolet mutagenesis of normal and xeroderma pigmentosum variant human fibroblasts. Mutation Res 1979;62:341–53. 14. Ward JF. DNA damage and repair. In: Glass WA, Varma MN, editors. Physical and chemical mechanisms in molecular radiation biology. New York, NY: Plenum Press, 1991:403–21. 15. Goodhead DT. Initial events in the cellular effects of ionising radiations: clustered damage in DNA. Int J Radiat Biol 1994;65:7–17. 16. Tchou J, Kasai S, Shibutani M-H, Chung J, Laval J, Grollman AP, Nishimura S. 8-Oxoguanine (8-hydroxyguanine) DNA glycosylase and its substrate specificity. Proc Natl Acad Sci USA 1991;88:4690–4. 17. Ward JF. DNA damage produced by ionizing radiation in mammalian cells: identities, mechanisms of formation and repairability. Progress in Nucleic Acid Research 1988;35:95–125. 18. Arlett CF, Harcourt SA, Lehmann AR, Stevens S, FergusonSmith MA, Morley WN. Studies on a new case of xeroderma pigmentosum (XP3BR) from complementation group G with cellular sensitivity to ionizing radiation. Carcinogenesis 1980;1:745–51. 19. DiGiovanna JJ, Patronas N, Katz D, Abangan D, Kraemer KH. Xeroderma pigmentosum: spinal cord astrocytoma with 9 year survival after radiation and isotretinoin therapy. J Cutaneous Med Surg 1998;2:153–8. 20. Giglia G, Bouffet E, Jouvet A, Ohgaki H, Kleihues P, Sarasin A. Molecular analysis of glioma and skin-tumour alterations in a xeroderma-pigmentsum child. Int J Cancer 1999;81:345–50.


21. Rogers PB, Plowman PN, Harris SJ, Arlett CF. Four radiation hypersensitivity cases and their implications for clinical radiotherapy. Radiother Oncol 2000;57:143–54. 22. Salob SP, Webb DKH, Atherton DJ. A child with xeroderma pigmentosum and bone marrow failure. Br J Dermatol 1992;126:372–4. 23. Leake J, Sheehan MP, Rampling D, Ramani P, Atherton DJ. Angiosarcoma complicating xeroderma pigmentosum. Histopathology 1992;21:179–81. 24. Arlett CF, Harcourt SA. Survey of radiosensitivity in a variety of human cell strains. Cancer Res 1980;40:926–32. 25. Henderson L, Cole H, Arlett CF, James SE, Cole J, Lehmann AR, et al. Diagnosis of ataxia-telangiectasia by T-lymphocyte cloning assay. Lancet 1985;ii:1242. 26. Mayne LV, Priestley A, James MR, Burke JF. Efficient immortalization and morphological transformation of human fibroblasts by transfection with SV40 DNA linked to a dominant marker. Exptl Cell Res 1986;162:530–8. 27. Arlett CF, Harcourt SA, Cole J, Green MHL, Anstey AV. A comparison of the response of unstimulated and stimulated T-lymphocytes and fibroblasts from normal, xeroderma pigmentosum and trichothiodystrophy donors to the lethal action of UV-C. Mutation Res 1992;273:127–35. 28. Teo IA, Arlett CF. The response of a variety of human fibroblast cell strains to the lethal effects of alkylating agents. Carcinogenesis 1982;3:33–7. 29. Blo¨cher D, Einspenner M, Zajackowski J. CHEF electrophoresis, a sensitive technique for the determination of DNA double-strand breaks. Int J Radiat Biol 1989;56:437–48. 30. Whitaker SJ, McMillan TJ. Oxygen effects for DNA doublestrand breaks determined by pulse field gel electrophoresis. Int J Radiat Biol 1992;61:29–41. 31. Whitaker SJ, McMillan TJ. Pulsed-field gel electrophoresis in the measurement of DNA double-strand break repair in xrs-6 and CHO cell lines: DNA degradation under some conditions interferes with the assessment of double-strand break rejoining. Radiat Res 1992;130:389–92. 32. Foray N, Priestley A, Alsbeih G, Badie C, Capulas EP, Arlett CF, et al. Hypersensitivity of ataxia-telangiectasia fibroblasts to ionizing radiation is associated with a repair deficiency of DNA double-strand breaks. Int J Radiat Biol 1997;72:271–83. 33. Badie C, Iliakis N, Foray N, Alsbeih G, Cedervall B, Chavaudra N, et al. Induction and rejoining of DNA double strand breaks and interphase chromosome breaks after exposure to x-rays in one normal and two hypersensitive human fibroblast cell lines. Radiat Res 1995;144:26–35. 34. Legerski R, Peterson C. Expression cloning of a human DNA repair gene involved in xeroderma pigmentosum group C. Nature 1992;359:70–3. 35. van der Spek PJ, Eker A, Rademakers S, Visser C, Sugasawa K, Masutani C, et al. XPC and human homologs of RAD23: intracellular localization and relationship to other nucleotide excision repair complexes. Nucleic Acids Res 1996;24:2551–9. 36. Chavanne F, Broughton BC, Pietra D, Nardo T, Browitt A, Lehmann AR, et al. Mutations in the XPC gene in families with xeroderma pigmentosum and consequences at the cell, protein and transcript levels. Cancer Res 2000;60:1974–82. 37. Arlett CF, Green MHL, Priestley A, Harcourt SA, Mayne LV. Comparative human cellular radiosensitivity: I. The effect of SV40 immortalisation on the gamma-irradiation survival of skin derived fibroblasts from normal individuals and from ataxia-telangiectasia patients and heterozygotes. Int J Radiat Biol 1988;54:911–28. 38. Riballo E, Kuhne M, Rief N, Doherty AJ, Smith GCM, Recio M-J, et al. A pathway of double strand break rejoining dependent upon ATM, Artemis and proteins locating to cH2AX foci. Mol Cell 2004;16:715–24.

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SHORT COMMUNICATION

Survival and initial chromatid breakage in normal and tumour cells exposed in vitro to gamma rays and carbon ions at the HIRFL 1,2

Y JIANSHE, PhD, MSc, 1L WENJIAN, PhD, 1J XIAODONG, W WEI, PhD and 3G QINGXIANG, BSc

MSc,

1

J XIGANG,

PhD,

1

G CHUANLING,

BSc,

1

1

Institute of Modern Physics, Chinese Academy of Sciences, 2Life Science School of North West Normal University and 3Life Science School of Lanzhou University, Lanzhou, China

ABSTRACT. Human hepatoma and normal liver cells were irradiated with 12C6+ ion beams (linear energy transfer (LET)596 keV mm21) and c-rays at the Heavy Ion Research Facility in Lanzhou (HIRFL). The numbers and types of chromatid breaks were detected using the premature chromosome condensation technique. Irradiation with 12C6+ ions produced a majority of isochromatid break types, while chromatid breaks were dominant for irradiation with c-rays. Experimental results showed that the initial level of chromatid breaks is clearly related to the absorbed dose from 12C6+ ions and c-rays. The 12C6+ ions are relatively more effective at inducing initial chromatid breaks when compared with the c-rays. A relative biological effectiveness (RBE) of about 2.5 resulted for the induction of initial chromatid breaks by 12C6+ ions relative to c-rays in both cell lines.

Heavy charged particle beams have been applied in tumour radiotherapy because of their clear radiobiological response in the destruction of malignant cells [1]. The induction of chromatid breaks in the cell by radiation is considered to be a major cause of damage, which can lead to cell death. It has been reported by Kawata et al [2] that low linear energy transfer (LET) radiation induced mostly chromatid breaks, whereas the most isochromatid breaks were generated by high LET radiation. Previous studies [3] by the present authors have shown that the chromatid break repaired more easily than the isochromatid break. A major project involving tumour radiotherapy using heavy ions is currently being conducted at the Heavy Ion Research Facility in Lanzhou (HIRFL). In the present communication, the use of the premature chromosome condensation technique to measure the level of G2 initial chromatid breaks of human hepatoma cells and normal liver cells is reported. It is a well known fact that the risk of liver cancer is high in Asia, especially in China. For this reason, one normal and one tumour human liver cell line were selected for exposure to carbon ions generated by the HIRFL. The main aim was to investigate and quantify the relationship between the radiation dose and the level of initial chromatid breakage. Such information is needed for predicting the radiosensitivity of these two liver cell lines, and for evaluating the relative biological effectiveness (RBE) of carbon ions for inducing the chromatid breaks. The experiments provided necessary and valuable data on liver cell radiosensitivity prior to the start of clinical radiotherapy. It was found that the two cell lines have moderate radiosensitivity when exposed to c-rays [3]. 518

Received 9 June 2005 Revised 9 November 2005 Accepted 15 November 2005 DOI: 10.1259/bjr/99734289 ’ 2006 The British Institute of Radiology

Materials and methods Cell culture and irradiation The human hepatoma cell line SMMC-7721 and the normal liver cell line L02 (obtained from the Chinese Center for Type Culture Collection, CCTCC) with moderate radiosensitivity, were grown in RPMI-1640 medium supplemented with 10% fetal calf serum at 37 ˚C in 5% CO2; additionally, 0.25 U ml21 of insulin was added to the L02 culture medium. Exponentially growing cells were irradiated with either c-rays obtained from a 60Co source (dose rate 0.2 Gy min21) or 12C6+ ion beams generated by the HIRFL facility with a dose range from 0 to 8 Gy. The initial energy of the 12C6+ ions was 80.55 MeV u21, which was attenuated by the 13.58 mm Lucite (r51.2 g cm23) to 20 MeV u21 before it reached the cells. The LET was 96 keV mm21 when the carbon ions interacted with the cells located in the region of the Bragg peak and the LET was calculated by the Trim Program 92. Dosimetry was performed with an air ionizing chamber and the uniformity of the carbon ion beams was found to be 85%, as measured with CR39 track etch plastic.

Colony assay After exposure to radiation, cells were washed three times using PBS (pH57.0), then trypsinized and put into 5 ml culture medium, and the cell density counted by using the light microscope. At each dose point 200, 400, or 600 cells were plated into 35 mm diameter culture dishes. 5 ml RPMI-1640 medium supplemented with The British Journal of Radiology, June 2006

Short communication: Survival and initial chromatid breakage

10% fetal calf serum was added and cells were incubated at 37 ˚C in 5% CO2 until colonies formed (7–14 days). Two dishes were plated per dose point and the experiments were repeated three times. All the data are the mean¡ standard deviation. Data were fitted using Origin 7.0 (Original Lab, America); the data from the cells exposed to c-rays were fitted using the non-linear fit program, while the data from the cells exposed to carbon ions were fitted using the linear fit program. The linear-quadratic 2 formula (S~e aD bD ) was used for the non-linear fit, where a and b represent the probability of a single or double photon interaction with the chromatids.

Chromosome preparation Calyculin-A (BIOMOL America) was used as the PCC inducer, which was dissolved in 100% ethanol as a 1 mmol l21 stock solution. In order to achieve the chromatid breaks, 50 nmol l21 of Calyculin-A was added to cell cultures 5 min before irradiation. Cells were then incubated for a further 30 min at 37 ˚C in 5% CO2. Chromosome spread was harvested by swelling the cells in 75 mmol l21 KCl for 20 min at 37 ˚C and fixed with Carnoy’s fixation. A final wash and fixation in the same fixative agent was completed before dropping the cells onto a glass slide and drying them in hot humidity. The chromosome was stained with 5% Giemsa for 20 min. According to the standard criteria [4], more than 40 G2 phase cells were scored for each dose level investigated. Briefly, a chromatid discontinuity or misalignment of the distal part to the lesion, or a non-stained region longer than the chromatid width was considered to be a chromatid break. Isochromatid breaks were considered to be two breaks occurring at the same position on the two sister chromatids, i.e. a penetrated lesion through the two q arms or p arms of the chromosome was regarded to be an isochromatid break. The total chromatid breaks were calculated by summing the numbers of chromatid and isochromatid breaks. 20 non-irradiated cells were scored and the mean of these chromatid breaks was recorded. There were very few spontaneous chromatid breaks. The number of chromatid breaks observed in the irradiated cells was reduced by the mean number of chromatid breaks observed in the non-irradiated cells, and this number forms the experimental data from which the results were evaluated. A small number of chromatid exchanges were also noticed, but these were not considered to be significant and did not have a linear relationship with absorbed dose.

Results Survival fraction of two cell lines exposed to gamma rays and carbon ions Survival curves for SMMC7721 and L02 cells exposed to c-rays and carbon ions are shown in Figure 1. Several trends can be seen: (1) the two cell lines have different survival curves for c-rays and carbon ion irradiation; (2) the cells irradiated with c-rays have a linear quadratic survival curve, while it is almost linear for carbon ion The British Journal of Radiology, June 2006

Figure 1. Survival curve of SMMC7721 and L02 cells exposed to gamma rays and carbon ions. Survival fraction value is the mean¡standard deviation. The survival curve of cells exposed to gamma rays was linear-quadratic, and was almost linear when cells were exposed to carbon ions. The survival fraction of SMMC7721 cells and L02 cells are significantly different when exposed to either gamma rays or carbon ions.

irradiation; (3) the surviving fraction of L02 cells was higher than that of the SMMC7721 cells for both exposure scenarios. For the cells exposed to c-rays, the fit parameters a and b were 0.03 and 0.06 for SMMC7721 cells and 0.04 and 0.05 for L02 cells, respectively.

Initial chromatid and isochromatid breaks after irradiation of G2-PCC Figure 2 shows the chromatid and isochromatid break frequency of hepatoma and liver cell lines after irradiation with c-rays and 12C6+ ion beams at different absorbed dose. For both cell lines, the G2 chromatid breaks and the isochromatid breaks increase linearly with the absorbed dose. The number of chromatid breaks and isochromatid break frequency in the SMMC7721 cell line was greater than that of the L02 cell line for each dose. The yields of two types of chromatid breaks irradiated with 12C6+ ion beams were more than that of c-rays. For both cell lines, the number of chromatid breaks induced by the carbon ions were about 2.5 times higher relative to the c-ray exposure. Hence, the biological effectiveness of carbon ions for inducing the chromatid breaks is about 2.5 times higher than the c-rays.

The proportion of chromatid breaks in two cell lines exposed to c-rays and 12C6+ ions Figure 3 shows the proportion of two types of chromatid breaks of two cell lines exposed to c-rays and 12C6+ ion beams. It can be seen in Figure 3 that for 12C6+ and c-ray irradiation of cells, the dominant 519

Y Jianshe, L Wenjian, J Xiaodong et al

Figure 3. Proportion of two types of chromatid breaks. For cells irradiated with gamma rays, the number of chromatid breaks in both cell lines was much more than that of isochromatid breaks; for cells exposed to carbon ions, the isochromatid breaks were much more frequent than the chromatid breaks. They were significantly different.

Figure 2. Correlation between absorbed dose and chromatid breaks. In both (a) L02 and (b) SMMC7721 cell lines, more isochromatid than chromatid breaks were seen when cells were exposed to carbon ions, and fewer isochromatid than chromatid breaks seen when cells were exposed to gamma rays. Fewer breaks were induced in L02 cells than in SMMC7721 cells after either carbon ion or gamma ray exposure. All the data were the mean¡standard deviation.

breakage types were isochromatid and chromatid, respectively.

ions. In the present study, it was observed that the surviving fraction of L02 and SMMC7721 cells irradiated with carbon ions were lower than for cells exposed to crays. The survival curve for two cell types exposed to carbon ions was linear while it was linear quadratic for cray exposure. After irradiation, G2 chromatid breaks and isochromatid breaks increased linearly with the absorbed dose for both cell lines, the increased ratio of L02 and SMMC7721 cell lines were 2.5 and 3.5 for chromatid breaks and 4.5 and 7 for isochromatid breaks, which was consistent with data from Kawata et al [2]. The yields of the two types of chromosome breaks irradiated with 12 6+ C ion beams were more than that of c-rays. The relative biological effectiveness (RBE) of 12C6+ ion beams in inducing the chromatid breaks was found to be about 2.5. These results are in agreement with several previous studies [2, 5, 6, 7]. The main type of chromatid break was isochromatid for cells exposed to 12C6+ ion beams, while the chromatid breaks predominated in cells exposed to c-rays. Kawata et al [2] reported that chromatid breaks dominated after low LET irradiation, while isochromatid breaks dominated for high LET exposures, such as heavy ions, suggesting that most isochromatid breaks resulted from two separate breaks on sister chromatids induced by independent electron tracks. For low LET radiation types, not enough energy can be deposited to penetrate sister chromatids simultaneously, so most breaks were found to be chromatid in nature. For cells exposed to heavy ions, more isochromatid breaks were induced. This explains why heavy ions have the higher RBE, which is relevant for the destruction of tumour cells.

Discussion Heavy ions have the advantage of destroying the tumour cells very effectively. Kawata et al [2] reported that the cell survival fraction of human fibroblast AG1552 was very much lower after exposure to heavy 520

Conclusion Our results suggest that heavy ions generated by the HIRFL have a higher efficiency in killing tumour cells, The British Journal of Radiology, June 2006

Short communication: Survival and initial chromatid breakage

and a high RBE for the induction of the chromatid breaks. Also, the chromatid breaks are tightly correlated with the cell surviving fraction and radiosensitivity.

Acknowledgment This work was supported jointly by National Natural Science Foundation of China and National Basic Research Program of China through fellowship No. 10335050 and 2003CCB00200. We express our thanks to all the workers in HIRFL, for providing us with the high quality heavy ion beams.

References 1. Kraft G. Tumor therapy with heavy charged particles. Progress in Particle Nuclear Physics 2000;45:s473–544.


2. Kawata T, Durante M, Frusawa Y, et al. Dose-response of initial G2-chromatid breaks induced in normal human fibroblasts by heavy ions. Int J Radiat Biol 2001;77:165–74. 3. Yang JS, Li WJ, Zhou GM, et al. A comparative study on radiosensitivity of various tumor cells and human normal liver cells. World J Gastroenterol 2005;11:4098–101. 4. Savage JR. Classification and relationships of induced chromosomal structural changes. J Med Genet 1976;13: 103–22. 5. Kawata T, Gotoh E, Durante M, et al. High-LET radiationinduced aberrations in prematurely condensed G2 chromosome of human fibroblasts. Int J Radiat Biol 2000;76:929–37. 6. Kawata T, Durante M, Furusawa Y, et al. Rejoining of isochromatid breaks induced by heavy ions in G2-phase normal human fibroblasts. Radiat Res 2001;156:598–602. 7. Kawata T, Ito H, Uno T, et al. G2 chromatid damage and repair kinetics in normal human fibroblast cells exposed to low- or high-LET radiation. Cytogenet Genome Res 2004;104: 211–5.

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SHORT COMMUNICATION

Usefulness of mini-tracheostomy and torque controlled insertion of applicator in fractionated endobronchial brachytherapy 1

K KISHI, MD, PhD, 2T YOSHIMASU, MD, PhD, 1S SHIRAI, MD, 3Y MINAKATA, 1 1 1 PhD, T SONOMURA, MD, PhD, Y SHIOYAMA and M SATO, MD, PhD

MD, PhD,

1

M KIMURA,

MD,

Departments of 1Radiology, 2Thoracic and Cardiovascular Surgery and 3Third Department of Internal Medicine, Wakayama Medical University (WMU), Kimiidera 811-1, Wakayama City, 641-0012, Japan ABSTRACT. Endobronchial brachytherapy was developed as effective treatment of endobronchial cancer and fractionated schedule is applied to decrease late toxicity. However, repeated bronchofiberscopy is onerous to the patient and restricts the treatment schedule itself. We applied mini-tracheostomy for a ready access route, and a torque controlled technique for easy insertion of the endobronchial applicator. Eight patients with tracheobronchial cancer invasion were treated with endobronchial brachytherapy of 18–30 Gy/3–5 fractions/1.5–2.5 weeks (median 24 Gy/4 fractions/ 2 weeks) at reference points of 5 mm from the bronchial surface. The averaged individual irradiation and single session times were 4 min and 24 min, respectively. There were no procedure-related complications. These technical improvements may facilitate flexible fractionated dose prescriptions.

Fractionated endobronchial brachytherapy was developed for both effective palliation of endobronchial cancer invasion [1–3], and eradication of localized superficial endobronchial cancers [4, 5]. This was performed with an endobronchial applicator inserted into the target lumen by a transoral or transnasal bronchofiberscopic approach [1–3, 5–8]. A shortcoming of this technique is the need to repeat the transoral or transnasal endobronchial approaches. Irrespective of possible fractionation patterns to decrease late radiation toxicity [9], a commonly used fractionation schedule has been once weekly [2, 4, 5, 7, 8, 10, 11]. To reduce the burden of this therapy we introduced a ready access procedure through a mini-tracheostomy, created without cartilage damage and maintained using a recently devised fast and less invasive device composed of a cannula-introducer assembly, which was originally devised for emergency use [12] and has been applied for various purposes [13, 14]. Usually the ministoma spontaneously closes in a few days. Additionally, the style of the applicator was curved to enable a torque control manoeuvre to insert it smoothly under fluoroscopy. The clinical feasibility and usefulness of these procedures were evaluated clinically.

Received 18 May 2005 Revised 14 January 2005 Accepted 24 January 2006 DOI: 10.1259/bjr/31613651 ’ 2006 The British Institute of Radiology

this brachytherapy treatment (Table 1) in our department of radiation oncology, where 60–80 patients a year were treated (to #7). The cancer ingrowths were the only clinical manifestation at the time of the referral. The participants consisted of six males and two females aged 51–70 years (median 55.9 years). Karnofsky performance status of each patient was 80% or more. There were three tracheal cancers, three lung cancers, and one mediastinal cancer and metastatic cancer each. All had previously received external beam radiation therapy at doses of 50.4–70 Gy (median 60 Gy). Therapeutic goals of the brachytherapy included symptomatic relief from bronchial obstruction in four and prevention of further trachobronchial obstruction in all (Table 1). Patients were first evaluated by bronchofiberscopy to determine indications and target locations [11]. Informed consent was obtained prior to treatment. A choice of two treatment protocols had been offered to the patients: once a week 5 Gy brachytherapy for 4 weeks each with bronchofiberscopy according to many other reports [2, 4, 5, 7, 8, 10, 11], or the present protocol of twice a week less than 2.5 weeks after creation of a mini-stoma.

Methods Patients

Mini-tracheostomy

Between January 1999 and June 2002, eight patients with inoperable advanced cancer growing into the endobronchial space, which developed 2.5–103.6 (median 46.7) months after disease onset, were subjected to

Under local anaesthesia, according to the general insertion procedure [15], the cricothyroid membrane was penetrated with a needle-guidewire set of the MiniTrach II system (Mini-Trach II; Portex Ltd, Kent, UK),

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Short communication: Mini-tracheostomy for endotracheobronchial radiotherapy

(a)

(b)

(c) Figure 1. (a) A cannula of Mini-Trach II was inserted to maintain a mini-stoma. (b) At fractionated endotracheobronchial radiotherapy, the cannula was temporarily withdrawn and the endobronchial or endotracheal mini-applicator was gently inserted through the mini-stoma (arrow) using a curved stylet (long arrow) under fluoroscopic monitoring. The soft elastic wings were stretched during insertion (shadowed arch indicated with a broken arrow). (c) Measuring wire with dots at 1 cm intervals was inserted in the bronchial applicator, the wings of which were opened to support the applicator position (Note the different opened width of the soft wings). Reference points were determined according to the vertical distance from the bronchial surface (one reference point is shown by an asterisk, the depth: vertical line).

and then a cannula-introducer assembly was introduced into the tracheal lumen with Seldinger’s procedure. Thus, a mini-stoma was created and then maintained with the cannula (Figure 1a).

Applicator deployment Patients were pre-medicated with inhalation of 0.2 g lidocaine via a nebulizer and intramuscular injection of 0.5 mg of atropine sulphate and 25 mg of hydroxyzine. An endobronchial mini-applicator (Clinical Supply, Tokyo, Japan) composed of a stylet, an elastic hard polyethylene inner tube and a soft silicone outer tube with self-expandable wings stretched by the inner tube (Figure 1b) [5, 7, 8], was introduced into the trachea via the stoma and inserted to the target lumen under fluoroscopic guidance, paying attention not to stimulate the tracheobronchical wall. The inner tube was curved to enable torque manipulation to reach the lesion more easily (Figure 1c) [16]. Then the inner tube was withdrawn to allow the wings to open softly and hold the The British Journal of Radiology, June 2006

catheter position in the centre of the lumen. Routine lidocaine spray topical anaesthesia, which is an irritant, was not necessary during the procedure because of the smooth insertion. Before each session, the cannula was temporarily withdrawn, and after each session the cannula was re-inserted to maintain the stoma until the final session, after which the stoma was left for spontaneous obliteration.

Dose prescription Treatment was scheduled twice weekly to a total dose of 18–30 Gy (median 24 Gy) in 3–6 fractions (median 4 fractions) with a single fraction dose of 5–6 Gy (median 6 Gy) for 1.5–2.5 weeks. These prescriptions were determined based on the necessity of treatment and previously irradiated dose at the reference point of the brachytherapy. In case no. 6, it was vital to suppress the widespread intratracheal tumour relapse, even though the total dose came to surpass the tolerance level of the 523

K Kishi, T Yoshimasu, S Shirai et al

normal tissue. The reference point was at 5 mm depth from the bronchial (tumour) surface (Figure 1c). The mean volume with the prescribed dose was 11 cm3, ranging from 4.3 cm3 to 18.3 cm3. The biological effective doses (BED) of the present schedules were the same as those corresponding to the once weekly schedules both at a/b ratio 3 and at 10.

Follow-up Patients were followed up at our out-patient clinic. Bronchofiberscopic follow-up study and X-ray CT study were scheduled for the third month, and every sixth month thereafter.

in four fractions at the outpatient clinic. The dyspnoea disappeared in 2 weeks. She survived 43.8 months after the brachytherapy. A 57-year-old-woman (patient no. 8) suffered from dyspnoea and obstinate cough due to relapsed tracheal cancer. 8 years previously she had undergone noncurative surgical resection of the tumour followed by 70 Gy in 35 fractions of external beam radiotherapy. During the recent 2 years, two metastatic nodules, both in the left lung, were excised and one nodule in the right lung was treated with stereotactic radiotherapy. A laser treatment and stent placement were attempted to relieve the present symptoms but were ineffective. The previous dose in the relapsed lesion was 40 Gy. An endobronchial irradiation of 24 Gy in four fractions was planned. She continues to maintain a good performance status for over 30.4 months after the brachytherapy.

Results Mini-tracheostomy was successfully performed and maintained in all patients. In the fractionated endobronchial treatment, the mean time from insertion to the targeted location of the applicator under fluoroscopic guidance was 8 s (range 6–18 s). The mean dwelling time of source in the irradiating position and the averaged single session time of each fractionated brachytherapy were 3.1 min (range 1 min 57 s to 8 min 6 s) and 24 min (range 14–35 min), respectively, excluding the time for treatment planning using the PLATO system. There were no procedure-related complications. The stomas closed in 3 days. All patients obtained symptomatic improvement, and four patients showed histologically negative bronchofiberscopic biopsy specimens. During the mean follow-up period of 18.9 months, no patient developed local relapse on the irradiated lesion. Three patients died of metastatic disease and two patients died of aggravated respiratory distress due to disease progression. One patient died of pneumonia due to tracheal fistula. Another died of haematoemesis at home. One patient is alive at 31 months presently without disease.

Case presentations A 53-year-old-man (patient no. 4) complaining of severe dyspnoea due to obstruction of the trachea and bilateral main bronchi was admitted to our hospital. He underwent 60 Gy in 30 fractions of external beam radiotherapy combined with intravenous administration of cisplatinum and doxorubicin. To the persisting tumour, an endobronchial brachytherapy of 24 Gy in four fractions was scheduled. 1 month after the brachytherapy, no tumour or viable cells were found in the trachea. He was asymptomatic for 12 months. A 52-year-old-woman (patient no. 7) complaining of mild dyspnoea due to metastatic left bronchial tumour obstruction was indicated for endobronchial brachytherapy. Since undergoing low anterior resection for rectal cancer 6.7 years previously, she had experienced gamma-knife therapy for brain metastases, conformal radiotherapy for pelvic relapse and right mediastinum mass, and stereotactic body irradiation for lung metastases. She underwent endobronchial irradiation of 24 Gy 524

Discussion The safety and less invasive nature of the minitracheostomy procedure have been well established [12]. The insertion procedure was fast and simple, and the reported mean procedural time was only 6.7 s from incision to removal of the guidewire [17]. During the endobronchial treatment, the mini-stoma was safely used as a ready access route. Through the mini-stoma, the applicator was inserted quickly and smoothly under fluoroscopy to the target position using a torque control technique. The patients were freed from the repeated burden of transoral or transnasal endobronchial manipulation under bronchofiberscopy. Thus, this fast fluoroscopic procedure would save time, cost and effort of the patients. Furthermore, the fast and convenient fluoroscopic procedure allows flexible schedules of fractionation unrestricted by bronchofiberscopy availability. Despite biological knowledge that fractionation decreases the late radiation toxicity [9], most of the reported schedules have been restricted to once a week: for instance, 30 Gy in two fractions [18], 30 Gy in 6 fractions [19], 22.5 Gy in three or 10 Gy in two [11], 20 Gy in five [10], 10 Gy in one or 18 Gy in three [8] and 14 Gy in 2 weeks as a boost. To our knowledge, only one schedule has been described in which 25 Gy for weekly five fractions was given [20]. Another merit of this procedure was that the applicator insertion to deployment was carried out under smooth fluoroscopic guidance, and anaesthesia such as local lidocaine spray, which is irritating to the airway, was not necessary. It is difficult to discuss the precise effects, including late toxicity, because of the limited number of observations in the present study and dosimetric inconsistency between reports due to different depth of reference points [1, 5]. However, efforts to control late radiation toxicity should be continued. Some patients survived for relatively long periods, suggesting that an appropriate selection of patients might result in a good outcome of the endotracheal brachytherapy in some cases, but at present any such conclusion is unwarranted due to the small number of cases studied. The possible benefits of well-fractionated endobronchial brachytherapy need to be determined in future investigations. The British Journal of Radiology, June 2006

Short communication: Mini-tracheostomy for endotracheobronchial radiotherapy

References 1. Stout R, et al. Clinical and quality of life outcomes in the first United Kingdom randomized trial of endobronchial brachytherapy (intraluminal radiotherapy) vs. external beam radiotherapy in the palliative treatment of inoperable non-small cell lung cancer. Radiother Oncol 2000;56:323–7. 2. Hennequin C, et al. [Endobronchial brachytherapy: technique and indications]. Cancer Radiother 2003;7:33–41. 3. Ikeda H. [Developments in brachytherapy]. Gan To Kagaku Ryoho 1995;22:1317–23. 4. Perol M, et al. Curative irradiation of limited endobronchial carcinomas with high-dose rate brachytherapy. Results of a pilot study. Chest 1997;111:1417–23. 5. Fuwa N, et al. The treatment results of 40 patients with localized endobronchial cancer with external beam irradiation and intraluminal irradiation using low dose rate (192)Ir thin wires with a new catheter. Radiother Oncol 2000;56:189–95. 6. Marsh BR. Bronchoscopic brachytherapy. Laryngoscope 1989;99(7 Pt 2 Suppl. 47):1–13. 7. Nomoto Y, et al. [Endobronchial brachytherapy with high dose rate 192Ir afterloading technique using a new applicator]. Nippon Igaku Hoshasen Gakkai Zasshi 1996;56:42–7. 8. Nomoto Y, et al. High dose rate endobronchial brachytherapy using a new applicator. Radiother Oncol 1997;45:33–7. 9. Nag S, et al. Consensus guidelines for high dose rate remote brachytherapy in cervical, endometrial, and endobronchial tumors. Clinical Research Committee, American Endocurietherapy Society. Int J Radiat Oncol Biol Phys 1993;27:1241–4.


10. Freitag L, et al. Sequential photodynamic therapy (PDT) and high dose brachytherapy for endobronchial tumour control in patients with limited bronchogenic carcinoma. Thorax 2004;59:790–3. 11. Celebioglu B, et al. High dose rate endobronchial brachytherapy effectively palliates symptoms due to inoperable lung cancer. Jpn J Clin Oncol 2002;32:443–8. 12. Cushing M. Minitrach in airway obstruction. Anaesthesia 1986;41:774. 13. Daykin AP. A new use of a ‘Minitrach’ introducer. Anaesthesia 1993;48:538. 14. Bonde P, et al. Sputum retention after lung operation: prospective, randomized trial shows superiority of prophylactic minitracheostomy in high-risk patients. Ann Thorac Surg 2002;74:196–202; discussion 202–3. 15. Corke C, Cranswick P. A Seldinger technique for minitracheostomy insertion. Anaesth Intensive Care 1988;16:206–7. 16. Kishi K, et al. Treatment of malignant tracheobronchial stenosis by Dacron mesh-covered Z-stents. Cardiovasc Intervent Radiol 1994;17:33–5. 17. Slots P, et al. Retrograde intubation with a Mini-Trach II kit. Acta Anaesthesiol Scand 2003;47:274–7. 18. Delclos ME, et al. Endobronchial brachytherapy with highdose-rate remote afterloading for recurrent endobronchial lesions. Radiology 1996;201:279–82. 19. Lorchel F, et al. [High dose rate brachytherapy: a potentially curative treatment for small invasive T1N0 endobronchial carcinoma and carcinoma in situ]. Rev Mal Respir 2003;20:515–20. 20. Yokoyama A, et al. Endobronchial brachytherapy for centrally located early-stage lung cancer. Haigan 1997;16:55–60.

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CASE REPORT

MR findings of penile lymphoma K-H CHIANG, MD, P-Y CHANG, and A S-B CHOU, MD

MD,

S-K LEE,

MD,

P-S YEN,

MD,

C-M LING,

MD,

C-C LIN,

MD,

C-C LEE,

MD

Department of Radiology, Buddhist Tzu Chi General Hospital, Hualien, Taiwan; ABSTRACT. Penile lymphoma is extremely rare and secondary involvement of the penis by lymphoma may be due to retrograde spread or to direct extension from neighbouring organ. The appearance of penile lymphoma varies and can be mistaken for other soft tissue tumours. We report on a case with malignant lymphoma of the penis. MRI findings revealed soft-tissue mass of homogeneous isointensity around the middle to distal part of penis on T1 weighted imaging and T2 weighted imaging. It was well encapsulated, minimally enhanced and distinct from corpus cavernosum and corpus spongiosum.

Penile lymphoma is an extremely rare neoplasm. Nodules, ulcers and diffuse penile swelling have been reported as the presentation of penile lymphoma, and the penile shaft is the most common site of involvement [1–4]. Full physical examination and radiological image investigations, including CT, MRI and PET (positron emission tomography), should be undertaken to stage the patient. Chemotherapy has the advantage of treating both the primary lesion and any systemic disease. Herein, we report on a case with malignant lymphoma of the penis and its imaging findings.

Case report A 77-year-old man presented with a more than 2-week history of a painless mass at penile base. He denied sexual exposure, previous sexually transmitted diseases, recent trauma and underlying medical disease. No difficulty with urination had been experienced. There was no fever, night sweats, fatigue or body weight loss in the previous 6 months. Physical examination revealed a firm nodule at penile base. The remainder of his physical examination was uneventful. Haematological and chemical panel values were within normal limits except elevated prostate specific antigen (PSA) and lactic acid dehydrogenase (LDH) (PSA: 23.0 ng ml21, normal range: 0–4 ng ml21; LDH: 751 IU l21, normal range: 15–400 IU l21). Pelvic MRI without and with Gd-DTPA enhancement was performed by a 1.5-Tesla MR scanner (Signa Excite; GE Medical System, Milwaukee, WI). The study showed soft-tissue mass of homogeneous isointensity around the middle to basal part of penis on T1 weighted and T2 weighted images. It was well capsulated, minimally enhanced by contrast medium, and distinct from both Address correspondence to: Dr Andy Shau-Bin Chou, Department of Radiology, Buddhist Tzu Chi General Hospital, 707, Section 3, Chung Yang Road, Hualien, Taiwan.

526

Received 8 March 2005 Revised 9 May 2005 Accepted 23 May 2005 DOI: 10.1259/bjr/55555394 ’ 2006 The British Institute of Radiology

the corpus cavernosum and corpus spongiosum (Figure 1). Cystoscopy with transrectal needle biopsy of prostate and scrotal incisional biopsy revealed a picture of large B-cell lymphoma, but there was no obvious abnormal finding of prostate while reviewing the MR images. CT of the chest, abdomen and pelvis showed only the penile mass (Figure 2) and bone marrow biopsy was negative. He underwent six courses of systemic chemotherapy with a CHOP regimen and tolerated the whole course of chemotherapy well. The penile mass disappeared gradually after six cycles of chemotherapy. At regular follow-up evaluations, the patient is free from recurrence and dissemination 8 months after the diagnosis.

Discussion Penile lymphoma is extremely rare and secondary involvement of the penis by lymphoma may be due to retrograde spread (haematic or lymphatic) or to direct extension from neighbouring organ [1–3, 5]. The most commonly affected site was the shaft, followed by the glans penis [6]. Diffuse large cell lymphoma was the most common histological subtype [7]. The appearance of penile lymphoma varies: it may appear as a mass, as plaques or ulcers in the skin of the organ, or as diffuse penile swelling [8]. Full physical examination and radiological image investigation, including CT, MRI and PET, should be undertaken to stage the patient [4]. Excision biopsy is essential to achieve the correct diagnosis, and histological analysis must include immunohistochemical tests to differentiate lymphoma from undifferentiated sarcomas or carcinomas and to distinguish between B- and T-cell lymphomas [7]. As in our case, MRI provides better tissue-contrast than CT. The margins between the mass and corpus cavernosum and corpus spongiosum are clear. MRI images of penile lymphoma had not been documented. It The British Journal of Radiology, June 2006

Case report: Lymphoma of the penis

(a)

(b)

(c)

(d)

Figure 1. (a) Axial T1 weighted image (repetition time (TR)/echo time (TE) 5 467/8) showed soft-tissue mass of homogeneous isointensity around the middle to basal part of penis. (b) Sagittal T2 weighted image (TR/TE 5 2350/87) showed the extension of the lesion. The capsule of corpus spongiosum is intact. (c) Coronal T1 weighted image (TR/TE 5 2650/87) showed the relationship between the mass and the corpus cavernosum and corpus spongiosum. (d) Coronal T1 weighted image after Gd-DTPA demonstrated the mass without obvious enhancement.


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K-H Chiang, P-Y Chang, S-K Lee et al

In conclusion, penile lymphoma is extremely rare and can be mistaken for other soft tissue tumours. The possibility of this diagnosis should be considered when evaluating a nodule or mass in the penis. MRI may play a role in the pre-treatment evaluation of penile lymphoma for local tumour involvement.

References

Figure 2. Contrast-enhanced CT showed soft-tissue mass of homogeneous isodense around the basal part of penis. There is also no obvious enhancement of the mass.

should be homogeneous isointensity on T1 weighted and T2 weighted images and minimally enhanced by contrast medium. Lymphoma of the penis seems to behave in a less aggressive pattern [9]. Chemotherapy has the clear advantage of obtaining good cosmetic and functional results; it will also be effective in patients with occult disseminated lymphomas [10]. Radical surgery should be used only after the failure of other modalities [7].

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1. Lo HC, Yu DS, Lee CT, Chen A, Chang SY, Sun GH. Primary B cell lymphoma of the penis: successful treatment with organ preservation. Arch Androl 2003;49:467–70. 2. Wang HT, Lo YS, Huang JK. Primary lymphoma of the penis. J Chin Med Assoc 2003;66:379–381. 3. Pomara G, Cuttano MG, Tripodo C, Carlino F, Selli C. Primary T-cell rich B-cell lymphoma of the penis: a first case. BJU Int 2003;91:889. 4. Israel O, Keidar Z, Bar-Shalom R. Positron emission tomography in the evaluation of lymphoma. Semin Nucl Med 2004;34:166–79. 5. Nakayama F, Sheth S, Caskey CI, Hamper UM. Penile metastasis from prostate cancer: diagnosis with sonography. J Ultrasound Med 1997;16:751–3. 6. Gough J. Primary reticulum cell sarcoma of the penis. Br J Urol 1970;42:336–9. 7. el-Sharkawi A, Murphy J. Primary penile lymphoma: the case for combined modality therapy. Clin Oncol 1996;8:334–5. 8. Bunesch Villalba L, Bargallo Castello X, Vilana Puig R, Burrel Samaranch M, Bru Saumell C. Lymphoma of the penis: sonographic findings. J Ultrasound Med 2001;20:929–31. 9. Moreno Aviles J, Salinas Sanchez AS, Gomez Gomez G, Server Falgas [Tumor of the penis: primary clinical manifestation of a lymphoma]. Actas Urol Esp 1988;12: 488–90. 10. Arena F, di Stefano C, Peracchia G, Barbieri A, Cortellini P. Primary lymphoma of the penis: diagnosis and treatment. Eur Urol 2001;39:232–5.



CASE REPORT

Deep tracheal laceration after balloon dilation for benign tracheobronchial stenosis: case reports of two patients 1

Y H KIM, MD, 1D J SUNG, J W UM, MD

MD,

1

S B CHO,

MD,

1

K B CHUNG,

MD,

1

S H CHA,

MD,

2

H S PARK,

MD

and

3

Departments of 1Radiology, 2Urology and 3Surgery, Korea University College of Medicine, Seoul, Korea ABSTRACT. We report two cases of deep tracheal laceration in female patients after balloon dilation for benign tracheobronchial stenosis. Immediate post-procedure bronchoscopy and CT including 3D reconstructions showed deep lacerations in the posterior tracheal wall. Clinically, the patients’ dyspnoea subsided and there has been no recurrence during follow-up after balloon dilation. On the follow-up 3D-reconstructed CT scans obtained 2 months and 8 months following balloon dilation, respectively, the lacerations had healed completely and there was considerable improvement in lumen size.

Since Cohen et al [1] initially reported balloon dilation of tracheobronchial stenosis in 1984, endoscopically or fluoroscopically guided balloon dilation has become an accepted treatment for benign tracheobronchial stenosis [2–7]. The major anticipated advantages of balloon dilation are lower morbidity and mortality than those of corrective surgery or bougienage. Although balloon dilation is safe and effective, excessive balloon inflation may theoretically lacerate or rupture the airway, thereby causing bleeding, pneumothorax, pneumomediastinum or mediastinitis. As far as we know, there is only one report [7] briefly describing superficial or deep lacerations of the tracheobronchial tree after balloon dilation for benign tracheobronchial strictures. However, there was neither a detailed description of the laceration nor information regarding the patients’ detailed clinical outcomes. We present two cases of deep tracheal laceration confirmed on post-procedure bronchoscopy and CT scans following balloon dilation for benign tracheobronchial stenosis, as well as the detailed clinical data.

Case reports Case 1 A 31-year-old female complained of aggravated cough, dyspnoea, and a 3 kg weight loss over a 7 month period. She had a past history of pulmonary tuberculosis 10 years ago. Chest radiography was normal. However, Address correspondence to: Ji Hoon Shin, MD, Department of Radiology, Asan Medical Center, University of Ulsan College of Medicine, 388-1, Pungnap-2dong, Songpa-gu, Seoul 138-736, Korea. This investigation has disclosed a potential conflict; none of the other authors have identified a conflict of interest.


Received 17 February 2005 Revised 6 May 2005 Accepted 23 May 2005 DOI: 10.1259/bjr/17839516 ’ 2006 The British Institute of Radiology

bronchoscopy with biopsy and chest CT including three-dimensional (3D) reconstructions (Figure 1) showed a 6 cm long tracheal fibrotic stenosis and a 1.5 cm long right main bronchial fibrotic stenosis, which was covered with whitish-yellow elevated lesions. The diameters of the narrowed segments of the trachea and right main bronchus were 6 mm and 2 mm, respectively, while those of the normal segment of the trachea and right main bronchus were 16 mm and 11 mm, respectively. A pulmonary function test (PFT) showed a forced expiratory volume in one second (FEV1) of 1.3 l (45%, predicted) and a forced vital capacity (FVC) of 2.0 l (54%, predicted). The detailed technique of balloon dilation is the same as described in the previous report [7]. We initially dilated the right main bronchial stenosis and subsequently the tracheal stenosis. A 6 mm diameter balloon catheter was used first for the severe right main bronchial stenosis in order to provide passage of the larger balloon catheter. Then the balloon (Boston Scientific/Medi-tech, Watertown, MA; 10 mm in diameter and 4 cm long for the bronchial stenosis, 18 mm in diameter and 10 cm long for the tracheal stenosis) was slowly inflated manually using a diluted water-soluble contrast medium until the waist formation in the inflated balloon catheter at the stricture segment disappeared. There was not much resistance during balloon inflation. Immediately after balloon dilation, the patient complained of mild chest pain and expectorated some bloodtinged sputum. On bronchoscopy (not shown) obtained immediately after balloon dilation, a tracheal laceration was observed at the posterior tracheal wall. Its length and depth were estimated at approximately 5 cm and 8 mm, respectively, on bronchoscopy. For further evaluation of the extent of the laceration and possible associated complications, CT scanning was performed and revealed a deep, longitudinal laceration on the 529

J H Kim, J H Shin, T S Shim et al

(a)

(b)

Figure 1. Case 1. (a) Anteroposterior and (b) lateral views of the three-dimensional (3D) reconstruction CT, obtained 4 days before balloon dilation, show a 6 cm long tracheal stenosis (arrowheads in (a) and (b)) from the mid-trachea to the carina and a 1.5 cm long right main bronchial stenosis (arrows in (a)).

posterior tracheal wall with the pneumomediastinum (Figure 2). As a result of the tracheal laceration and separation of the tracheal wall at the tear point, the diameter of the trachea became widened. The patient’s vital signs were stable, she was without fever or chills and there was no change in her haemoglobin level. Her chest pain and blood-tinged sputum disappeared within 24 h. Therefore, she was prescribed only oral antibiotics to prevent possible infection and was discharged from the hospital 5 days after the procedure. On follow-up CT scans (Figure 3) obtained 8 months following balloon dilation, the deep laceration had completely healed and the widened tracheal lumen was maintained. Tests, also obtained 8 months following balloon dilation, showed an FEV1 of 1.8 l (63%, predicted) and an FVC of 3.3 l (89%, predicted); the FEV1 and FVC increased as much as 18% and 35%, respectively, compared with those of the predilation PFT. The patient has maintained her symptomatic improvement without recurrence for 10 months.

Case 2 A 27-year-old woman had complained of respiratory difficulty for 3 months. She had a past history of tracheal intubation due to decreased mental ability by reason of 530

diabetic ketoacidosis for 6 months. Bronchoscopy and CT (Figure 4) obtained 3 weeks before balloon dilation revealed two focal fibrotic stenoses in the mid and lower levels of the trachea. The diameters of the narrowed segments in the mid and lower levels of the trachea were 5 mm and 12 mm, respectively, while the diameter of the normal segment in the trachea was 16 mm. PFT’s performed 1 week before balloon dilation were as follows: the FEV1 was 1.1 l (39%, predicted) and the FVC was 2.3 l (65%, predicted). Dilation was performed without much resistance, using an 8 cm long and 16 mm diameter balloon for the two focal tracheal stenoses. Immediately after balloon dilation, the patient complained of mild chest pain. On bronchoscopy and CT (Figure 5) obtained immediately after balloon dilation, a deep longitudinal laceration (4.5 cm in length, 1 cm in depth) was detected at the mid-level of the posterior tracheal wall. Her vital signs were stable without fever, and her chest pain disappeared within 24 h. She did not receive any further therapy except for preventative oral antibiotics. On follow-up CT scans (Figure 6) obtained 2 months following balloon dilation, the deep laceration had nearly disappeared and the tracheal lumen was widened. Follow-up PFT, obtained 2 months after balloon dilation, showed an FEV1 of 1.9 l (69%, predicted) and an FVC of 2.3 l (65%, predicted); the The British Journal of Radiology, June 2006

Case report: Deep tracheal laceration after balloon dilatation

(a)

(c)


(b)

Figure 2. Case 1. Immediately after balloon dilation, the lumens of the stenoses appear to be widened on (a) the anteroposterior view of the 3D reconstruction CT. However, a long and deep laceration (arrowheads in (b) and (c)) on the posterior tracheal wall with pneumomediastinum (arrows in (c)) is well visualized on (b) the lateral view of the 3D reconstruction CT and (c) axial CT scan.

531


(a)

(b)

Figure 3. Case 1. (a) Anteroposterior and (b) lateral views of the follow-up 3D reconstruction CT scans show marked improvement of the right main bronchial (arrows in (a)) and tracheal stenosis (arrowheads in (a)) without further visualization of the deep laceration on the posterior tracheal wall (arrowheads in (b)).

FEV1 had increased as much as 30% compared with that of the pre-dilation PFT. The patient’s symptoms were much improved and had not recurred for 5 months at the time of writing after balloon dilation.

Discussion Although expandable metallic stent deployment has been an effective and minimally invasive procedure for benign tracheobronchial stenosis and may overcome the problem of short-term recurrence after balloon dilation, stent placement has notable drawbacks, including migration, recurrence of stenosis from tissue hyperplasia, stent fracture and difficulty of stent removal [4, 7– 10]. Therefore, balloon dilation is usually performed as the initial treatment for benign tracheobronchial stenosis because it is a safe and simple procedure [2–7]. However, several complications, such as bronchospasm or lung atelectasis after balloon dilation, have been reported by several authors [7, 11–13]. Furthermore, overdilation of the stenosis may cause tracheobronchial rupture similar to the rupture or bleeding reported for balloon dilation of the upper gastrointestinal tract [14–16]. Although the predominant fibrotic process can be successfully dilated and has a potentially successful clinical outcome [3, 6], there may be potential complications such as laceration 532

or even free perforation of the tracheobronchial tree after sudden balloon dilation of fibrotic stenosis because of the hardness or stiffness of the fibrotic stenosis. To the best of our knowledge, Lee et al [7] initially reported tracheal or bronchial laceration after balloon dilation, i.e. two cases of deep mucosal laceration and 15 cases of superficial mucosal laceration of the 59 patients treated with balloon dilation for benign tracheobronchial stenosis. In their report, most patients (45 of 59, 76%) had chronic fibrotic stenosis due to tuberculosis, and balloons for bronchial and tracheal stenosis were 6–12 mm and 14– 20 mm in diameter, respectively. Their patients experienced tracheobronchial laceration (25%, 17/59) with relatively high frequency after balloon dilation, although the procedures were performed with caution. However, in Lee’s report, the clinical outcome and follow-up data of the deep tracheal or bronchial laceration were not documented in detail. Interestingly, in our two cases the long and deep tracheal laceration disappeared completely during the follow-up period after balloon dilation. Lee et al [7] also briefly stated that two deep lacerations left no subsequent clinical sequelae. We assume that, if a deep laceration heals soon without much growth of granulation tissue, a good clinical outcome can be achieved after the deep tracheobronchial laceration because the lumen of the stenotic segment will be widened enough. Balloon dilation for congenital tracheal stenosis is another illustration of The British Journal of Radiology, June 2006

Case report: Deep tracheal laceration after balloon dilatation

Figure 4. Case 2. Anteroposterior view of the three-dimensional (3D) reconstruction CT obtained 3 weeks before balloon dilation, shows two focal stenoses (arrows) at the mid and lower levels of the trachea.


Figure 5. Case 2. Immediately after balloon dilation, a deep laceration (arrowheads) on the posterior tracheal wall is clearly visualized on the axial CT scan.

533


Figure 6. Case 2. (a) Anteroposterior view of 3D reconstruction and (b) axial scan of the follow-up CT show improvement of the two focal tracheal stenoses (arrows in (a)) and reveal a completely healed deep laceration on the posterior tracheal wall.

further proof to support our assumption that rupture of the complete cartilaginous rings by balloon dilation represents a prerequisite step for increasing the luminal diameter itself [17, 18]. As in our cases, even the deep and large tracheal laceration can be treated conservatively if patients show stable conditions and a minimal and asymptomatic pneumomediastinum. However, if patients present with acute respiratory distress secondary to a tracheal laceration, surgical or interventional treatment is indispensable [19]. Bronchoscopy can detect early tracheal laceration and determine the location. However, bronchoscopy may not offer sufficient information about anatomical location and morphology of deep tracheal laceration due to limited sight of view. In contrast, CT including 3D reconstructions can not only delineate the precise extent and morphology of the deep laceration, but also detect pneumomediastinum or mediastinal bleeding, which are important factors in deciding the treatment plan.

References 1. Cohen MD, Weber TR, Rao CC. Balloon dilatation of tracheal and bronchial stenosis. AJR Am J Roentgenol 1984;142:477–8.

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2. Carlin BW, Harell JH, Moser KM. The treatment of endobronchial stenosis using balloon catheter dilatation. Chest 1988;93:1148–51. 3. Ferretti G, Jouvan FB, Thony F, Pison C, Coulomb M. Benign noninflammatory bronchial stenosis: treatment with balloon dilation. Radiology 1995;196:831–4. 4. Lee KW, Im JG, Han JK, Kim TK, Park JH, Yeon KM. Tuberculous stenosis of the left main bronchus: results of treatment with balloons and metallic stents. J Vasc Interv Radiol 1999;10:352–8. 5. Carre P, Rousseau H, Lombart L, et al. Balloon dilatation and self-expanding metal Wallstent insertion for management of bronchostenosis following lung transplantation. Chest 1994;105:343–8. 6. Sheski FD, Mathur PN. Long-term results of fiberoptic bronchoscopic balloon dilation in the management of benign tracheobronchial stenosis. Chest 1998;114:796–800. 7. Lee KH, Ko GY, Song HY, Shim TS, Kim WS. Benign tracheobronchial stenoses: long-term clinical experience with balloon dilation. J Vasc Interv Radiol 2002;13:909–14. 8. Kim JH, Shin JH, Shim TS, et al. Results of temporary placement of covered retrievable expandable nitionl stents for tuberculous bronchial strictures. J Vasc Interv Radiol 2004;15:1003–8. 9. Kim JH, Shin JH, Shim TS, et al. Efficacy and safety of a retrieval hook for removal of retrievable expandable tracheobronchial stents. J Vasc Interv Radiol 2004;15: 697–705.


Case report: Deep tracheal laceration after balloon dilatation 10. Petersen BD, Uchida BT, Barton RE, Keller FS, Rosch J. Gianturco-Rosch Z stents in tracheobronchial stenoses. J Vasc Interv Radiol 1995;6:925–31. 11. Brown SB, Hedlund GL, Glasier CM, Williams KD, Greenwood LH, Gilliland JD. Tracheobronchial stenosis in infants: successful balloon dilation therapy. Pediatr Radiol 1987;164:475–8. 12. Hebra A, Powell DD, Smith CD, Othersen HB Jr. Balloon tracheoplasty in children: results of a 15-year experience. J Pediatr Surg 1991;26:957–61. 13. Elkerbout SC, van Lingen RA, Gerriten J, et al. Endoscopic balloon dilatation of acquired airway stenosis in newborn infants: a promising treatment. Arch Dis Child 1993;68:37–40. 14. Kang SG, Song HY, Lim MK, Yoon HK, Goo DE, Sung KB. Esophageal rupture during balloon dilation of strictures of benign or malignant causes: prevalence and clinical importance. Radiology 1998;209:741–6.


¨ P. Long-term results of 15. Solt J, Bajo J, Szabo´, Horva´th O balloon catheter dilation for benign gastric outlet stenosis. Endoscopy 2003;35:490–5. 16. Kim JH, Shin JH, Di ZH, et al. Benign duodenal strictures: treatment by means of floroscopically guided balloon dilation. J Vasc Interv Radiol 2005;16:543–8. 17. Brown SB, Hedlund GL, Glasier CM, Williams KD, Greenwood LH, Gilliland JD. Tracheobronchial stenosis in infants: successful balloon dilation therapy. Radiology 1987;164:475–8. 18. Kim HJ, Shin JH, Hong SJ, Park SJ, Lee SY, Song HY. Treatment of congenital tracheal stenosis with balloonassisted posterior tracheal splitting and temporary placement of a covered retrievable metallic stent. J Vasc Interv Radiol 2005;16:287–91. 19. Madden BP, Datta S, Charokopos N. Experience with ultraflex expandable metallic stents in the management of endobronchial pathology. Ann Thorac Surg 2002;73:938–44.

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Any fool can understand IMRT

(Received 13 January 2006 and accepted 27 January 2006) DOI: 10.1259/bjr/77476500

The Editor—Sir, In his paper, Intensity-modulated radiation therapy (IMRT): a clinical reality for cancer treatment, ‘‘any fool can understand this’’ (The 2004 Silvanus Thompson Memorial Lecture) [1], Prof. Webb directs your readers to the famous book by Silvanus Thompson, Calculus made easy, from which his ‘‘any fool…’’ quotation is taken. I, too, would strongly recommend this book to anyone interested in mathematics, but would caution that the book of that title to be ordered via Amazon, as suggested by Prof. Webb, is not the original book as displayed in his Figure 2, but is a re-write of the book by Martin Gardner [2]. Martin Gardner is an outstanding mathematician and popularizer of mathematics, but I have to say I do not believe he has left very much of the original or entirely succeeded in keeping its spirit in his new version. I would recommend looking for a second-hand copy of the original. It was this book which the teenage Richard Feynman, having discovered that ‘‘calculus is a big thing,’’ found and taught himself from. As Feynman later wrote: ‘‘I have since realised that that particular calculus book had its especially screwy methods…and it invented proofs that weren’t proofs…and there were errors in proofs…but of course the proofs were not important’’. Yes indeed – but Thompson’s was one of the earliest attempts to present in an accessible form a subject once thought too difficult to be attempted outside a university mathematics course. Yours etc., P DAWSON Directorate of Imaging UCL Hospitals NHS Trust 235 Euston Road London NW1 2BU

536

References 1. Webb S. Intensity-modulated radiation therapy (IMRT): a clinical reality for cancer treatment, ‘‘any fool can understand this’’. Br J Radiol 2005;78:S64–S73. 2. Thompson S. Calculus made easy. Gardner M, editor. Palgrave Macmillan, 1999.

Author’s reply The Editor—Sir, Thank you for this useful clarification. As may be seen from my Figure 2 [1], I had a version of the original when I was writing. It was probably not too wise to mention Amazon anyway and I concur with Peter that obtaining sight of the original is sound advice. Yours etc., S WEBB Institute of Cancer Research Royal Marsden NHS Trust Downs Road Sutton Surrey SM2 5PT UK (Received and accepted 30 January 2006) DOI: DOI: 10.1259/bjr/77476500 10.1259/bjr/69383091

References 1. Webb S. Intensity-modulated radiation therapy (IMRT): a clinical reality for cancer treatment, ‘‘any fool can understand this’’. Br J Radiol 2005;78:S64–S73.



CASE OF THE MONTH

Neck pain: an unusual presentation of a common disease A C PANKHANIA, MBChB, MRCS, T PATANKAR, MBChB, DMRD, DMRE,

DNBE, FRCR

and D DU PLESSIS, MRCPath

Department of Neuroradiology and Neuropathology, Hope Hospital, Salford, Manchester M6 8HD, UK

Received 15 November 2004 Revised 10 June 2005 Accepted 11 July 2005 DOI: 10.1259/bjr/28763793 ’ 2006 The British Institute of Radiology

A 68-year-old man presented with 1 month history of neck pain, progressively worsening sensory dysfunction in the right hand, weakness of both hands and difficulty walking. On clinical examination he was found to have reduced power in both upper and lower limbs, with upgoing plantar reflexes. MRI was performed using a Philips Gyroscan 1.5 T machine using sagittal T1 weighted turbo spin echo (TSE; repetition time (TR) 400/echo time (TE) 10, matrix 512, field of view (FOV) 2556255, slice thickness 3.0 mm/ 0.3 mm, number of signal averages (NSA) 4), T2 weighted TSE (TR 3500/TE 120, matrix 512, FOV 2556255, slice thickness 3.0 mm/0.3 mm, NSA 4) and axial T2 weighted turbo field echo (TFE, TR 7.8/TE 3.9, flip angle 45 ˚, matrix 512, FOV 2256225, slice thickness 3.5 mm/21.8 mm, NSA 3) sequences. Post-contrast sagittal (as above) and axial T1 weighted TFE (TR 9.4/ Address correspondence to: Dr Tufail Patankar, 30 Windy Hill Drive, Bolton BL3 4TH, UK.


TE 4.6, flip angle 25 ˚, matrix 512, FOV 24562.4, slice thickness 4.0 mm/22.0 mm, NSA 4) sequences were also performed from C3 to D1 level. MRI demonstrated a destructive lesion involving the right facet joint of C4/5 associated with a medially placed extradural mass of intermediate signal on T1 weighted images, intermediate heterogeneous signal on T2 weighted images which showed peripheral contrast enhancement postgadolinium (Figure 1). Enhancement was also present in the joint and surrounding soft tissues. The soft tissue mass was compressing and displacing the spinal cord and intrinsic high signal was present in the cord on T2 weighted sequences. Similar but less severe changes were also present in the right C2/C3 facet joint. A CT scan performed to look for bony changes revealed subtle eggshell calcification noted around the extradural mass and well-defined erosive changes involving the facets of C2/C3, C3/C4 and C4/C5 (Figure 2). What is the differential diagnosis?

537

A C Pankhania, T Patankar and D Du Plessis

Figure 1. MRI showing a destructive lesion involving the right facet joint of C4/5 associated with a medially placed extradural mass compressing the spinal cord and of intermediate signal on (a) T1 weighted images, (b) heterogeneous intermediate signal on T2 weighted images which showed (c) peripheral contrast enhancement post-gadolinium (arrow demonstrates well-defined facet joint erosion).

Figure 3. Low power image showing tophaceous deposits

Figure 2. CT scan shows well defined erosive change involving the right C4/C5 facet joint.

538

[black arrows] against a background of fibrosis and focal chronic inflammation (haematoxylin and eosin stain, original magnification 650). Inset: Birefringent needle shaped urate crystals [white arrow] demonstrated by polarised light (original magnification 6630).


Case of the month: Neck pain

Imaging findings suggested a diagnosis of spinal gout. The patient had no history or evidence of gout and denied weight loss or trauma. Subsequently, the serum urate level was found to be 0.49 mmol l21 (normal values 0.24–0.50 mmol l21). Inflammatory markers were normal except for a minimally raised C-reactive protein, which was 22 mg l21 (normal ,10 mg l21). Posterior surgical decompression and debulking of the extradural mass was undertaken. Histology of the mass showed areas of chronic inflammation and necrosis. Some of the areas of the necrosis showed birefringent needle shaped crystalline structures consistent with urate crystals (Figure 3). The diagnosis was therefore made of gout related arthropathy with tophus formation. A good post-operative recovery was made and the patient received medical therapy for gout.

Discussion Gout is a common metabolic disorder characterized by episodes of recurrent arthritis and the presence of monosodium urate in the affected tissues. The disease tends to affect distal joints but involvement of the axial skeleton, though rare, has been reported [1, 2] with less than 40 cases reported in the world literature [3]. The distribution between cervical, thoracic and lumbar spine is debatable [1, 2]. The imaging features of spinal gout can be non-specific and can mimic infectious, inflammatory, degenerative or neoplastic disease. The MR appearances are defined to a great extent by the tophus but are variable. The tophi are low to intermediate signal on T1 weighted images, but


may be homogeneously low or high in signal on T2 weighted images [2, 4]. The most common pattern is homogeneous intermediate signal on T1 and heterogeneous intermediate to low signal on T2 weighted imaging [2], which are similar to the appearances seen on MR in our patient. The variability of signal characteristics on MR is thought to be due to variable levels of calcium deposition within the tophus [2, 4]. The tophus may show homogeneous enhancement or heterogeneous peripheral enhancement following gadolinium [2] as in this case. Our case is unusual in its unilateral involvement of the facets and that the patient had no radiological or clinical evidence of gout [2, 5]. Infection and neoplastic process were excluded on imaging because of multilevel involvement centred on the facet joints. An inflammatory condition such as rheumatoid disease was considered unlikely as there was no involvement of other joints.

References 1. Duprez TP, et al. Gout in the cervical spine: MR pattern mimicking disk vertebral infection. AJNR Am J Neuroradiol 1996;17:151–3. 2. Hsu C-Y, et al. Tophaceous gout of the spine: MR imaging features. Clin Radiol 2002;57:919–25. 3. Barrett K, Miller ML, Wilson JT. Tophaceous gout of the spine mimicking epidural infection: report and review of the literature. Neurosurgery 2001;48:1170–3. 4. Yu JS, et al. MR imaging of tophaceous gout. AJR Am J Roentgenol 1997;168:523–7. 5. Kaye PV, Dreyer MD. Spinal gout: an unusual clinical and cytological presentation. Cytopathology 1999;10:411–4.

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Book reviews Computer imaging: digital image analysis and processing. By S E Umbaugh. pp. 696, 2005 (CRC Press, Taylor & Francis Group, Boca Raton, FL), £38.99 ISBN 0-8493-2919-1 This is a comprehensive text-book dealing with all aspects of the acquisition and processing of visual information by computer. The treatment given is oriented towards application of the techniques and examples are taken from a wide range of fields, which includes medical imaging but deals with other areas such as the entertainment industry and space exploration. The book is intended primarily for use by university students and staff in teaching image processing courses or research, but will be useful for professionals working in the commercial sector, government research or the health service. The book is designed for use by scientists and engineers. It is well written and a strong computing background is not required in order to gain a conceptual understanding of the subject from the book. However, far more can be gained if the reader has computing and programming skills. Tutorial exercises are included with each chapter to provide hands-on experience and enable the reader to gain insight into the use of various image analysis and processing algorithms. The book is illustrated throughout with images to demonstrate the effects of the processing techniques described. Since the book is not specifically aimed at the medical sector, it is not a book to dip into to find the techniques applicable to a particular aspect of medical imaging, but rather to obtain a grounding in image analysis and processing techniques. The book is divided into four sections. The introduction to computer imaging presents a global picture to enable the reader to gain an understanding of the overall process. It contains the basic concepts required to understand computer imaging, including optics, imaging analysis and presentation. The second and third sections on digital image analysis and image processing make up over three-quarters of the text. Section 2 describes the tools, concepts and models required for analysing digital images, including segmentation, transforms and feature analysis. Section 3 on image processing starts with visual perception and discusses the application of processing images for human consumption, including topics such as enhancement, restoration and compression. Section 4 contains information on computer vision and image processing tools software developed at the author’s department. A windows version of the software is on a CD accompanying the book. This has a menu-driven user interface to facilitate applications and is designed to allow the reader to apply the algorithms. A knowledge of C and C++ programming will allow the user to develop the algorithms further for their own applications. The book provides a valuable text for those wishing to study and develop skills in image analysis and processing. It differs from other texts by including an insight into the application of the techniques in many different fields. C J MARTIN 540

Coronary radiology. By M Oudkerk. pp. x+254, 2004 (Springer-Verlag, Berlin, Heidelberg, New York), £107.50 ISBN 3-540-43640-5 This book provides an overview of the various imaging modalities used to assess the coronary vessels. Five main areas are covered: anatomy, invasive imaging, noninvasive imaging, calcification and multidimensional computed coronary visualization. The invasive imaging section includes intracoronary ultrasound. Non-invasive imaging principally outlines coronary CT, but also addresses MR coronary angiography and electron beam CT. The calcification chapter makes up a significant proportion of the book and outlines coronary calcification imaging, its epidemiology and the clinical implications. In depth coverage of imaging rendering techniques is found in the final chapter. This particular section would also be useful to imaging in other areas of the body. The large sections are divided into clearly set short chapters which lend this book well to reading in multiple short time periods. Furthermore, the text flows well and it is an easy read. Correlation between modalities is frequently displayed particularly with regard to conventional angiography positioning and CT angiography positioning. This is evident in the demonstration of anatomy with direct correlations between diagrammatic angiographic positioning with conventional angiograms and CT angiographic images displayed together. The images are clear, up to date and well set out. The use of diagrams to assist explanation is frequent and works well. References are displayed throughout the text with lists found at the end of each chapter. Overall this book allows the reader to gain an understanding rather than simply use for reference. B HOLLOWAY Molecular basis of breast cancer: prevention and treatment. By J Russo and I H Russo. pp. xiv+448, 2004 (SpringerVerlag, Berlin, Heidelberg, New York), £154.00 ISBN 3-540-00391-6 This magnum opus covers the entire molecular basis of breast cancer in ten self-contained chapters. Epidemiological considerations in breast cancer are covered in chapter one. This reviews geographical influences, ionizing radiation, electromagnetic field, reproductive aspects and environmental factors, including smoking and alcohol. This is an interesting introduction to the subject. The development of the breast is described in the second chapter in great detail. A knowledge of breast development is critical in understanding breast cancer risks, such as its inverse relationship with early parity. The level and depth of information is excellent. The changes in breast structure from intrauterine life to old age are described in detail. Endocrine control of breast development is covered in chapter three. The response of breast tissue to oestrogen The British Journal of Radiology, June 2006


July 2006 Volume 79 Issue 943

July 2006, Volume 79, Issue 943 ●

Hypoxia in biology and medicine: the legacy of L H Gray

●

Controversies in non-accidental head injury in infants

Advanced imaging applied to radiotherapy planning in head and neck cancer: a clinical review ●

Radiofrequency ablation in pig lungs: in vivo comparison of internally cooled, perfusion and multitined expandable electrodes ●

PET/CT detects abdominal wall and port site metastases of colorectal carcinoma ●

Quantitative colorectal cancer perfusion measurement by multidetector-row CT: does greater tumour coverage improve measurement reproducibility? ●

Prophylactic implantation of inferior vena cava filter during interventional radiological treatment for deep venous thrombosis of the lower extremity ●

●

A survey of MRI quality assurance programmes

Influence of menopausal status and use of hormone replacement therapy on radiation dose from mammography in routine breast screening ●

Kodak EDR2 film for patient skin dose assessment in cardiac catheterization procedures ●

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Randomized phase II study of GM-CSF to reduce mucositis caused by accelerated radiotherapy of laryngeal cancer

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Magnetic resonance urography: a pictorial overview

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Focal pancreatic lesion: can a neoplasm be confidently excluded?


COMMENTARY

Hypoxia in biology and medicine: the legacy of L H Gray 1

P P DENDY,

PhD

and 2P WARDMAN,

DSc

1

1A Coppice Avenue, Great Shelford, Cambridge CB2 5AQ and 2Gray Cancer Institute, PO Box 100, Mount Vernon Hospital, Northwood, Middlesex HA6 2JR, UK

Received and accepted 15 March 2006. DOI: 10.1259/bjr/13634453 ’ 2006 The British Institute of Radiology

On 10 November 2005, the centenary of the birth of Louis Harold Gray FRS (1905–1965), a meeting was held at Blenheim Palace near Oxford to mark the anniversary and to bring together scientists who had worked with Gray and others specializing in the field of tumour hypoxia – which is arguably his main legacy. This Commentary cannot summarize all the presentations, but aims to provide, in this context, a brief overview of the current understanding of hypoxia in relation to radiotherapy and other wider biological and medical implications. We illustrate how progress in imaging hypoxia and understanding changes in gene expression linked to hypoxia provide important avenues towards molecular targeting in radiation oncology, and should eventually lead to individualization of patient treatment to an extent that Gray could only have dreamed of.

The scientific career of L H Gray Barry Michael opened the meeting by outlining Gray’s scientific career. ‘‘Hal’’, as he was known to colleagues, studied physics under Rutherford at the Cavendish Laboratory, Cambridge. His mentors at Cambridge included J J Thompson, Chadwick, Cockcroft, C T R Wilson, Aston and Kapitza; a veritable ‘‘who’s who’’ of physicists of the early 1930s. At an early stage, Gray decided he wished to use his knowledge in a practical way and turned to biology and medicine. He was elected President of the British Institute of Radiology for 1949– 1950 and in 1952 he delivered the 32nd Silvanus Thomson Memorial Lecture, entitled The initiation and development of cellular damage by ionising radiations [1]. The link between tumour blood supply, oxygen delivery and radiosensitivity had been discussed by Mottram in this Journal as early as 1936 [2], but in 1953 and 1955 Gray and his colleagues published two seminal papers [3, 4], recognizing that it might be possible to manipulate the oxygen status to improve radiotherapy. Thus the first paper [3] stated: ‘‘Consideration is given to the supply of oxygen to tissues as a factor in radiotherapy, and it is The British Journal of Radiology, July 2006

concluded, on the basis of existing knowledge, that in certain circumstances the effectiveness of X-ray treatment might be increased if the patient were breathing oxygen at the time of irradiation.’’ In the later paper [4], Tomlinson and Gray showed that the histological pattern of necrosis in sections of some human lung tumours was consistent with that expected if the supply of oxygen was the limiting factor determining the onset of necrosis. For the remainder of his life, Gray led an enthusiastic team establishing radiobiology as a new, rigorous, scientific discipline. They worked in many areas, but the challenge of hypoxic cells always played a central role and, to this day, remains an increasingly complex problem.

Hypoxia and radiotherapy: the position in the late 1990s The Fifth Edition of Eric Hall’s monograph [5] summarizes the state of knowledge on hypoxic cells and hypoxic cell radiosensitizers in the late 1990s. Some key points are: ‘‘Oxygen ‘‘fixes’’ (i.e. makes permanent) the damage produced by free radicals. In the absence of oxygen, damage produced by indirect action may be repaired. Chronic hypoxia results from the limited diffusion range of oxygen through respiring tissue. Acute hypoxia is a result of the temporary closing of tumour blood vessels and is therefore transient. … There is good evidence that human tumours contain hypoxic cells … evidence includes histologic appearance, oxygen probe measurements, the binding of radioactive nitroimidazoles … There is clinical evidence that hypoxia may play an important role in malignant progression.’’ Hall also summarizes progress on radiosensitizing hypoxic cells, and work towards hypoxia-selective cytotoxins such as tirapazamine. He describes how ‘‘Adams and his colleagues listed properties that would be essential for a clinically useful hypoxic cell sensitizer. Firstly, it had to selectively sensitize hypoxic cells at a concentration that would result in acceptable toxicity in 545

P P Dendy and P Wardman

normal tissues … be chemically stable and not subject to rapid metabolic breakdown … highly soluble in water or lipids … capable of diffusing a considerable distance through a non-vascularized cell mass to reach hypoxic cells … effective at … doses of a few grays.’’ Several drugs of the nitroimidazole family had been tested as radiosensitizers. Misonidazole had a higher electron affinity and was more effective than metronidazole, being very effective in cells in culture and with animal tumours but with poor results in clinical trials. Related compounds evaluated included etanidazole and nimorazole, which had less toxicity because of shorter biological half-lives or reduced passage across the blood– brain barrier. Nimorazole was shown to be of benefit in head-and-neck cancer in Danish trials, consistent with an earlier meta-analysis by the Danish group of all randomized trials of hypoxia modification indicating an odds ratio of 1.3 [6]. A conclusion from these studies has been that the clinical trials of hypoxic cell radiosensitizers were compromised by a lack of knowledge of the hypoxic status of an individual patient’s tumour. Now, however, newer ‘‘bioreductive’’ drugs selectively toxic to hypoxic cells without radiation have been identified. Currently, tirapazamine is the lead compound in clinical trial. Interestingly, these bioreductive drugs also have a dependency on electron affinity for their activity.

Changes in gene expression In recent years, hypoxia-induced changes in gene expression have been extensively demonstrated; two speakers (Bradly Wouters and Christopher Pugh) discussed recent developments. Hypoxia inducible factor (HIF)-1 is a heterodimeric transcription factor made up of a and b subunits and it was first recognized as the DNA binding factor that mediates the hypoxia-induced expression of the erythropoietin gene. HIF-1 may promote either directly or indirectly the expression of as many as 60 target genes. Hydroxylation of HIF-1a to an inactive form has an absolute requirement for molecular oxygen and prolyl hydroxylase enzymes: HIF-1a escapes inactivation in hypoxia and hence can bind HIF-1b, form the functional HIF-1 complex and drive gene expression. Hypoxia-inducible genes are known to be involved in regulation of biological processes associated with malignancy [7]. HIF-1 can regulate expression of many enzymes in the glycolytic pathway, as well as processes involved in genetic instability, tissue invasion and metastases. HIF-1 also has a clear role in the regulation of genes involved in angiogenesis, both in normal development and in tumours [8]. In all the above roles HIF-1 is likely to promote tumour growth but there are a few situations where HIF-1 (and other isoforms of HIF) can have a negative regulatory effect on tumour response. The two speakers also discussed different geneticallyrelated issues. Bradly Wouters pointed out that other mechanisms are required to explain the biological response to acute, rather than chronic, hypoxia. He presented new evidence that a potentially important point for regulating gene expression that is able to respond rapidly to changes in the microenvironment is 546

the process of RNA translation (protein synthesis). Global mRNA translation is severely, but reversibly, inhibited during hypoxic conditions, but Wouters and colleagues have shown [9] that, in HeLa cells and prostate carcinoma cells in vitro, this averaging process obscures wide variations in behaviour at the level of individual genes. Indeed a significant number of mRNA species are not dependent on the translation factors that are inhibited during hypoxia and in this efficiently translated fraction of mRNA, 120 genes were more than 4-fold up-regulated by hypoxia. Pugh pointed out that although HIF is frequently upregulated in cancer, genetic studies have not always supported a simple model in which up-regulation of HIF promotes a specific biological process associated with malignancy, e.g. angiogenesis, and hence tumour growth [8]. He postulated that genetic mutation may affect the function of an extensive physiological pathway. Thus clonal selection of a particular property affects a package of properties which, individually, could contribute positively, negatively, or not at all to the overall advantage driving selection of the clone. For example, cellular proliferation, HIF activation and angiogenesis might be co-selected because they were linked by pathways that operate physiologically to preserve oxygen homeostasis. Pugh concluded that understanding the HIF system has many implications for cancer biology and interference with this system may have therapeutic uses. However, we will need to know a lot more about the consequences of intervention at any particular point in the development of malignancy to be certain of a net positive benefit.

Labelling and mapping hypoxic cells There have been important developments in markers of hypoxia, especially markers detectable by immunohistochemistry. These markers require no additional intervention beyond an initial pre-treatment biopsy, which is used to generate formalin-fixed or frozen sections, and may be very suitable for widespread clinical use. These markers are more applicable than oxygen electrodes and provide a high resolution assay of the distribution of hypoxia at the microregional level. Albert van der Kogel discussed imaging the dynamics of tumour hypoxia using immunohistochemical markers [10]. Pimonidazole was injected before biopsy into patients with head and neck cancer, and this can be regarded as an exogenous marker of hypoxia. Following sectioning, carbonic anhydrase CA9 (a HIF-dependent gene that has been investigated as a putative endogenous hypoxia marker) and blood vessels could also be visualized by immunohistochemistry. Beautiful triple-staining images of pimonidazole, CA9 and vessels were shown. Kaanders et al [10] confirmed the widespread variation in hypoxia in human tumours of the head and neck – in 43 squamous cell carcinomas, tumour area staining positive for pimonidazole ranged from 0.3% to 17.2%. Follow up of patients after treatment showed a significant correlation between vascular density and locoregional control, and a significant negative correlation between pimonidazole binding and both locoregional control (15 months) and disease-free survival (2 years). The British Journal of Radiology, July 2006

Commentary: Hypoxia – the legacy of L H Gray

In the laboratory, more sophisticated experiments are possible because of the use of experimental models. For example, van Laarhaven et al [11] injected two nitroimidazole markers of hypoxia, CCI-103F and pimonidazole, before and after treatment with nicotinamide and carbogen (95% O2, 5% CO2) singly and in combination into two different murine carcinomas. Bromodeoxyuridine and Hoechst 33342 were also used as proliferation and perfusion markers, respectively. The two tumours showed extensive differences in vascular architecture, distribution patterns of hypoxia and bromodeoxyuridine labelling. All treatment combinations caused a decrease in the hypoxic fraction, but the responses of the two tumours were quantitatively very different. As an alternative to image analysis of histological sections, flow cytometry was used by Bennewith and Durand [12] to obtain important information about the transient nature of hypoxia. Pimonidazole was injected hourly into a human tumour xenograft for 8 h. This causes a time-integrated asymptotic rise in the number of cells showing the hypoxic marker. 1 h before sacrifice, a second hypoxic marker (CCI-103F) was injected. Examination by flow cytometry of the single and dual fluorescence peaks from a large number of cells showed that substantial numbers of cells that had been hypoxic were no longer hypoxic immediately before sacrifice. This dynamic behaviour is likely to have a big influence on tumour management and particularly the application of hypoxia-directed therapies. Methods of monitoring hypoxia using non-invasive imaging techniques are being explored. Nitroimidazoles labelled with I-123 may be suitable for conventional gamma camera imaging or with F-18 for PET imaging. A recent paper by No¨th et al [13] described the use of 15C5loaded alginate capsules as fluorine-19 oxygen sensors in MRI for in vivo determination of tumour oxygenation during growth and in response to carbogen breathing.

Molecular targeting in radiation oncology Addressing this subject, Michael Baumann pointed out that radiotherapy is very efficient at reducing a tumour to a small number of surviving clonogenic cells. There is therefore scope for developing novel therapeutic agents which, although perhaps not curative in themselves, may be highly effective if used in combination with an appropriate therapeutic regimen. Hypoxia is known to have a role in: (i) selecting for cells that have lost sensitivity to the tumour suppressor gene p53; (ii) regulating genes involved in drug resistance; (iii) a tendency to select for a more malignant phenotype; (iv) increasing the mutation rate; (v) increasing expression of genes associated with angiogenesis; and (vi) tumour invasion. There is therefore considerable potential for developing new approaches to therapy based on targeting hypoxia. Two examples discussed by Baumann are summarized below. A number of groups have shown a link between expression of epidermal growth factor receptor (EGFR) and tumour growth. For example, Ang et al [14] showed that in a large series of patients with head and neck squamous cell carcinomas, EGFR expression was a The British Journal of Radiology, July 2006

strong independent prognostic indicator of overall and disease-free survival. It was also a robust predictor of locoregional relapse, but not of distant metastases. The authors recommended that EGFR immunohistochemistry should be considered for selecting patients for more aggressive combined therapies or enrolment in trials targeting EGFR signalling pathways. Baumann and Krause [15] had reviewed the evidence that inhibition of EGFR can increase radiosensitivity of clonogenic cells and tumour cell proliferation, and Krause et al [16] have shown that in nude mice the anti-EGFR monoclonal antibody can significantly reduce the tumour control dose with decreased repopulation and reoxygenation. The second example targets tumour angiogenesis. Adjuvant inhibition of the vascular endothelial growth factor receptor (VEGFR) with a tyrosine kinase inhibitor after fractionated irradiation prolongs tumour growth [17]. Zips et al [18] have confirmed this finding for a human squamous cell carcinoma grown in nude mice and exposed to the VEGFR inhibitor for 75 days after irradiation, but showed that the 50% tumour control dose was no different. The authors conclude that recurrences depend on VFGF-driven angiogenesis but surviving tumour cells retain their clonogenic potential during this treatment. Other strategies for targeting therapies based on hypoxia, not discussed in detail at the meeting, include hypoxia-activated prodrugs, hypoxia-selected gene therapy, and the use of genetically-engineered anaerobic bacteria [19]. Molecular-based approaches targeted to hypoxia-mediated processes add a further dimension to the complexity of the problem since overall and diseasefree survival, locoregional control, and control of metastases rarely give concordant results in animal experiments and clinical trials.

Clinical trials Van der Kogel also outlined the background to the present Phase 3 clinical trials of ARCON (accelerated radiotherapy with carbogen and nicotinamide) in head and neck and bladder cancers [20]. This concept, which was pioneered at the Gray Laboratory, is designed to attack tumour cells that may have more than one resistance mechanism (i.e. acute and chronic hypoxia) and therefore may be responsive to combined modality treatments. Furthermore, tumour repopulation is a major cause of poor local tumour control, which generally deteriorates as treatment time increases (at fixed dose). Thus the first arm of the ARCON attack is to use accelerated fractionated radiotherapy with several fractions per day. Inhalation of carbogen is designed to decrease diffusion-limited hypoxia and nicotinamide is used to decrease perfusion-limited hypoxia, although it may also have other effects. ARCON is a good example of translational research. Pre-clinical studies, mainly on animals, have shown that each of the three components can be effective, both in isolation and in combination. For example, tumour control rate for a mouse mammary carcinoma was the same with the combination treatment with almost 50% less radiation dose. Phase 1 and 2 clinical trials showed that the ARCON regimen was feasible and tolerable in 547

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patients and produced promising results in terms of tumour control. For the Phase 3 trial that is nearing completion, selection of patients has been mainly based on clinical and histopathological tumour characteristics. It is already clear that, whatever the outcome, improved selection of patients based on measures of hypoxia will be desirable, if not essential, to achieve good therapeutic outcomes, as may now be appreciated from previous hypoxic cell radiosensitizer trials.

Widening horizons Two speakers took a broader view of some aspects of radiobiology. Using the title Science spanning the generations, Ian Stratford reviewed a number of other areas in which early work by Gray was taken forward by colleagues and later generations with varying degrees of success. One phenomenon that attracted the attention of Gray and co-workers was the effect of nitric oxide on the radiosensitivity of tumour cells [21, 22]. Nitric oxide appeared to be able to take the place of oxygen in enhancing sensitivity to X-rays in bacteria, plant meristem and mammalian tumour cells. Furthermore, cell survival curves showed that sensitization by nitric oxide was exactly the same as that achieved with oxygen when added to anoxic conditions. It is now known that nitric oxide is a biochemical signalling molecule, generated in tissues by nitric oxide synthase (NOS). NOS levels are elevated in a wide range of tumour types (see, for example, Ambs et al [23]). Stratford showed that cytokine-mediated induction of NOS radiosensitized tumour cells at intermediate oxygen tensions in vitro, hence suggesting that varying levels of NOS in human tumours could be a significant factor in radiotherapy. Another topic studied by Gray was the dissociation of oxyhaemoglobin [24]; Adams, Stratford and colleagues subsequently explored enhancement of hypoxia by compounds such as BW12C which influenced this dissociation. They showed that these compounds could increase tumour hypoxia and thereby be useful when combined with hypoxia-selective bioreductive drugs [25]. An alternative approach has been taken in the development of efaproxiral (RSR13), which reduces haemoglobin–oxygen binding to decrease hypoxia [26]; this agent is now in clinical trial in combination with radiotherapy. Stratford also discussed the extensive research on the development of bioreductive drugs worldwide, some pioneered at the Gray Laboratory [19, 25, 27], and the potential for utilizing the presence of hypoxia in tumours to drive selective gene therapy [28]. Finally, he returned to the impact of modulating HIF function on tumour response to radiotherapy, presenting data to support the hypothesis that in HIF-1 deficient tumours, hypoxic cells have an extremely limited lifetime and therefore may not contribute to radiation sensitivity [29]. Gillies McKenna outlined the research programme of the newly-established Department of Radiation Oncology and Biology at the University of Oxford, which will incorporate the Gray Laboratory when it is moved from its current location at Mount Vernon Hospital. The vision of the Department is to continue the tradition of the ‘‘Gray Lab’’; be research focused and 548

clinically-relevant; learn lessons from the past; be flexible to encompass all relevant areas in the future; complement existing groups in Oxford and elsewhere in the UK; and be a centre for training all health care personnel in radiation oncology and biology. Multidisciplinary collaboration will be a key feature of the Department, and two examples were given. First, in the 21st century it will be necessary to move on from empirical therapy, tailored by experience, to targeted therapy. Maximizing the information from tumour imaging will involve diagnostic radiologists, medical physicists, radiographers, and others. Pushing back the boundaries of physical treatment involves radiotherapists, radiation physicists and industry. To these groups must be added a broad spectrum of biologists who will open up new frontiers. McKenna and his colleagues are seeking to define biological processes that regulate the radiation responsiveness of solid tumours, and take agents that target these processes into the clinic in order to improve therapeutic outcome. Second, there will be strong emphasis on translational research. Examples of areas where laboratory workers and clinicians must work in close collaboration include signal transduction inhibitors with radiation, systematic targeting of radiation with Auger electronemitting radionuclides, molecular targets for modifying the radiation response, and the molecular basis of metastases.

Conclusions Hal Gray was a true polymath: introducing their biography of him, Loutit and Scott discussed Gray’s ‘‘escape from this confinement’’ of the ‘‘cult of specialization for learning more and more about less and less’’ [30]. One of us (PW), writing in 1982 in a booklet to commemorate the silver jubilee of the opening of the Gray Laboratory, asserted ‘‘… [Gray] must have been the first – and quite possibly the last – scientist to have a thorough appreciation of current activity in all four sectors of radiation research – physics, chemistry, biology and medicine.’’ While no one subject can sum up his work as a scientist, overcoming tumour hypoxia is a challenge that requires such a broad approach. Had he not died so young he might well have received a Nobel Prize. Some of his contributions are now well understood, for example radiation dosimetry, where Gray’s name is recorded for posterity as the SI unit of absorbed dose. Other areas, especially hypoxia in tumours, represent a far more complex problem than Gray could possibly have imagined, with heterogeneity of behaviour in almost every aspect investigated. In future, human tumours must be better characterized on an individual basis. Performing clinical trials on unselected patients who have a mixture of hypoxic and better-oxygenated tumour cells runs the clear risk of rejecting a treatment that could be of significant benefit to a sub-set of patients. Loutit and Scott described Gray as ‘‘the Fellow [of The Royal Society] who fathered radiobiology’’ [30], but it is arguable that drawing attention to the importance of tumour hypoxia is the most important legacy of L H Gray. The British Journal of Radiology, July 2006

Commentary: Hypoxia – the legacy of L H Gray

Acknowledgments We thank Prof. I J Stratford for helpful comments on a draft manuscript. PW is supported by Cancer Research UK.

15.

References

16.

1. Gray LH. The initiation and development of cellular damage by ionising radiations. Br J Radiol 1953;26:609–18. 2. Mottram JC. A factor of importance in the radiosensitivity of tumours. Br J Radiol 1936;9:606–14. 3. Gray LH, Conger AD, Ebert M, Hornsey S, Scott OCA. The concentration of oxygen dissolved in tissues at the time of irradiation as a factor in radiotherapy. Br J Radiol 1953;26:638–48. 4. Tomlinson RH, Gray LH. The histological structure of some human lung cancers and the possible implications for radiotherapy. Br J Cancer 1955;9:539–49. 5. Hall EJ. Radiobiology for the radiologist, 5th edn. Philadelphia, PA: Lippincott, Williams and Wilkins, 2000:108–9, 458. 6. Overgaard J, Hansen HS, Overgaard M, Bastholt L, Berthelsen A, Specht L, et al. A randomized double-blind phase III study of nimorazole as a hypoxic radiosenstizer of primary radiotherapy in supraglottic larynx and pharynx carcinoma. Results of the Danish Head and Neck Cancer Study (DAHANCA) Protocol 5-85. Radiother Oncol 1998;46:135–46. 7. Harris AL. Hypoxia – a key regulatory factor in tumour growth. Nat Rev Cancer 2002;2:38–47. 8. Pugh CF, Ratcliffe PJ. Regulation of angiogenesis by hypoxia – role of the HIF system. Nat Med 2003;9:677–84. 9. Koritzinsky M, Seigneuric R, Magognin MG, van der Beuchen T, Lambin P, Wouters BG. The hypoxic proteome is influenced by gene specific change in m-RNA transcription. Radiother Oncol 2005;76:177–86. 10. Kaanders JH, Wijffels KI, Marres HA, Ljungkvist AS, Pop LA, van den Hoogen FJ, et al. Pimonidazole binding and tumour vascularity predict for treatment outcome in head and neck cancer. Cancer Res 2002;62:7066–74. 11. van Laarhoven HWM, Bussink J, Lok J, Punt CJA, Heerschap A, van Der Kogel AJ. Effects of nicotinamide and carbogen in different murine colon carcinomas: immunohistochemical analysis of vascular architecture and microenvironmental parameters. Int J Radiat Oncol Biol Phys 2004;60:310–21. 12. Bennewith KL, Durand RE. Quantifying transient hypoxia in human tumour xenografts by flow cytometry. Cancer Res 2004;64:6183–9. 13. No¨th U, Rodrigues LM, Robinson SP, Jork A, Zimmermann U, Newell B, Griffiths JR. In vivo determination of tumour oxygenation during growth and in response to carbogen breathing using 15C5 loaded alginate capsules as fluorine19 magnetic resonance imaging oxygen sensors. Int J Radiat Oncol Biol Phys 2004;60:909–19. 14. Ang KK, Berkey BA, Tu X, Zhang HZ, Katz R, Hammond EH, et al. Impact of epidermal growth factor receptor expression on survival and pattern of relapse in patients

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with advanced head and neck carcinoma. Cancer Res 2002;62:7350–6. Baumann M, Krause M. Targeting the epidermal growth factor in radiotherapy: radiobiological mechanisms, preclinical and clinical results. Radiother Oncol 2004;72: 257–66. Krause M, Ostermann G, Petersen C, Yaromina A, Hessel F, Harstrick A, et al. Decreased repopulation as well as increased reoxygenation contribute to the improvement in local control after targeting of the EGFR by C225 during fractionated irradiation. Radiother Oncol 2005;76:162–7. Williams KJ, Telfer BA, Brave S, Kendrew J, Whittaker L, Stratford IJ, et al. ZD6474, a potent inhibitor of vascular endothelial growth factor signaling, combined with radiotherapy: schedule-dependent enhancement of antitumor activity. Clin Cancer Res 2004;10:8587–93. Zips D, Hessel F, Krause M, Schiefer Y, Hoinkis C, Thames HD, et al. Impact of adjuvant inhibition of vascular endothelial growth factor receptor tyrosine kinase on tumour growth delay and local tumour control after fractionated irradiation of human squamous cell carcinoma in nude mice. Int J Radiat Oncol Biol Phys 2005;61:908–14. Brown JM, Wilson WR. Exploiting tumour hypoxia in cancer treatment. Nat Rev Cancer 2004;4:437–47. Kaanders JH, Bussink J, van der Kogel AJ. ARCON A novel biology-based approach in radiotherapy. Lancet Oncol 2002;3:728–37. Gray LH, Green FO, Howes CA. Effect of nitric oxide on the radiosensitivity of tumour cells. Nature 1958;182:952–3. Dewey DL. The effect of oxygen and nitric oxide on the radiosensitivity of human cells in tissue culture. Nature 1960;186:780–2. Ambs S, Bennett WP, Merriam WG, Ogunfusika MO, Oser SM, Khan MA, et al. Vascular endothelial growth factor and nitric oxide synthetase expression in human lung cancer and the relation to p53. Br J Cancer 1998;78:223–39. Gray LH, Steadman JM. Determination of the oxyhaemoglobin dissociation curves for mouse and rat blood. J Physiol 1964;175:161–71. Adams GE, Stratford IJ. Bioreductive drugs for cancer therapy – the search for tumour specificity. Int J Radiat Oncol Biol Phys 1994;29:231–8. Choy H, Nabid A, Stea B, Scott C, Roa W, Kleinberg L, et al. Phase II multicenter study of induction chemotherapy followed by concurrent efaproxiral (RSR13) and thoracic radiotherapy for patients with locally advanced non-smallcell lung cancer. J Clin Oncol 2005;23:5918–28. Foster JL, Conroy PJ, Searle AJ, Willson RL. Metronidazole (Flagyl): characterization as a cytotoxic drug specific for hypoxic tumour cells. Br J Cancer 1976;33:485–90. Dachs GU, Patterson AV, Firth JD, Ratcliffe PJ, Townsend KM, Stratford IJ, et al. Targeting gene expression to hypoxic tumor cells. Nat Med 1997;3:515–20. Williams KJ, Telfer BA, Xenaki D, Sheridan MR, Desbaillets I, Peters HJ, et al. Enhanced response to radiotherapy in tumours deficient in the function of hypoxia-inducible factor-1. Radiother Oncol 2005;75:89–98. Loutit JF, Scott OCA. Louis Harold Gray 1905–1965. Elected F.R.S. 1961. Biogr Mem Fellows R Soc 1966;12:195–217.

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COMMENTARY

Controversies in non-accidental head injury in infants N STOODLEY,

FRCS, FRCR

Consultant Neuroradiologist, Department of Neuroradiology, Frenchay Hospital, Bristol BS16 1LE, UK ABSTRACT. Non-accidental head injury in infants is not uncommon and is associated with significant morbidity and mortality. It is therefore important to identify it at the earliest opportunity so that appropriate intervention can be made which protects the child from further harm. The whole topic is controversial and the aim of this paper, in question and answer format, is to review some of the more controversial areas to give an overview of the neuroimaging features of this condition. The author has drawn on his clinical and medicolegal experience of these cases, and the review is based upon questions commonly encountered in Court.

Child abuse is a controversial subject and aspects relating to non-accidental head injury (NAHI) are no exception, not least because the limited evidence base hampers professionals working in the field. The main role of the Expert Witness is to interpret and explain to the Court material which is within the Expert’s area of expertise, but which is outwith the experience of the Judge or the jury. The Expert Witness is in the privileged position of being allowed to give opinion evidence to the Court, but this brings a responsibility to ensure that the opinion given is both reasonable and capable of withstanding logical analysis. An Expert Witness has an overriding duty to the Court that takes precedence over any obligation to those Instructing the Expert. The Expert should be independent, impartial and confine their opinions to their particular area of expertise. All sides in Court, will rightly question this evidence and the Expert must be able to justify their opinion. However, just as lawyers’ opinions on the interpretation of points of law sometimes differ, doctors reading the same scientific papers may come to different conclusions; hence controversy!

What are the ‘‘typical’’ neuroimaging features of NAHI? There are no absolutely pathognomonic features that could only be due to NAHI in all cases. Why should there be when we are only looking at the effects of a traumatic episode on specific tissues in a biological system? The brain has limited ways in which it can respond to various insults, so it is not surprising that some cases of accidental trauma show imaging appearances very similar to those seen in NAHI. A fairly common finding in cases of NAHI is of shallow subdural haematomas (SDH) at various, separate sites over the cerebral hemispheres and often in the posterior fossa. The subdural blood may be of different ages, although this is by no means always the case. The 550

Received 23 January 2004 Revised 2 February 2006 Accepted 27 March 2006 DOI: 10.1259/bjr/23921951 ’ 2006 The British Institute of Radiology

most common site for SDH following both accidental trauma and NAHI is over the cerebral convexities. SDH following all but severe accidental trauma, however, tends to be unifocal whereas SDH in NAHI tends to be multifocal. SDH at sites other than over the convexities, especially if in the posterior interhemispheric fissure or posterior fossa, is more likely to occur following NAHI than following accidental trauma [1]. Subdural blood can be seen following an impact injury from any cause, but it is usually related to the site of the impact and is often associated with a fracture. Subdural blood seen away from the point of impact is unusual in accidental trauma, unless severe. Focal parenchymal lesions such as contusions, haematomas and shearing injuries may be seen following both accidental and non-accidental head trauma; when seen following accidental trauma there is usually a very clear history of a major traumatic event. Shearing injuries (diffuse axonal injury or the larger gliding contusions) were thought to be very common in NAHI. However, recent neuropathological and neuroimaging evidence suggests that axonal injury is uncommon even in infants who die following NAHI and that hypoxic–ischaemic changes are more common in these infants [2, 3]. Often there is evidence of reduced grey-white differentiation focally or more generalized on head scans following NAHI and, given the neuropathology, it would seem likely that these scan changes are related to hypoxic–ischaemic changes in the brain.

There is no absolute scientific evidence that shaking causes the injuries in ‘‘shaken baby syndrome’’ is there, doctor? Preferable terms are, abusive head trauma or NAHI which describe what has happened, but do not make any assumptions about mechanism. However, shaking may be an important mechanism in the majority of cases. There must be something very different about the The British Journal of Radiology, July 2006

Commentary: Non-accidental infant head injury

mechanism of injury in NAHI as opposed to accidental head trauma as the clinical presentation is often different, the neuroimaging appearances are different and the outcome, both in terms of mortality and morbidity in the short and the long term, is different [4]. The main differentiating factor could well be the different mechanism of injury. There is no absolute scientific proof that shaking causes these injuries because it is not possible to perform the relevant scientific experiment, of shaking infants of different ages and sizes with different degrees of force and then performing sequential scans; and then studying them clinically and correlating this with scan appearances. In the absence of such data it is necessary to rely on various other sources of evidence including animal experiments, from the neuroimaging appearances of infants following witnessed accidental trauma and imaging evidence from cases of NAHI in which there has been a reliable confession. Primate experiments have been performed in which the animals were subjected to pure translational and/or rotational forces without impact. The head injuries seen in the animals were very similar to those seen in NAHI [5]. Scans of infants and children who have sustained head injuries following accidental head trauma may show the same pattern of SDHs at different sites (as in NAHI), but usually only where the accidents involved severe forces such as following road traffic accidents or major falls. Most domestic accidents do not involve this degree of force and, given the number of domestic incidents that occur daily involving children bumping their heads, the fact that our departments are not inundated with scan requests for infants suggests that the majority of these falls are neurologically benign, as has been shown in the literature [6]. There is debate as to whether shaking alone is sufficient to cause all of the features of NAHI or whether some impact (even against a soft surface) is always necessary. Some of the animal work cited above suggests that impact is not required. There are biomechanical models that show that the amount of force generated during a pure shake is much less than following an impact, but there are also reports in the literature of fatal cases of NAHI where there has been no evidence of impact injury even on post-mortem. The focal brain injuries sometimes seen in NAHI may reflect associated impact or the result of very severe shaking. Some perpetrators admit to shaking and many a loving parent must have been pretty close to it, having been up for hours at night with an inconsolable infant that they have done everything to try to settle. It is possible that only the minority of cases are due to wilful cruelty and some authors have advocated changing the way in which we classify abuse partly because of factors such as these [7].

The subdurals are very thin and not causing significant mass effect, so what is the cause of the presenting symptoms? These infants present with varied and non-specific signs and symptoms which may be the same or similar to those seen in infants who are unwell for a variety of The British Journal of Radiology, July 2006

reasons. These range from being ‘‘off-colour’’ and grizzly, off feeds to vomiting, with various degrees of reduced levels of consciousness, through fits to frank coma and death. The severity of the symptoms and signs does not relate to the size or number of SDH, but does have some relation to the degree of associated brain hypoxic–ischaemic injury and/or swelling. Those infants presenting in coma and fitting are more likely to have generalized brain changes on their initial CT scans when compared with infants presenting with lesser symptoms [8]. There have also been a few publications showing that the pattern of changes on diffusion weighted imaging in NAHI is that of hypoxic–ischaemic change rather than diffuse axonal (shearing) injury [9]. This neuroimaging evidence is supported by the neuropathological literature. Geddes et al published two papers after studying a cohort of fatal cases of NAHI. One paper [2] was a review of 53 cases, 37 infants ranging in age from 20 days to 9 months and 16 children ages ranging from 13 months to 8 years. The authors demonstrated that most of the brain damage seen in these fatal cases was due to hypoxic vascular damage rather than traumatic axonal damage. Traumatic diffuse axonal injury was only seen in three cases. A second paper by this group [3] was based upon 39 infants aged 9 months or under where the authors compared the neuropathology with that seen in 14 control patients. Traumatic diffuse axonal injury was only found in two of the NAHI patients, both of who had suffered severe head injury with multiple fractures. If hypoxic–ischaemic injury is the major cause of death in fatal cases, it is logical to assume that it is an important factor in non-fatal cases, and that the degree of associated hypoxic–ischaemic change largely determines presenting symptoms in the short term as well as outcome in terms of mortality and morbidity. If this is correct then it is likely that infants who sustain a severe injury will lose consciousness at the time of the injury and some may not recover. At the other end of the spectrum, the hypothesis would suggest that there should be a group of infants who sustain an injury which may cause a minor change in behaviour insufficient to cause their carers to take them to a doctor and where the brain (and therefore the infant) recovers from the injury spontaneously. In between these extremes lies a group of infants in whom there is a probably quite a marked change in behaviour following the injury. In less susceptible infants, the brain may be able to recover spontaneously from the injury whereas in more susceptible infants some threshold may be reached beyond which the brain cannot recover from the insult by itself. This latter group of infants would go on to deteriorate clinically and present with an encephalopathic illness. In this way, hypoxic–ischaemic change following NAHI may be in some way analogous to birth asphyxia where infants may develop hypoxic–ischaemic encephalopathy some hours after the insult. It is the different degree and extent of hypoxic– ischaemic injury that determines whether and how these infants will present to medical attention and that the SDHs are just markers of the mechanism of injury. Sometimes the SDHs appear so insignificant on the scan 551

N Stoodley

of an extremely sick infant that they may be overlooked and their true significance not realised.

What else could have caused the SDH? All other possible causes of subdural effusions and haematomas have to be considered in all of these cases. These other causes can usually be diagnosed or excluded with a combination of proper history taking and relevant investigations. Severe previous accidental or birth trauma can lead to the development of SDH. SDH is probably quite common following all modes of delivery but, in otherwise normal infants, they do not persist beyond the first few weeks of life [10]. Infections such as meningitis can lead to infected subdural collections (empyemas) and these could be associated with some haemorrhage, although clinical experience would suggest that bleeding into empyemas is not very common. Infants with coagulopathies may bleed spontaneously or after minimal trauma, but this again is surprisingly uncommon. The pattern of intracranial haemorrhage in infants with clotting disorders also tends to be different as intraparenchymal haemorrhage is much more common than SDH in these infants [11]. Congenital abnormalities including vascular malformations or metabolic conditions such as glutaric aciduria may predispose to SDH.

When did the bleeding occur? Estimating the age of subdural blood on CT and MR is not precise because various factors influence the appearance of blood on both modalities. On CT, acute blood should be of high attenuation, but this assumes that the blood has clotted, that the patient is not severely anaemic and that there has been no significant dilution of the subdural blood by CSF following a traumatic tear of the arachnoid. In severe shaking injuries, CSF dilution of the subdural blood may make acute SDH appear of low attenuation and be misinterpreted as chronic. The time course of attenuation changes is also variable depending on such factors as the volume of blood and the haemoglobin level of the patient at the time of the bleed. Different factors apply to the dating of blood on MR scans, but when there is a combination of CT and MR scans it is often possible to give a range of probable ages for the subdural collections present. This does not necessarily help with the timing of the injury, however, as the SDH is unlikely to be responsible for the presenting symptoms. Usually, the history gives a better assessment of the timing of injury than the radiology because an infant that has suffered a NAHI of sufficient severity to lead to admission to hospital is in my experience extremely unlikely to have behaved completely normally after that injury was inflicted.

What about re-bleeding into chronic SDH? In the elderly population, minimal trauma can trigger re-bleeding into a chronic SDH. The basic pathology must be similar in infants, but the incidence of chronic 552

subdural haematoma in infants is extremely low. In the elderly population, the brain is undergoing involutional change, so any blood within the subdural space may persist allowing the haematoma to become chronic. In a normal infant, the situation is completely different as the brain is growing rapidly and this growth may act to minimize the potential for any material to persist in the subdural space. If the brain does not grow properly due to a congenital or acquired insult of sufficient severity, SDH may also become chronic in infants. The most common previous insult is NAHI. The presence of a chronic subdural haematoma in an infant who has not had an insult such as severe birth injury, a major accident, meningitis or a known underlying brain metabolic abnormality should raise the possibility of previous NAHI [12].

What degree of force is required to produce these injuries? This is unknown. These injuries are not seen following ‘‘normal’’ domestic trauma or rough play, otherwise there would be many cases every day throughout the UK. Therefore, it is likely the degree of force is such that an independent witness would realise that it was likely to cause harm. By the same token, it is possible to severely injure an infant without intending to cause them harm. The active intent in most of these cases is to stop them crying. As we see similar appearances to NAHI following severe accidental trauma, it is sometimes suggested that the degree of force involved in NAHI must be similar to that occurring during road accidents. This is incorrect, as the mechanism of the two injuries is different: high velocity impact with acceleration/deceleration in accidents; lower velocity rotation and apnoea in NAHI.

The radiological investigation of NAHI This has been comprehensively covered in recent publications [13, 14] and the rationale for the imaging approach discussed [15], emphasising the importance of using both CT and MR in these cases. CT should be the initial investigation because it is good at demonstrating acute blood, it is more widely available than MR and it is much simpler to perform a CT scan on a sick infant. MR is better at showing older collections of blood, blood in sites not well seen on CT (such as the middle cranial fossa), low volume haematomas and, of course, is vastly superior to CT in the demonstration of parenchymal brain injuries. Both modalities are therefore required for a full neuroradiological assessment of these infants. Many of these cases are obvious, others are not and we won’t find the evidence unless we look. Think of SDH in the same way as metaphyseal and rib fractures: we look for these fractures assiduously with skeletal surveys but, although they are markers of mechanisms of injury, they are of little long-term consequence. Shouldn’t we be looking for the SDH that is also marker of a mechanism of head injury that causes most of the morbidity and mortality in these cases? The British Journal of Radiology, July 2006

Commentary: Non-accidental infant head injury

Conclusions To miss abuse risks sending the child back into an abusive environment; to suggest abuse where there has been none can tear a family apart. These difficult decisions can only be taken if we have sufficient information on which to base them. The radiologist may be the first clinician to suggest the possibility of child abuse and may therefore become involved in the legal proceedings that inevitably, and rightly, follow many of these cases. In the summer of 2005, four joined appeals against convictions in cases of alleged non-accidental head injury were heard in the Court of Appeal. The appellants had been convicted of murder (1); manslaughter (2) and grievous bodily harm (1). The Court heard from 21 experts and the Judgment [16] contains much invaluable guidance on the approach to take in possible NAHI cases. The Judgment emphasises that each of these cases is fact-specific and each should be determined on their individual facts, and also that not all cases where the socalled triad (encephalopathy, subdural and retinal haemorrhage) is present will be due to NAHI. Having heard the evidence in these four cases, one murder conviction was reduced to manslaughter, two convictions were quashed (manslaughter and grievous bodily harm) and an appeal against a conviction for manslaughter was dismissed.

References 1. Ewing-Cobbs L, Prasad M, Kramer L, Louis PT, Baumgartner J, Fletcher JM, et al. Acute neuroradiologic findings in young children with inflicted or non inflicted traumatic brain injury. Child’s Nervous System 2000;16: 25–34. 2. Geddes JF, Hackshaw AK, Vowles GH, Nickols CD, Whitwell HL. Neuropathology of inflicted head injury in children. 1. Patterns of brain damage. Brain 2001;124: 1290–8.


3. Geddes JF, Hackshaw AK, Vowles GH, Nickols CD, Whitwell HL. Neuropathology of inflicted head injury in children. II: Microscopic brain injury in infants. Brain 2001;124:1299–306. 4. Duhaime AC, Christian CW, Moss E, Seidl TS. Long term outcome in infants with the shaking impact syndrome. Pediatr Neurosurg 1996;24:292–8. 5. Ommaya AK, Gennarelli TA. Cerebral concussion and traumatic unconsciousness. Brain 1974;97:633–54. 6. Warrington SA, Wright CM. Accidents and resulting injuries in premobile infants: data from the ALSPAC study. Arch Dis Childhood 2001;85:104–7. 7. Southall DP, Samuels MP, Golden MH. Classification of child abuse by motive and degree rather than type of injury. Arch Dis Childhood 2003;88:101–4. 8. Kemp AM, Stoodley N, Cobley C, Coles L, Kemp KW. Apnoea and brain swelling in non-accidental injury. Arch Dis Childhood 2003;88:472–6. 9. Biousse V, Suh DY, Newman NJ, Davis PC, Mapstone TB, Lambert SR. Diffusion weighted magnetic resonance imaging in shaken baby syndrome. Am J Ophthalmol 2002;133:249–55. 10. Whitby EH, Griffiths PD, Rutter S, Smith MF, Sprigg A, Ohadike P, et al. Frequency and natural history of subdural haemorrhages in babies and relation to obstetric factors. Lancet 2004;362:846–51. 11. Vorstman EBA, Anslow P, Keeling DM, Haythornthwaite G, Bilolikar H, McShane T. Brain haemorrhage in five infants with coagulopathy. Arch Dis Childhood 2003;88:1119–21. 12. Feldman KW, Bethel R, Shugerman RP, Grossman DC, Grady MS, Ellenbogen RG. The cause of infant and toddler subdural hemorrhage: a prospective study. Pediatrics 2001;108:636–46. 13. Jaspan T, Griffiths PD, McConachie NM, Punt JAG. Neuroimaging for non-accidental head injury in childhood: A proposed protocol. Clin Radiol 2003;58:44–53. 14. Kemp AM. Investigating subdural haemorrhage in infants. Arch Dis Childhood 2002;86:98–102. 15. Stoodley N. Neuroimaging in non-accidental head injury: if, when, why and how. Clin Radiol 2005;60:22–30. 16. R v Harris and Ors [2005] EWCA Crim 1980.

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REVIEW ARTICLE

Advanced imaging applied to radiotherapy planning in head and neck cancer: a clinical review 1

K NEWBOLD, MRCP, FRCR, 2M PARTRIDGE, PhD, CPhys, 2G COOK, MD, FRCP, FRCR, 1S A SOHAIB, BSc, MRCP, 2 1 1 FRCR, E CHARLES-EDWARDS, MSc, CSci, P RHYS-EVANS, FRCS, K HARRINGTON, MRCP, FRCR and 1 C NUTTING, MD, MRCP, FRCR 1

The Royal Marsden NHS Trust, Fulham Road, London SW3 6JJ and 2The Royal Marsden NHS Trust, Downs Road, Sutton, Surrey SM2 5PT, UK ABSTRACT. Head and neck squamous cell carcinoma represents an ideal model to investigate the application of recent advances in medical imaging to radiotherapy planning. Tumours usually remain localized, and are potentially curable with local radiation. The steep radiation dose–response relationships support the strategies of radiation dose escalation to increase local control. Two-dimensional simulator-based planning and CT planning have significant drawbacks in terms of accurate target volume definition. MRI has enhanced soft tissue delineation, but has to be fused with CT to allow dose calculation. Functional imaging using dynamic contrast enhanced CT or MRI sequences may allow improved knowledge of tumour function. Positron emission tomography (PET) may allow further physiological information to be determined. This review summarizes the current techniques in clinical development in this area.

Head and neck cancer (HNC) is the seventh most common cancer in the UK. Whilst early stage disease has a high cure rate, stage III and IV HNC still have poor rates of local control and survival [1]. Surgery and/or radiotherapy aim to achieve locoregional control, and where it enables organ preservation with maintenance of function, radiotherapy is the modality of choice. Failure to achieve locoregional control may be due to geographical miss of primary tumour or nodes due to understaging disease, intrinsic radioresistance of the tumour, or factors such as hypoxia and proliferation known to reduce radiosensitivity of a tumour. Three-dimensional conformal radiotherapy (3D-CRT) and intensity-modulated radiotherapy (IMRT) aim to achieve better locoregional control and improve survival by radiation dose escalation, but such techniques demand more accurate localization of tumour and normal tissues with noninvasive imaging techniques.

Anatomical imaging for head and neck radiotherapy planning Current standard radiotherapy imaging In the UK, the most common methods of target localization for radiotherapy planning include the use of the simulator and CT scanning. In the simulator, standard radiotherapy field borders are placed on identifiable bony landmarks, and simple outlines are taken at levels though the treatment volume. For most 554

Received 3 February 2005 Revised 23 March 2006 Accepted 6 April 2006 DOI: 10.1259/bjr/48822193 ’ 2006 The British Institute of Radiology

patients, this approach gives the radiation oncologist a ‘‘one size fits all’’ treatment. There is limited opportunity for individualization of treatment volumes, although simple customization is possible with reference to bony landmarks within individual cases. It results in treatment fields with large margins around tumours to account for uncertainty of target position and conservative doses due to uncertainties in dose to organs at risk (OARs). CT planning provides a 3D representation of the target volume and OARs allowing improved target definition and also accurate OAR delineation. The use of iodinated contrast agents increases sensitivity, and CT with contrast remains the best modality for defining nodal disease, cartilage invasion, or bone destruction [2, 3]. Reconstruction of high resolution coronal and sagittal sections may aid RT planning. CT images are most accurate when tumour interfaces are with air cavity, fat, or bone such as in the paranasal sinuses or neck, but are less accurate when the tumour interfaces with normal soft tissue such as in the tongue, or when there may be mucosal spread for example in the larynx and hypopharynx [4]. CT planning generates electron density data for photon dose calculation, and calculation of dose– volume histograms. CT planning allows accurate conformal shaping of the radiotherapy portals around the target and shielding of the OARs (Figure 1). CT planning removes many of the uncertainties of position of tumour and OAR. However, CT planning per se does not necessarily allow radiation dose escalation, especially if the adjacent OARs are neurological. The spinal cord, brain stem and optic nerves are assumed to have a serial The British Journal of Radiology, July 2006

Review article: Advanced imaging applied to radiotherapy planning in HNC

Figure 1. Three-dimensional conformal radiotherapy planning. CT plan of a left sided oropharyngeal tumour. Blue colourwash: planning target volume (PTV), isodose levels indicated by coloured linear boundaries. ANT: anterior beam, LAO: left anterior oblique, LPO: left posterior oblique.

organization of functional subunits, where inactivation of a single subunit causes loss of function of the whole organ. In this setting, the risk of late toxicity is determined by the maximum dose received by the OAR, and therefore partial reductions of the volume of tissue irradiated, such as those offered by 3D-CRT, do not reduce the risk of late toxicity. In such cases the more advanced radiation delivery techniques such as IMRT, which can produce dose distributions with concave isodose surfaces, may be required [5, 6].

Use of MRI for treatment planning in head and neck cancer The main advantage of MRI over CT is significantly greater soft tissue contrast, permitting better definition of disease extent and OAR [7]. T1 weighted images give good anatomical detail whilst T2 weighted images can differentiate between normal and pathological tissues. Image contrast can be enhanced by intravenous gadolinium. Artefacts arising from dental amalgam can be made significantly less conspicuous on MRI than CT. MRI is the modality of choice for imaging tumours of the base of tongue and lesions arising at the skull base [8, 9], also visualization of important OAR such as the orbit, optic nerves and chiasm and central nervous system. Emami et al showed that the co-registration of two anatomical modalities, MRI and CT improved the target definition of cancer of the nasopharynx [10]. The use of MRI in radiotherapy treatment planning is limited by the presence of geometrical distortions, including inhomogeneities within the magnetic field and non-linearities in the spatial-encoding gradients. These distortions increase as the distance from the isocentre increases and have been reported to be 5 mm at 12 cm from isocentre [11]. These therefore are of clinical significance for head and neck planning, and need to be either removed by a post-processing technique [12, 13], or by CT/MR fusion and image correction. Electron density information necessary for treatment dosimetry cannot be obtained from MRI, and the images have to be fused with CT data if this information is to be used in treatment planning calculations. Therefore fusion of distortioncorrected MRI and CT images could provide both the improved target definition and dosimetric accuracy required for treatment planning. The British Journal of Radiology, July 2006

Intrinsic susceptibility-weighted or blood oxygenation level dependent (BOLD) MRI exploits the differences in magnetic susceptibility of oxyhaemoglobin and deoxyhaemoglobin. Changes in blood oxygenation can therefore be characterized by looking at differences between T2* weighted images during a change in oxygenation, and may help to identify hypoxia in tumours. No intravenous contrast agents are required; however, BOLD effects are very short lived, and require rapid sequences. Some of the changes in signal can be small and are difficult to reproduce.

Functional imaging Functional imaging is defined as characterizing tumours radiologically in terms of their biochemistry or physiology. Such imaging modalities include positron emission tomography (PET), single photon emission computed tomography (SPECT), magnetic resonance spectroscopy (MRS), and dynamic contrast-enhanced MRI and CT. Functional imaging may improve staging of disease by detecting occult carcinoma, or give clearer delineation of areas of previously known tumour (Figure 2). Furthermore, it may provide information on tumour parameters such as blood flow, vascular permeability, proliferation rate and oxygenation. The introduction of functional imaging to radiotherapy planning adds a new concept, termed the biological target volume (BTV) [14].

PET (18)F-Fluoro-2-deoxy-D-glucose (18FDG) is the most commonly used imaging tracer in the diagnosis and staging of HNC. Table 1 summarizes studies showing the sensitivity and specificity in diagnosis of HNC which is superior to CT and MR in assessing lymph nodes, distant metastases and second primaries in a single study [15–23], although the number of false positives makes the specificity suboptimal.

PET-CT PET alone has been of limited value in radiotherapy planning because of its limited spatial resolution 555

K Newbold, M Partridge, G Cook et al

(a)

(b)

(c)

18 FDG PET/CT case study. This patient presented with a right sided level II node containing metastatic squamous cell carcinoma. Conventional imaging with a diagnostic contrast enhanced CT scan showed the right neck node but the primary tumour was occult. (a) The non-contrast enhanced CT component of the PET/CT scan illustrates this (arrow points at node). (b) 18 FDG PET/CT demonstrated the right neck node and but also identified an area of 18FDG uptake in the right tongue base which on biopsy was proven to be the primary site. (c) The planning target volume (PTV) defined in blue shows the target volume planned with conventional data only whereas the PTV defined in green shows the expansion of this when the PET/CT identifies the tongue base tumour and therefore includes the oropharynx and contralateral neck. The PTV increased by 467 cm3 with the addition of data from the PET/CT.

Figure 2.

(4–5 mm for 18FDG-PET) and a lack of anatomical landmarks. Software image co-registration of non-contemporaneous PET images and CT or MRI is possible, but may generate significant matching errors which make it unsatisfactory for radiotherapy planning. Integrated PET/CT scanners produced hardware fused images (Figure 3) which reduces these errors and increases accuracy compared with PET alone in head and neck cancer (96% vs 90%, p 50.03, [15]). The CT component of a standard PET/CT scan is not usually of diagnostic resolution in order to minimize the radiation dose received by the patient. Optimal CT scanning, for example with multislice scanners, are sometimes advised to supplement the PET/CT examination. PET/CT is likely to have a maximum impact in the head and neck region because accurate co-registration enables differentiation between pathological and physiological areas Table 1.

18

FDG-PET in staging head and neck cancer (HNC)

Author

Year

Schmid [16] Dizendorf [17]

2003 2003

Kresnick [18] Hanasano [19] Kau [20] Nowak [21] Wong [22]

2001 1999 1999 1999 1997

556

of 18FDG uptake. Common causes of false positives are Waldeyers ring, salivary glands, brown fat and fast twitch muscles, which have the potential to cloud the diagnostic picture. Using PET, an area of tracer uptake may be seen which is distinct from, or overlap the conventionally defined gross target volume (GTV). In lung cancer planning, the addition of PET led to 26–100% of patients having a change in radiotherapy management when compared with CT plans alone [24]. The changes were not consistent. 15–64% of patients showed an increase in the planning target volume (PTV) and 21–36% had a decrease. Ciernik et al [25] took 39 patients with mixed primaries (12 with HNC) and compared GTV and PTV when localizing using CT alone compared with 18FDGPET/CT. They observed that the GTV in HNC changed in 32% with the PET data (either an increase or decrease)

Number of patients

48 202 all sites 24 146 70 71 54

T stage Sensitivity (%)

N stage Specificity (%)

Sensitivity (%)

Specificity (%)

Management change (%)

8 27 21 50

83

87

67

86 87 80 67

73 94 92 100



Figure 3. Hardware-fused

18

FDGPET/CT showing primary tumour in the left oropharynx. The position of the 18FDG uptake can be seen in relation to the CT derived anatomy in three dimensions.

and the mean PTV change was 20%. Interestingly, the interobserver variability was reduced when the PET data were included compared with CT alone. Nishioka [26] looked at image fusion between 18FDG-PET and MRI/CT for radiotherapy planning in 12 patients with oropharyngeal carcinoma and nine patients with nasopharyngeal carcinoma. They concluded that the fusion was useful in GTV and CTV determination, and enabled sparing of normal tissues. Scarfone et al observed an average increase in GTV of 15% when defined by PET/CT versus CT alone [27]. Table 2 summarizes radiotherapy planning studies with PET imaging.

Other PET tracers for functional imaging DNA precursors, such as 11C or 124I labelled thymidine or deoxyuridine, are incorporated into DNA during repair or S phase and can be imaged to non-invasively identify regions of cell proliferation. 11C-labelled methionine or choline are substrates for protein synthesis and have been tested in prostate cancer [28]. Imaging of tumour hypoxia with tracers is now possible. An example is the group of imidazole containing agents (e.g. 18F-misonidazole) which are bioreductive molecules that accept an electron to form a free radical that is incorporated into the cell constituents under hypoxic conditions. 62Cu-diacetylbis (N-4-methyl-thiosemicarbazone, Cu-ATSM) is a non-imidazole bioreducible radiopharmaceutical which has been used clinically to image hypoxia and to define a potential target for therapy [29].

planning. However, the detection of single photons, rather than two coincident rays, reduces the spatial resolution compared with PET. In view of this, SPECT has remained in the research arena in the head and neck region.

Magnetic resonance spectroscopy (MRS) MRS provides a non-invasive method with which to identify and quantify the presence of specific chemicals within a tissue, e.g. tissue metabolites or the presence of a specific drug. It could feasibly be used to identify chemicals associated with hypoxic tissue. Current limitations of MRS include sensitivity, limited spatial localization and organ motion, but advances in data acquisition techniques have the potential to make spectroscopy an increasingly important clinical tool [30, 31].

Dynamic contrast-enhanced imaging of tumour vascularity, blood flow and permeability Rapid scanning sequences during the administration of contrast agents for CT and MRI have allowed dynamic scans that can impart information regarding the biology of the tumour and its microenvironment, such as blood flow, vascular permeability, hypoxia and pH. Two techniques are currently in clinical research.

SPECT

Dynamic contrast enhanced (DCE) MRI

SPECT produces a three-dimensional tomographic image of the distribution of an injected radioisotope, and therefore is of potential interest for radiotherapy

DCE-MRI involves the acquisition of multiple sequential MRI scans of an area of interest following injection of a contrast agent. This enables study of pharmacokinetics

Table 2.

18

FDG-PET and radiotherapy planning

Author

Year

n

Tumour type

Fusion

Results

Scarfone [27]

2004

6

HNC

Software fusion 18 FDG-PET and CT Hardware fusion 18 FDG-PET/CT Software fusion 18 FDG-PET with MRI/CT

Modified GTV by a mean 15% increase GTV Change-56%, Reduced interobserver variability Useful in GTV and CTV, and normal tissue sparing

Ciernik [25]

2003

39

Various

Nishioka [26]

2002

21

HNC

n, number of patients; HNC, head and neck cancer; NSCLC, non small cell lung cancer; GTV, gross tumour volume; CTV, clinical target volume; PTV, planning target volume.


557

K Newbold, M Partridge, G Cook et al

of para-magnetic contrast agents and provides information on tumour vascularity, blood volume and vessel permeability. MRI sequences can be designed to be sensitive to the initial, largely intravascular phase of contrast delivery [32, 33]. Cooper et al [34] examined the relationship between DCE-MRI parameters and Eppendorf pO2 histographic measurements in 30 patients with cervical carcinoma and found a correlation between maximum enhancement over baseline and rate of enhancement, using T1 weighted sequences, with both median pO2 and proportion of pO2 with values less than 5 mmHg. They concluded that DCE-MRI could be used to measure hypoxia in human tumours in vivo. Hoskin et al [35] examined tumour perfusion in patients with advanced HNC using DCE-MRI (T1 weighted sequences) and found a correlation between local tumour control and maximum tumour enhancement following accelerated radiotherapy.

Perfusion CT Similarly, rapid acquisition of images by spiral or multislice CT, as contrast is given, can estimate tissue perfusion based on the contrast density changes over time [36]. Histological assessment of tumour neovascularization such as microvessel density correlates with contrast enhancement parameters in lung and renal cancer [37, 38]. Hermans et al [39] used this method in patients with HNC who had undergone radical radiotherapy or chemotherapy and concluded that tumour perfusion was an independent predictor of local control, with decreased perfusion levels associated with a higher local failure rate. Possibly these tumours had a reduced blood supply rendering them relatively hypoxic and therefore radioresistant.

Application of functional imaging to head and neck cancer treatment planning Hypoxia targeting Low oxygen levels are associated with reduced apoptotic potential, increased angiogenesis and increased frequency of mutations [40], and are associated with poor local control and survival. Hypoxic radioresistance may be overcome to some extent by increasing tumour oxygenation, hypoxic cell sensitizers, or increasing the radiation dose [41, 42]. Some studies both in vitro and in vivo suggest that a radiation dose 2.5–3 times current dose levels are required to overcome the effects of hypoxia [43]. However, others suggest that more modest dose increases of 1.2–1.5 times the primary dose may result in equivalent tumour control [44]. These doses, in the region of 90–100 Gy, are potentially achievable to areas within tumours with concomitant boost techniques deliverable with IMRT [45]. The spatial and temporal stability of the hypoxic volume during radiotherapy is critical for such approaches and is the subject of ongoing studies. Chronic, or diffusion limited hypoxia is defined by reduced pO2 over hours to days, thought to be due to the distance from a capillary, the oxygen content, the rate of blood flow and the oxygen 558

metabolism of that capillary. Acute, or perfusion limited hypoxia is defined when the variation is over minutes and is thought to be due to intermittent reductions in capillary flow. Dose escalation strategies based on targeting areas of chronic hypoxia are only likely to succeed if this is the dominant cause of treatment failure. PET imaging with 18F-Miso, or CuATSM have been used in this context to quantify hypoxia in head and neck cancer [46, 47]. In a study by Taylor et al [48], BOLD MRI, in conjunction with carbogen-breathing in patients with head and neck carcinoma, suggested improved tissue oxygenation and blood flow; however, difficulties with measurement and reproducibility of BOLD signals have made it impractical for radiotherapy planning.

Validating hypoxia imaging using histological markers Direct measurement of tissue oxygen tension is possible using the Eppendorf polarographic electrode. This technique is not ideal for validating hypoxia because it is invasive, requires accessible tumours and is highly user dependent. 50–150 readings are required per sample, and spatial heterogeneity remains a problem [49]. Sequential readings in assessment of temporal changes are unreliable due to tissue damage following initial formation of tracks [50]. Despite this, DCE-MRI parameters have been correlated to Eppendorf measurements in carcinoma of the cervix [34]. Exogenous and endogenous markers are more promising. Exogenous markers are chemicals that accumulate or are bioreducible in hypoxic conditions, e.g. nitroimidazoles [51]. These retained bioreductive products can be detected by immunohistochemistry, e.g. pimonidazole [52]. Pimonidazole staining has been correlated with outcome in HNC [53]. Endogenous markers are gene products that are up-regulated in the presence of hypoxia. Aebersold et al reported 94% of a cohort of 98 oropharyngeal squamous cell carcinomas overexpressed the transcription factor, hypoxia-inducible factor 1-a (HIF 1-a) [54]. Carbonic anhydrase-9 (CA9) has been shown to have the greatest magnitude of expression in response to hypoxia among a range of 12 genes [55]. CA9 levels rise from 4 h to 24 h of levels of pO2 at 20 mmHg and less. Expression has been concordant with pimonidazole in head and neck cancer [53], and with polarographic (electrode) measurements [56] in cervical cancer. The examination of tumour specimens for hypoxia with the above markers can be used to validate imaging methods performed in the pre-operative period, as long as careful attention is paid to the orientation of the specimen at the time of surgery and histological sections are cut in the same plane as the test images.

What is the likely impact of functional imaging on radiotherapy planning? Functional imaging adds to anatomical imaging for radiotherapy planning in a number of ways. First, functional imaging may alter disease stage, which may have a major impact on disease management. This is The British Journal of Radiology, July 2006


Figure 4. Intensity-modulated radiotherapy (IMRT) plan, illustrating dose boost to a biological target volume (BTV).

Figure 5. Interaction of anatomical and functional imaging modalities and their compatibility with radiotherapy planning (RTP).

more likely to be an issue in diseases such as lung cancer with a high metastatic potential, than with head and neck cancer. Second, functional imaging may improve our localization of the target volume by detection of unexpected tumour extension, or presence of occult locoregional metastases in lymph nodes. This will lead to changes in GTV and CTV definition which may be clinically significant. Finally, functional imaging may determine radioresistant sub-regions within the conventional GTV due to hypoxia or accelerated proliferation. Such biological target volumes may be suitable for radiation dose escalation delivered by techniques such as simultaneous IMRT boosts (Figure 4). These allow increase in total dose and dose-per-fraction, which radiobiologically may be of particular benefit.

Conclusion Several new imaging techniques, both anatomical and functional are currently being evaluated for treatment planning for head and neck cancer (Figure 5). As well as improved conventional target volume definition, new biological target volumes may be generated by these The British Journal of Radiology, July 2006

imaging methods. Careful validation of these imaging methods against histological parameters is urgently required before they can be integrated into clinical treatment planning. Functional imaging will most likely be used in conjunction, rather than as an alternative to, conventional imaging techniques.

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Radiofrequency ablation in pig lungs: in vivo comparison of internally cooled, perfusion and multitined expandable electrodes 1,2

J M LEE, MD, 1,2J K HAN, and 1,2B I CHOI, MD

MD,

1,2

J M CHANG,

MD,

1,2

S Y CHUNG,

MD,

1,2

S H KIM,

MD,

1,2

J Y LEE,

MD

1

Department of Radiology, and Institute of Radiation Medicine, Seoul National University College of Medicine, 28 Yongon-dong, Chongno-gu, Seoul 110-744 and 2Clinical Research Institute, Seoul National University Hospital, Seoul, Korea ABSTRACT. The purpose of this study was to compare the amounts of in vivo coagulation obtained by radiofrequency (RF) ablation in porcine lung, using three types of electrodes. 15 in vivo ablation procedures were performed in the lungs of five pigs using three kinds of currently available RF devices under CT guidance. After placing an electrode in the lung, three ablation zones were created at each of three different regimens: Group A: RF ablation with an internally cooled electrode; Group B: RF ablation with a perfusion electrode, with instillation of 0.9% NaCl solution at a rate of 1.5 ml min21; Group C: RF ablation with a multitined expandable electrode. According to the manufacturer’s recommendations, RF application times were 12 min in group A and 20 min in group B. In group C, RF energy was delivered for 7 min after a mean temperature of 110 ˚C was reached at 5 cm deployment. 36 min after the procedures, contrast-enhanced CT scans were obtained to evaluate the volume of zone of coagulation, and lungs were harvested for gross measurements. After macroscopic and histopathological analyses of 5 mm-thick lung sections, diameters, volumes and variation coefficients of regions of central coagulation were assessed. During RF ablation, the perfusion electrode allowed a larger energy delivery than the internally cooled or the multitined expandable electrodes, i.e. 33.6¡4.7 kJ in group A, 40.0¡8.2 kJ in group B and 23.5¡6.1 kJ in group C (p,0.05). On gross observation, the cut surface of the gross specimen containing RF-induced coagulation showed that the ablated tissue appeared to be a central, firm, dark-brown area surrounded by an irregular outer margin (approximately 3–10 mm thick) of bright red tissue. In vivo studies showed that RF ablation using the perfusion electrode achieved larger coagulation volume than RF ablation using the other electrodes (p,0.05): 7.2¡4.1 cm3 in group A; 16.9¡5.5 cm3 in group B; 7.5¡3.3 cm3 in group C. The corresponding variation coefficients were 0.55, 0.31, and 0.45, respectively. Our study shows that RF ablation using a perfusion electrode achieves a larger coagulation volume with an irregular margin than RF ablation using internally cooled or multitined expandable electrodes in the porcine lung. Lung cancer is one of the major leading causes of cancer-related mortality, with more than 1.2 million new cases of lung cancer being diagnosed worldwide each year [1–3]. In addition, the lung is also the second most frequent site of metastatic disease. Until recently, therapeutic options for primary lung cancers, depending on tumour grading and staging and the presence of comorbidities, included a combination of surgical resection, chemotherapy, and/or XRT. Indeed, surgical resection is the treatment of choice for early-stage non-small cell lung cancer (NSCLC), but unfortunately, patients with NSCLC are frequently poor surgical candidates due to coexistent medical diseases such as chronic obstructive bronchopneumopathy or cardiac disease [4]. Several This study was supported by grant No. 21-2005-021-0 from the Seoul National University Hospital Research Fund.

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Received 12 September 2005 Revised 24 November 2005 Accepted 19 January 2006 DOI: 10.1259/bjr/51844219 ’ 2006 The British Institute of Radiology

studies have documented survival benefits in patients with pulmonary metastases with favourable histologies who received complete resection as compared with unresectable individuals [5, 6]. In patients with lung cancer who are not surgical candidates, the treatment options are primarily XRT with or without chemotherapy. However, for stage I/II NSCLC, the effectiveness for XRT relative to surgery remains uncertain [7], and a meta-analysis of trials comparing primary treatment with or without chemotherapy showed that chemotherapy provided only a modest benefit [8]. Therefore, less invasive therapies that can accomplish tumour destruction without the use of general anaesthesia may complement, improve, or even replace existing therapies. Radiofrequency (RF) ablation has received much recent attention as a minimally invasive strategy for the The British Journal of Radiology, July 2006

In vivo radiofrequency ablation in pig lungs

treatment of various neoplasms of the liver, kidney and bone [9–12]. Also, a number of recent experimental and clinical studies have demonstrated the feasibility and safety of RF ablation for the treatment of inoperable lung malignancies [13–18]. Some preliminary studies have shown that RF ablation enables the successful treatment of relatively small lung malignancies with a high rate of complete response and acceptable morbidity [15–18]. However, others have shown limitations in achieving complete necrosis in large tumours measuring 3 cm or more in diameter [19–21]. Lee et al [20] treated 32 lesions by RF ablation and achieved complete necrosis in 100% of tumours smaller than 3 cm in diameter, and in 23% of larger tumours. To successfully ablate malignant tumours, it is essential to ensure the coagulation of the entire targeted volume with as few complications as possible. Indeed, the acquisition of a large volume of ablation in a safe manner is of paramount importance if RF is to be accepted as routine form of intervention. Previous clinical studies on lung RF ablation have examined internally cooled needle electrodes [18, 20] and multitined expandable electrodes [19, 21]. Currently in Korea, Valleylab, BerchtoldTM, and RITAH RF systems have been used (personal communication with Drs KY Jin, YK Kim, and GS Jung) for the treatment of primary and secondary lung cancers. Recently, Lee et al [22, 23] demonstrated improved RF ablation efficacy using saline infusion to induce a large volume of coagulation both ex vivo and in vivo rabbit lungs. However, no study has compared the in vivo efficacies of different types of electrodes in the lung using large animals. The purpose of our study was to compare in vivo coagulation obtained with currently available RF ablation devices, namely an internally cooled needle electrode, a multitined expandable electrode and a perfusion electrode.

Materials and methods Animals and preparations The experimental protocol was approved by the Animal Use and Care Administrative Advisory Committee of our institution. All experiments were performed according to a protocol approved by the local institutional committee on animals, in accordance with the general guidelines issued by the National Institute of Health for the care of laboratory animals. Five female farm pigs were used in this study (weight range 30– 40 kg). The animals were fasted overnight, but had free access to water before the experiments. Each of the five pigs was anaesthetised using an intramuscular injection of 50 mg kg21 of ketamine hydrochloride (Ketamine; Yuhan, Seoul, Korea) and 5 mg kg21 of Xylazine (Rumpun, Bayer Korea), and prepared for RF ablation. Booster injections of up to half of the initial dose were administered as needed. Ringer’s lactate solution was continuously infused during the experiment (500 ml h21). Endotracheal intubation was performed and anaesthesia was maintained with inhaled enfluorane (Gerolan; Choongwae Pharma Corporation, Seoul, Korea). Mechanical ventilation was used throughout the procedure. The British Journal of Radiology, July 2006

Cardiac and respiratory parameters were monitored throughout the procedures. Each animal’s lateral hindquarters were shaved bilaterally, and two 8 cm 6 12 cm wire-mesh grounding-pads coated with conductive gel were placed on each hind limb. For the RF ablation procedures, animals were placed in the supine position on the CT scanner. Animals were euthanized approximately 1–2 h after the final RF ablation procedure with a pentobarbital overdose of 60 mg kg21. The lungs were then removed.

Study design To minimize potential variations in the RF ablation procedures, all procedures were performed by consensus between two radiologists who had extensive routine experience with all three systems. Electrodes were placed under the same experimental conditions with CT guidance (Somatom plus 4 scanner or Sensation 16; Siemens Medical Solutions, Forchheim, Germany). CT enabled an electrode tip to be positioned at least 2 cm from the pleura, from a previous ablation site in the same lung, and from large vessels (. 3 mm in diameter). RF ablation was performed using one of the three RF systems: Group A: a 200 W generator (CC3: Valleylab TM ) and an internally cooled electrode; Group B: a 60 W generator (HiTT 106 BerchtoldH, Tuttlingen, Germany) and a perfusion electrode; Group C: a 150 W generator (1500 TM model; RITAH medical Systems; Mountain View, CA) and a multitined expandable electrode. The RF systems, lung lobes, position and the order in which each ablation procedure was performed were randomly assigned. A new electrode was used for each ablation session. All settings were performed according to manufacturer’s recommendations. The automated control mechanisms functioned by measuring the total impedance between the electrodes for the BerchtoldH, and ValleylabTM systems and on the temperature of the electrode tip for the RITAH system. Applied current, power output and impedance were continuously monitored using a generator system during RF ablation and were recorded. The technical aspects of the RF ablation, including impedance and wattage changes, and the dimensions of the RF-coagulated area for each system were compared.

RF devices, ablation protocols and procedures Group A (ValleylabTM system) A 480 kHz generator (CC3; ValleylabTM) capable of a maximum power of 200 W was used with a 1.6 mmdiameter internally cooled electrode (single cool-tip needle), which has a 3 cm-long active distal region (Figure 1a). Electrode cooling was ensured by the peristaltic perfusion of chilled saline using a peristaltic pump (PE-PM; ValleylabTM), which allowed the electrode to maintain a tip temperature of below 25 ˚C during RF delivery. This RF system requires the applications of four neural pads. The circuitry incorporated into the generator allowed continuous monitoring of the impedance between the active electrode and the grounding 563

J M Lee, J K Han, J M Chang et al

Figure 1. Different electrodes used for monopolar radiofrequency ablation. (a) ValleylabTM internally cooled electrode with a 30 mm long exposed tip (Cool-tip). (b) BerchtoldH perfusion electrode with a 15 mm long exposed tip and side holes. (c) RITAH multitined expandable electrode (Starburst XL) with nine curved tines.

pads. RF current was passed for 12 min at a maximum generator setting for the impedance control method. This method allows the maximum power to be delivered until impedance rises to 10 V above the baseline value. At this point, the current is switched off automatically to avoid a further local increase in temperature, which would result in tissue charring. 15 s later, the current is automatically switched on again, thus being referred to as the pulsed RF technique, which increases RF ablation area [24].

Group B (BerchtoldH system) This perfusion RF system is supported by a 375 kHz generator (HiTT 106; BerchtoldH) capable of 60 W at maximum through a 1.7 mm-diameter monopolar electrode with an active tip length of 1.5 cm. The electrode used for this saline-enhanced technique was double walled at its distal part, and the inner wall had small perfusion holes (Figure 1b). A 0.9% saline solution was used as the perfusion liquid. This flowed through the hollow shaft of the electrode and permeated through the perfusion holes into the space between the inner and 564

outer walls of the needle tip. Continuous interstitial perfusion of saline was started at 30 s at 90 ml h21 prior to RF application, and was maintained during RF ablation using a digitally controlled syringe pump (Pilot C; Frensenius Vial, Brezins, France). RF current was applied for 20 min at 40 W, in accordance with the manufacturer’s recommendations. The associated control mechanism stabilized RF power within moderate impedance changes of between 100 V and 350 V. When the impedance exceeded 900 V, an additional saline bolus was administered [25].

Group C (RITAH system) Model 1500TM (RITAH Medical systems) utilizes a 150 W generator operating at 460 kHz. In this RF system, the expandable electrode (Starburst XL; Rita Medical Systems) consists of an insulated outer needle with a diameter of 2.2 mm that houses nine deployable curved tines, which have a maximum diameter of 5 cm when fully expanded (Figure 1c). The tips of five of these tines contain a thermocouple and allow temperature monitoring during The British Journal of Radiology, July 2006


the ablation procedure. Electrodes were progressively extended deeper into the lung parenchyma with temperature monitoring, and power was controlled according to average temperature. Tines were first deployed at 2 cm with a pre-selected target temperature of 80 ˚C, then advanced to 3 cm with a target temperature of 105˚C and, finally, extended to 4–5 cm with a target temperature of 110˚C. Target temperatures were maintained for 7 min and then post-ablation temperatures were monitored [26].

Imaging follow-up A multirow detector CT (Sensation 16, Siemens Medical Solutions) was used to monitor ablations at 30–60 min after RF ablation. Axial CT scans were obtained using a 0.75 mm detector collimation, a reconstruction increment of 3 mm and a 1.0 pitch, and included both lungs, before and after injecting 70 ml of contrast medium (Ultravist 370H; Schering Korea, Seoul, Korea). Contrast medium was injected at a rate of 2 ml s21 through an ear vein; post-contrast CT scans were obtained at 60 s after contrast administration and CT images were reconstructed at an interval of 3 mm in the axial plane and at an interval of 1 mm to obtain a high-quality data set for multiplanar reconstruction images. The thin section data set was forwarded to a PC containing dedicated 3D software (Rapidia, INFINITT, Seoul, Korea), and data was reconstructed into 3 mm-thick coronal and sagittal slices.

Assessment of coagulation zone (imaging and pathological studies) Pigs were euthanized after obtaining CT images. Once harvested, lungs were serially sectioned at 5 mm intervals along the axial plane. The histopathological study included staining for mitochondrial enzyme activity, which was performed by incubating thin representative tissue sections for 30 min in 2% 2,3,5,-triphenyl tetrazolium chloride (TTC; Sigma, St Louis, MO), at 20–25 ˚C. This test is a vitalline stain for mitochondrial enzyme activity [27] and can be used to determine irreversible cellular injury during the early stages of RF-induced necrosis [28]. As the unstained area of an RF-induced coagulation has been shown to correspond to the zone of necrosis [28], two observers measured axial diameters along the axis of the electrodes (Dax) and transverse diameters (Dtr) of ablation areas in the axial plane before reaching consensus. The number of slices containing an RF ablated region determined the vertical diameter (Dv). In addition, the slices were photographed using a digital camera (Canon EOS 300D; Canon Inc., Tokyo, Japan), and images were saved to image management software (PhotoShop; Adobe, San Jose, CA). Area analysis was performed on a computer equipped with NIH Image J software (National Institutes of Health; http://rsb.info. nih.gov/ij/) [29]. The area of coagulation, on each slice was calculated using this computer program, and volumes were calculated by multiplying areas by slice thickness and summed to obtain total lesion volumes. The shapes of RF-induced coagulations were characterized using the ratio between the long axis diameter The British Journal of Radiology, July 2006

and the short-axis diameter: Dax/Dtr. Thus, a ratio near 1 indicates a near spherical shape. In addition, volume variations in each group were determined using coefficients of variation, calculated as follows: standard deviation of the ablation volume/mean value of the ablation volume. The closer this ratio is to 0, the more reproducible the coagulation [26]. The RF-induced ablated regions of representative cases in each group were fixed in 10% formalin for routine histological processing, and processed by paraffin sectioning and haematoxylin-eosin staining for light microscopic studies. On contrast-enhanced CT images, the diameters of hyperattenuated ablation areas in treated lungs were measured on axial images and on sagittal reconstructed images using a dedicated 3D program (Rapidia: INFINITT, Seoul, Korea). The Dax and Dtr values of hyperattenuated coagulation were measured in axial slices showing maximum ablation zone dimension, and Dv was measured in sagittal reconstruction images showing maximum ablation zone dimension. Using the Image J software (http://rsb.info.nih.gov/ij/) [29], the area of coagulation, on each CT image was calculated, and volumes were calculated by multiplying areas by slice thickness, summed to obtain total lesion volumes.

Statistical analysis The dimensions of thermal ablation areas and the technical parameters such as accumulated RF energy of the three groups were averaged for each group and compared using one-way analysis of variance (ANOVA) test. In comparisons between groups, the Bonferroni multiple comparison test was applied. The volume of the hyperattenuated area seen at CT was correlated with the volume of central white area of the RF-induced coagulation at pathological specimen, and then the degree of correlation between them analysed using Pearson’s correlation coefficient. Values are expressed as means¡SD. For all statistical analyses, a p-value of ,0.05 was considered significant. Statistics were performed using the Instat program (GraphPad Software, Inc., San Diego, CA).

Results Evaluation of RF methods All pigs tolerated the RF ablation procedures well. In terms of the duration of energy application, the mean duration of RF energy delivery was 12 min in Group A, 20 min in Group B and 23¡1.7 min in Group C (p,0.05). Impedance values decreased linearly during the procedure in group B. However, in Group C, impedance values gradually increased over the first 10 min and fluctuated between 150 V and 200 V over the final 10 min. In Group A, the impedance intermittently increased and activated the pulsed RF algorithm (Figure 2). In Group B, we were able to deliver 40 W and keep the impedance below 150 V. As a consequence, RF ablation using the perfusion electrode (Group B) allowed larger energy delivery than RF ablation using the internally cooled (Group A) or multitined expandable (Group C) electrodes, i.e. 33.6¡4.7 kJ in 565


(a)

(b) Figure 2. Graphic depiction of the electrical parameters during radiofrequency ablation. (a) Tissue impedance (lower row), radiofrequency (RF) current (middle row), and power changes (upper row) during radiofrequency ablation using an internally cooled electrode. (b) Power and impedance changes (lower row), and tissue temperature changes (upper row) during radiofrequency ablation using the multitined expandable electrode.

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Group A, 40.0¡8.2 kJ in Group B and 23.5¡6.1 kJ in Group C (p,0.05).

CT and gross examination of the RF-induced coagulation area On gross observation, the cut surface of gross specimens containing RF-induced coagulation showed that ablated tissue had a two layered structure presenting as a central, firm, dark-brown area surrounded by an outer, irregular margin (approximately 3–10 mm thick) of bright red tissue (Figure 3). When the specimens were fixed with 2% 2,3,5-TTC to determine if mitochondrial function was irreversibly damaged after RF ablation, the central dark-brown area did not show any staining with TTC (no enzyme activity), but the outer bright red zone showed inhomogeneous staining with TTC in red colour (admixture of stained and not-stained areas). The viable lung tissue outside of this area also showed homogeneous staining in red with TTC (positive enzyme activity). The lung tissue on CT, RF-ablated regions appeared as irregularly shaped hyperattenuated areas, which agreed with gross examination findings. In some cases, RF-induced hyperattenuated areas consisted of heterogeneous central zones containing air and surrounded by a broad hyperattenuated region (Figure 3).

Evaluation of volumes of coagulation using gross specimens and CT images The mean Dax’s of the RF induced central white zone (unstained area with TTC) in axial planes measured in gross specimens of the three groups, were 1.7¡0.4 cm in Group A, and 2.7¡0.5 cm in Group B, and 1.9¡0.5 cm in Group C (p50.018) (Table 1). The corresponding mean Dtrs were 2.6¡0.6 cm, 3.5¡0.8 cm and 2.9¡1.0 cm (p.0.05), and the mean Dvs of ablated regions were 3.1¡0.5 cm, 3.6¡0.5 cm and 2.9¡1.0 cm (p.0.05) (Figure 3). Furthermore, the volumes of coagulation areas in the three groups were 7.2¡4.1 cm3 in Group A, 16.9¡5.5 cm3 in Group B and 7.5¡3.3 cm3 in Group C. The differences were significant between Groups A and B (p50.002), and between Groups B and C (p50.018). The coefficients of variability for internally

cooled, perfusion and multitined expandable electrodes were 0.55, 0.31, and 0.45, respectively (p.0.05). Dax/Dt ratios of coagulated areas produced with each RF system were 0.7¡0.1 in Group A, 0.78¡0.2 in Group B and 0.74¡0.4 in Group C. Although there was no significant difference in Dax/Dt ratios between the groups (p.0.05), the perfusion electrode created more irregularly bordered coagulations than the internally cooled or multitined expandable electrodes. On CT, the volumes of RF-induced hyperattenuated areas in the three groups were 9.2¡4.5 cm3 in Group A, 20.4¡6.3 cm3 in Group B and 10.2¡4.0 cm3 in Group C, respectively. The volume of RF-induced hyperattenuated areas in Group B was larger than those of Groups A or C (p,0.05). The volumes of RF-induced hyperattenuated areas on CT were larger than the volumes of RFinduced coagulation (central unstained area of the gross specimen with TTC staining) measured on the gross specimens, but there was a significant correlation between them (p,0.05).

Histopathological results Histological specimens showed central tissue loss at the electrode insertion tract and in the inner ablation zone; lung structures seemed to remain, but parenchymal cells showed eosinophilic cytoplasm with pyknotic nuclei and alveolar exudates, which correspond to the known early changes of coagulation necrosis [30]. The outer ablation zone showed haemorrhagic congestion, which was accompanied by neutrophil infiltration, and pulmonary alveoli were filled with exudates. In this zone, there was an admixture of abnormal cells with pyknotic nuclei and eosinophilic cytoplasm and normal looking parenchymal cells. Beyond this area, lung tissue showed normal histological findings, except acute inflammatory cell infiltrate.

Discussion The efficacy and safety of percutaneous RF ablation have been firmly established for the treatment of hepatic malignancies [9, 10]. However, this is not true in the lung [18, 19]. The most promising application for lung RF ablation is in the treatment and local control of primary

Table 1. Measured values of RF-induced coagulation necrosis in gross specimens in the three groups Coagulation

Group A (12 min RFA)

Group B (20 min RFA)

Group C (23 min RFA)

P-value

Dax (cm) Dtr (cm) Dv (cm) Volume (cm3) Ratio of Dax/Dtr Coefficient of variation Mean applied energy (kJ)

1.7¡0.4 2.6¡0.6 3.1¡0.5 7.2¡4.1 0.7¡0.1 0.55 33.6¡4.7

2.7¡0.5 3.5¡0.8 3.6¡0.5 16.9¡0.5 0.78¡0.2 0.31 40.0¡8.2

1.9¡0.5 2.9¡1.0 2.9¡1.0 7.5¡3.3 0.74¡0.4 0.45 23.5¡6.1

p p p p p p p

, . . , . . ,

0.05+ 0.05 0.05 0.05+ 0.05 0.05 0.01*

Dax, the axial diameter of the central coagulation along the axis of the electrodes; Dtr, transverse diameter of the central coagulation in the axial plane; Dv, vertical diameter of the central coagulation. Group A, RF ablation using an internally cooled electrode; Group B, RF ablation using a perfusion electrode; Group C, RF ablation using a multitined expandable electrode. +Differences between groups A and B, and groups B and C were significant (p , 0.05). *Differences between groups A and C, and groups B and C were significant (p,0.05).


567


(a)

(b)

(c)

(d)

(e)

568


In vivo radiofrequency ablation in pig lungs Figure 3. Contrast-enhanced CT scans and photographs of lung treated by radiofrequency (RF) ablation using an internally cooled electrode and a multitined expandable electrode in a pig model. (a) Contrast-enhanced axial CT scan obtained immediately after RF ablation showing an RF-induced hyperattenuated region (arrows) in both lungs. Note that the hyperattenuated region in the right lung treated with an internally cooled electrode is similar to that in left lung treated with a multitined expandable electrode. Photographs of gross specimen containing RF-induced coagulation areas created with (b) an internally cooled electrode and (c) a multitined expandable electrode. Ablated tissues appeared as a central, firm, dark-brown area (asterisk) surrounded by an outer, irregular margin (arrows) of bright red tissue. (d) Contrast-enhanced coronal reconstruction image obtained immediately after RF ablation using a perfusion electrode showing RF-induced hyperattenuated regions in the left lung (arrows). Note that the hyperattenuated region in the left lung treated with a perfusion electrode is larger than those produced using an internally cooled electrode (arrowheads) in the right lung. (e) Photographs of a gross specimen containing an RF-induced coagulation area created with a perfusion electrode. Note that the short- and long-axis diameters of the coagulation area were larger in Group B using a perfusion electrode than in the other Groups (a and c).

NSCLC. Published studies [16–21] indicate that, although RF ablation has been offered to a heterogeneous cohort of lung cancer patients, initial results are encouraging. According to previous studies, significant differences are apparent in terms of the complete tumour necrosis rates of tumours smaller and larger than 3 cm in diameter [20, 21]. This limitation of lung RF ablation is primarily due to the fact that, when using a single electrode device in monopolar mode, it is limited by the precipitous drop in current density that occurs with distance from the energy source, which makes the periphery of the RF lesion particularly prone to vascular cooling [10, 30]. To circumvent these problems, strategies to increase the dimension of RF-induced ablation zones are needed, and several researchers have suggested that saline-enhanced RF ablation using a perfusion electrode can expand RF-induced coagulation [30–32]. In our study, RF ablation according to manufacturer’s recommendations leads to larger volumes of coagulation using the perfusion electrode than using multitined or internally cooled electrodes. These results confirm those of previous studies [22, 23], i.e. that RF ablation using perfusion electrodes produces significantly larger coagulation areas than internally cooled or multitined electrodes, when RF ablation is performed according to manufacturer’s recommendations. In our study, the high efficacy of the perfusion electrode in creating a large volume coagulation could be attributed to the low impedance kept during the ablation procedure, and high energy delivery during the RF ablation. This phenomenon could be explained by the presence of highly ionic saline around the electrode with continuous saline infusion during the RF procedure, which improves tissue conductivity [30, 31]. In addition, the larger volumes of coagulation obtained with this device are probably the result of the effects of heated fluid and its higher thermal conductivity [22]. The shape of a coagulated area is at least as important as the coagulation volume because the ablation of a tumour requires that the entire tumour, and a safety margin of grossly normal tissue, are encompassed by the ablation. Thus, the creation of large but complex asymmetric coagulation shapes does not reflect the effectiveness of a device. In clinical practice, coagulation shape is determined by the configuration of the RF electrode, the location of the tumour, tumour consistency, perfusion-mediated cooling effects and, in the case of open-perfusion devices, the saline distribution [26]. One potential disadvantage of RF ablation with saline infusion is the possibility of irregular zones of The British Journal of Radiology, July 2006

coagulation, due to uneven distribution of saline and a higher complication rate [32–35]. According to a previous study of RF ablation using a perfusion electrode for treating human pulmonary tumour by Kim et al [34], hypertonic saline-enhanced RF ablation was powerful and efficient in local ablation, but it was difficult to predict the exact extent of ablation. In our study, although there was no significant difference between the Dax/Dtr ratio of RF-induced coagulation areas for the three kinds of electrodes, the perfusion electrode created rather irregular bordered coagulations compared with the internally cooled or multitined expandable electrodes. Given that prospective surgical data demonstrate 3 and 2.4 fold increases in local-regional recurrence rates for local wedge resection and segmental resection, respectively, compared with lobectomy, RF ablation alone for the treatment of primary lung cancer may not be validated [36]. However, for tumours under 2 cm in diameter (stage IA NSCLC), a recent study that compared limited resection (segmentectomy) with lymph node assessment versus lobectomy showed equivalent 5-year survival and local recurrence rates [37]. Recent studies on lung RF ablation [18–20] have led to the opinion that, for small tumours (,3 cm), RF ablation might provide a viable alternative to surgical resection for local disease control, especially in the nonsurgical patient cohort. RF ablation can also be used in conjunction with other treatment modalities. In addition to the previously mentioned combination with XRT, there are ongoing studies on combined RF ablation and brachytherapy in patients with either metastatic lung malignancies or a history of prior treatment that precludes additional external beam radiotherapy [17]. The rationale involves the enhancement of local control by magnifying the cytoreductive and radiation effect by destroying the central hypoxic area of the target tumour. As demonstrated in this study, the larger volume of coagulation created by RF ablation using a perfusion electrode may increase the clinical utility of RF ablation therapy by allowing the successful treatment of larger lung tumours, or by reducing the number of sessions needed to treat a given tumour. In our study, the measured volume of RF-induced hyperattenuated region on CT was greater than that of the RF-induced coagulation at gross specimen. Differences between the measured volumes of RFinduced hyperattenuation on CT and on gross specimens were attributed to the fact that the hyperattenuated area on CT corresponded to a region including both central 569


and peripheral discoloured zones of ablation area on gross specimen; but on the gross specimens with TTC staining, the only central whitish area showed no enzyme activity (no staining). Several limitations of this study must be addressed. First, because of the small sample size, interpretations are limited. Second, differences between the results obtained for the three RF devices are valid in healthy lungs, but not for lung tumours. Therefore, the extent to which our findings reflect the clinical situation is limited. Third, the duration of RF energy application was different between the groups. It would be fair to compare the same duration of energy application, because the volume of coagulation using internally cooled electrodes could be increased if RF energy is applied for more than 12 min. However, in our study, all RF ablation procedures were performed according to manufacturer’s recommendations as in previous studies of porcine liver [25, 26]. Fourth, although there are modifications of expandable electrodes and internally cooled electrodes (cluster) available, we did not include those electrodes in the present study. Given that the developmental speed of RF technology is rapid, this study represents a snapshot in time because further refinements and improvements of current techniques will undoubtedly increase the effectiveness and further expand the role of RF ablation. In addition, we evaluated the geometry of RF-induced coagulation by the two-dimensional measurements. Three-dimensional virtual modelling of RF-induced coagulation in the gross specimens would be beneficial for the evaluation of coagulation geometry. However, given the explanted lung collapse, its value may not be as good as in the liver. Finally, large volumes of coagulation may not always be beneficial or desirable. In certain circumstances, coagulation extending beyond the tumour boundaries could be detrimental if surrounding structures are damaged or if insufficient tissue is preserved to permit normal organ function. In summary, RF ablation according the manufacturer’s recommendations leads to larger volumes of coagulation using perfusion electrodes. Based on our study results, we believe that large lung tumours could be treated by RF ablation using a perfusion electrode, with longer energy application, more effectively than with other electrode types.

8.

9.

10. 11.

12.

13.

14.

15.

16.

17.

18.

19.

20.

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In vivo radiofrequency ablation in pig lungs 25. Denys AL, De Baere T, Kuoch V, Dupas B, Chevallier P, Madoff DC, et al. Radio-frequency tissue ablation of the liver: in vivo and ex vivo experiments with four different systems. Eur Radiol 2003;13:2346–52. 26. Pereira PL, Trubenbach J, Schenk M, Subke J, Kroeber S, Schaefer I, et al. Radiofrequency ablation: in vivo comparison of four commercially available devices in pig livers. Radiology 2004;232:482–90. 27. Goldlust EJ, Placzynski RP, He YY, Hsu CY, Coldberg MP. Automated measurement of infarct size with scanned images of triphenyltetrazolium chloride-stained rat brains. Stroke 1996;27:1657–62. 28. Goldberg SN, Gazelle GS, Compton CC, Mueller PR, Tanabe KK. Treatment of intrahepatic malignancy with radiofrequency ablation: radiologic-pathologic correlation. Cancer 2000;88:2452–63. 29. ImageJ. http://rsb.info.nih.gov/ij/download.html/ [Accessed 27 March 2006]. 30. Goldberg SN. Radiofrequency tumor ablation: principles and techniques. Eur J Ultrasound 2001;13:129–47. 31. Schmidt D, Trubenbach J, Brieger J, Koenig C, Putzhammer H, Duda SH, et al. Automated saline enhanced radiofrequency thermal ablation: initial results in ex vivo bovine livers. AJR Am J Roentgenol 2003;180:163–5.


32. Lee JM, Han JK, Choi SH, Kim SH, Lee JY, Shin KS, et al. Comparison of renal ablation with monopolar radiofrequency and hypertonic-saline-augmented bipolar radiofrequency: in vitro and in vivo experimental studies. AJR Am J Roentgenol 2005;184:897–905. 33. Boehm T, Malich A, Reichenbach JR, Fleck M, Kaiser WA. Percutaneous radiofrequency (RF) thermal ablation of rabbit tumors embedded in fat: a model for RF ablation of breast tumors. Invest Radiol 2001;36:480–6. 34. Kim TS, Lim HK, Kim H. Excessive hyperthermic necrosis of a pulmonary lobe after hypertonic saline-enhanced monopolar radiofrequency ablation. Cardiovasc Intervent Radiol 2005;29:160–3. 35. Gillams AR, Lees WR. CT mapping of the distribution of saline during radiofrequency ablation with perfusion electrodes. Cardiovasc Intervent Radiol 2005;28:476–80. 36. Kodama K, Doi O, Higashiyama M, Yokouchi H. Intentional limited resection for selected patients with T1 N0 M0 non-small-cell lung cancer: a single-institution study. J Thorac Cardiovasc Surg 1997;14:347–53. 37. Okada M, Yoshikawa K, Hatta T, Tsubota N. Is segmentectomy with lymph node assessment an alternative to lobectomy for non-small cell lung cancer of 2 cm or smaller? Ann Thorac Surg 2002;71:956–61.

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PET/CT detects abdominal wall and port site metastases of colorectal carcinoma 1

E GOSHEN,

MD,

2

T DAVIDSON,

MD,

3

D ADERKA,

MD

and 1S T ZWAS,

MD

Departments of 1Nuclear Medicine, 2Radiology and 3Oncology, The Chaim Sheba Medical Center, Tel-Hashomer and Sackler School of Medicine, Tel Aviv University, Israel ABSTRACT. Abdominal wall metastases from colorectal cancer (CRC) may be resected with curative results. Such lesions, often indicators of additional intra-abdominal lesions, may appear in surgical scars, stomas and port site metastases after laparoscopeassisted surgery (LAS). Post-operative changes, primarily surgical scars, alter local physical findings making early detection of small lesions challenging. The purpose of this study was to retrospectively evaluate the contribution of PET/CT to the diagnosis of recurrent colorectal cancer in the post-operative abdominal wall. 120 patients were referred for PET/CT with suspected recurrent CRC based on clinical, radiological or laboratory findings. All underwent whole body PET/CT imaging. 12 of these 120 (10%), were found to have abdominal wall lesions. A total of 16 abdominal wall lesions were detected, located to surgical scars, stomas, drain and laparoscope ports. Additional findings on PET/CT in this group included liver metastases, intra-abdominal lesions and retroperitoneal lymph node involvement. In general, the patients in this small group were young with high grade tumours presenting in advanced stages. In conclusion, PET/CT appears to be a sensitive tool for the diagnosis of abdominal wall recurrence of CRC. The accuracy of localization afforded by the fused functional and anatomic images makes PET/CT a likely tool for diagnosing abdominal wall lesions, including port site metastases of other aetiologies.

Disease recurrence in the abdominal wall from a primary colorectal cancer is a poorly studied and little understood phenomenon. Aggressive resection of disease restricted to the abdominal wall and associated adherent viscera can result in local disease control with little morbidity and no mortality [1], yet abdominal wall metastases are often indicators of recurrent intra-abdominal cancer. Certainly, the timely diagnoses of abdominal wall lesions and possible concomitant intra-abdominal disease are crucial to optimize patient management. Positron emission tomography (PET) with 18F-FDG allows functional imaging of malignant tissue. The welldocumented mechanism of FDG uptake is based on increased glucose metabolism and increased expression of glucose transporters in malignant cells as compared with normal tissue. Thus, the metabolic rate of suspicious lesions as determined from PET can be used to define the presence and extent of active disease. While the sensitivity of dedicated PET has consistently been reported to be in the range of 85–100% for the detection of recurrent colorectal lesions the lack of anatomical detail affects image interpretation and remains a major limitation of PET. The combined PET/CT technique now helps overcome this drawback by providing fused images of functional PET and anatomic CT studies. The almost synchronous image acquisition and exact coregistration of anatomical and metabolic data improves the anatomic localization of PET abnormalities and reduces the number of equivocal PET interpretations [2–4]. PET/CT imaging has been found to increases the accuracy and certainty of locating lesions in colorectal 572

Received 10 August 2005 Revised 1 November 2005 Accepted 24 November 2005 DOI: 10.1259/bjr/25287790 ’ 2006 The British Institute of Radiology

cancer with a consequent improvement in staging and restaging accuracy from 78% to 89% [5]. This report describes the use of PET/CT in the detection of abdominal wall lesions, including port site metastases from colorectal carcinoma.

Patients and methods This was a retrospective evaluation of positive PET scans. Patients referred for PET/CT with a diagnosis of colorectal cancer were listed in a prospective database beginning 01/2004. Review of this database (10/2004) indicated 120 patients, 12 of whom (3 female, 9 male, aged 42–73 years, average 57 years) were found to have recurrent disease in the abdominal wall. The study population consisted of these 12 patients, who underwent 13 PET/CT scans. Primary disease was located in caecum (n53), transverse colon (n52), sigma (n54) and rectum (n53). All presented initially with advanced disease (all T3 or T4 according to TNM classification), and 10 of the 12 had moderate to poor tumour differentiation (1 was well differentiated, 1 unknown). Only 5 of the 12 had documented initial lymph node involvement. Three patients had intestinal perforations, yet there was no evidence of peritoneal spread at surgery. Three had liver metastases at presentation. All were reported to have clean surgical margins. Of the 12 patients, 8 had originally undergone open laparotomy, and 4 had laparoscope-assisted surgery (LAS). One patient (pt #4) had both LAS and open surgery, and The British Journal of Radiology, July 2006

PET/CT detects abdominal wall and port site metastases of CRC Table 1. Demographic and clinical data Patient sex, age

Primary site

Clinical presentation

Surgery

Tumour differentiation

Stage

Lymph nodes

Other Sites at presentation

1(F, 42)

sigma

open*

mod-poor

T3N2M1

6 of 15

liver

2(F, 67) 3(M, 71) 4(M, 73)

cecum cecum cecum

abdominal abscess vomiting obstruction obstruction

mod-poor well-mod mod

T3N1M0 T3N2M0 T4N2M0

3 of 39 4 of 8

5(M, 56) 6(M, 45) 7(M, 46) 8(M, 73) 9(F, 49) 10(M, 57) 11(F, 59) 12(M, 46)

rectum transverse sigma sigma rectal transverse sigma rectum

LAP open open LAP open* open open* LAP open LAP open open

mod well diff. mod mod NA mod mod mod

T3N0M0 T3N1M0 T4N0M1 T3N0M0 NA T3N1M0 T3N0M0 T3N0M1

0 of 5 3 of 9 0 of 26 0 of 6 NA 1of 4 0 of 46 0 of 4

perforation into bladder anaemia pain pain obstruction pain

CEA

Treatment

De Gramont 31

7.9 liver

5FU-LCV 5FU-LCV Folfiri no Folfiri XRT+UFT no

appendix peritonitis liver

29

5FU-LCV refused Folfiri

*Patients with perforations at surgery.

also underwent two PET scans. The clinical data of these patients are summarized in Table 1. All 12 patients presented at follow-up with clinical, radiological or laboratory findings suggestive of recurrent disease. Four patients were referred for PET/ CT due to palpable masses, one had intestinal obstruction, one had CT findings suggestive of local recurrence and the remaining six were referred for PET/CT due to marker elevation. Patient #4 was referred both times due to palpable abdominal lesions. The interval between initial surgical procedures and PET/CT for suspected recurrence in this group ranged between 4 months and 37 months, the average being 15 months. The indications for scan and intervals from surgery are presented in Table 2, together with PET/CT results.

PET/CT imaging protocol All patients were imaged with a Gemini PET/CT Imaging System (Philips Co., USA) which provides

three-dimensional acquisition, processing and display of CT, PET and PET/CT images, with 6.0 mm PET spatial resolution and a dual slice MX800 EXP CT scanner. Whole body PET/CT imaging was performed in fasting patients (4–6 h) following intravenous administration of 370 MBq (10 mCi) of 18F-FDG. Neither oral nor intravenous contrasts were administered. After a 60 min uptake period, during which patients were instructed to rest silently, images were acquired. First, a CT surview, (30 mA, 120 kVp, FOV 500 mm, length of scan 1.0–1.5 m with a speed of 100 mm s21 and a spatial resolution of 1 mm) was performed from the orbital level to the level of the proximal thigh. This was followed by a dual slice CT (50 mAs per slice, 120 kVp, with a slice thickness of 6.5 mm, length of scan according to result of surview, with a bed speed of 20 mm s21, rotation time of 0.75 s and a pitch of 1.5, and FOV of 600 mm). Finally, the acquisition of PET emission images was performed (2–3 min per bed position of 8.4 cm). The total acquisition time, accumulating between 100 and 150 million useful events, varied between 25 min and 35 min per patient. The CT data were used for attenuation correction of PET emission images.

Table 2. Scan indications, recurrence time and PET/CT findings Pt#

Indication for PET/CT

Post-op interval (months)

Lesion #

Lesion Size (cm)

18

F-FDG uptake intensity

Abdominal wall location

Other recurrence sites

1 2 3 4

CEA CEA Ca-19.9 Palpable

10 9 21 28

5 6

Palpable CT Obstruction

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16

1.3 1.4 2.061.0 1.663.0 1.061.2 2.063.0 1.0 1.761.0 2.063.0 1.3 567 362 1.4 1.0 1.0 1.061.0

+ +++ + +++ +++ +++ +++ +++ +++ + +++ +++ +++ ++ ++ ++

Scar LAP Scar (hernia) Scar Scar LAP Stoma Scar Stoma Scar LAP Scar LAP Drain Drain Drain (hepatic)

None Retroperitoneal LN Intra-abdominal spread Local recurrence * Local recurrence none Local recurrence * Liver none Liver Intra-abdonimal spread Liver * Local Recurrence

7 8 9 10 11

CEA

4 10 29 37 6 27 12 18 10

12

CEA

7

CEA Palpable Palpable

*Same site of recurrence as listed above in previous scan of same patient.


573

E Goshen, T Davidson, D Aderka and S T Zwas

Non-attenuated data was reconstructed after scan acquisition was completed. Reconstruction of attenuation corrected data was executed concurrently. The lower limit of resolution of the PET system used is 0.45 cm.

Image interpretation Both attenuation corrected and non-attenuation corrected scans were coregistered with the CT for interpretation using Syntegra (Version 2.0j, Philips) software. All studies were visually interpreted independently by individual physicians from the two specialties (Nuclear Medicine and Diagnostic Radiology). Studies were interpreted for PET alone and for combined PET/CT data. The low-dose non-enhanced CT data was used only in conjunction with PET data. Comparisons between the CT data and PET regarding contributions in the specific clinical context reported here were not made.

Evaluation of accuracy and statistical analysis Results from combined PET/CT data were validated by comparison with concurrent contrast-enhanced CT, histopathological findings and at least 3 months of clinical follow-up. The contrast-enhanced CT scans used for comparison were not performed in our department, and were provided by the patients. PET/CT findings were interpreted on both lesion and patient levels, and determined to be true-positive (presence of cancer), or false-positive (increased FDG uptake unrelated to cancer). As the patients were selected on the basis of a positive PET result, there can be no meaningful assessment of true or false negatives.

Results A total of 16 abdominal wall lesions were detected in 13 scans performed in 12 patients. Three patients had 2

lesions each (pts 6, 11 and 4 who had 3 sites in 2 scans). The lesions ranged in size from 1.0–567 cm, with most (9/16) measured less than 1.5 cm. Intensity of 18F-FDG uptake was variable, yet the majority of lesions (10/16) had intense 18F-FDG avidity. The least uptake was observed in the smallest lesions. All lesions were noted from PET findings and were subsequently localized anatomically by CT. Seven findings were localized to mid-abdominal surgical scars (1/7 in the site of a hernia), four localized to laparoscope insertion sites, two were in stomas and three in sites of surgical drains (1/3 in a hepatic drain site). The nonenhanced CT findings generally depicted heterogeneous nodular lesions, described as soft tissue masses with ill defined borders located near muscle. In some cases, fat infiltration was observed adjacent to the soft tissue mass, consistent with post-operative changes. No increased uptake of 18F-FDG was observed in any such region of fat infiltration. The abdominal lesions were divided nearly equally between open-surgery scars and drain/LAP sites. The lesions depicted essentially the same functional and anatomical characteristics (demonstrated in Figure 1). There were no suspicious findings on CT which were undetected by PET (no negative PET scans), and no CT findings which were below the resolution of the PET. All abdominal wall lesions were previously unreported. Concomitant disease was present in 9 of 12 patients. The PET/CT findings demonstrated additional sites of active disease in 11 of 13 scans as follows: retroperitoneal lymphadenopathy (n51), intra-abdominal spread (n52), liver metastasis (n53) and local recurrence (n54), including repeated findings twice in pt #4. Findings from PET/CT are summarized in Table 2. All PET/CT findings in this group were interpreted as true-positive, i.e. consistent with presence of cancer. Two abdominal wall lesions (pts #5 and #12) were highly suspicious clinically, while all other occult sites of abdominal wall and local recurrences were validated histologically.

Figure 1. (a) Non enhanced CT demonstrates small irregularity in right anterior abdominal wall. (b) Attenuation-corrected PET image demonstrates focus of increased uptake of 18F-FDG in corresponding site in abdominal wall. The additional focus of intraabdominal uptake on the right is due to physiological uptake in the ureter. (c) Fused PET/CT image yields functional and anatomic information.

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PET/CT detects abdominal wall and port site metastases of CRC

The patient profile from the data of this small group included young age (8/12 below 60 years of age, of whom 5 were in their 50s), advanced disease at presentation (8/12 patients T3; 3/12 patients T4, and 1 unknown), with tumour differentiation being mostly moderate to poor.

Discussion Changes in tumour metabolism precede morphologic changes, and the functional information obtained from PET is therefore essentially independent of tumour location and size. The additional anatomic information provided by the CT contributed to the localization and characterization of lesions, thereby increasing the accuracy of the PET diagnoses. Overall, PET scanning has been reported to be contributory in the initial staging as well as in the evaluation of recurrent colorectal carcinomas. Recent works have reported the successful use of FDG-PET for the assessment of tumour aggressiveness, staging of disease, evaluation of treatment response, and the detection of recurrent disease [6]. While abdominal wall lesions and port site metastases are well documented entities, to the best of our knowledge the use of PET or PET/CT specifically addressing their diagnostic contribution has not been previously described. Abdominal wall recurrences have been reported to occur in approximately 60% of patients with advanced colorectal cancer [7]. The importance of early and accurate diagnosis of such recurrence is not only in the localization of disease, allowing resection with curative intent, but also in the fact that abdominal wall metastases are often indicators of recurrent intra-abdominal cancer [7]. In the small study group evaluated here in fact, all but three patients (9/12 patients, in 11/13 scans) had additional sites of active disease. In three patients the additional sites were previously diagnosed by CT (in two patients referred for evaluation of palpable masses, and one with obstruction on previous CT) whereas previously unsuspected sites of disease were disclosed in another six patients. This finding is particularly striking as it was observed in the asymptomatic patients referred for evaluation due to elevated markers. In this subgroup, the lack of disease documentation on their previous CT was, in fact, the reason for referral to PET. When dividing our population into subgroups according to indication for the scan, this was also the largest subgroup. Abdominal wall lesions, long recognized in surgical scars and stomas, have received renewed attention due to the recognition of port site metastases in patients following laparoscopic colorectal resections [1]. Although beneficial to the patient in the immediate post-operative period, the initial descriptions of port site lesions had put the adequacy of laparoscopic-assisted colectomy for tumour under question [8]. The incidence of port site metastases, however, is now recognized to be close to the incidence of wound metastases after open surgery [9]. Clinical evidence indicates that wound recurrence rates are 0.60% and 0.85% for open and laparoscopic colon cancer operations, respectively [10]. The group investigated here was derived from a larger group of 120 patients with suspected recurrent CRC, yet The British Journal of Radiology, July 2006

no evaluation of the entire population was performed, particularly regarding initial surgical procedure. Interestingly, however, the limited results in this small group indicate the abdominal lesions to be divided nearly equally between open-surgery scars and trocar/ LAP sites. Excluding the two lesions localized to stomas, of the remaining lesions, half (7/16) were in incision sites from open surgery, and half (7/16) in trocar and laparoscope insertion sites. The lesions were essentially the same functionally and anatomically, and differed, in fact, only by location. Increased uptake in stoma sites is frequently observed and well recognized. Physiological intestinal uptake is generally mild, but may be more increased if compounded by inflammation, complicating scan interpretation. Anatomic localization of the uptake by CT, and particularly the definition of soft tissue changes (nodular lesions, or infiltration) surrounding the stoma, allows clearer diagnoses with greater confidence when tumour is suspected. The retrospective evaluation of all postoperative CT images helps to discern physiological findings from pathology, including inflammatory changes (subcutaneous fat infiltration, or fluid collections suggestive of abscess). Two tumours in stoma were demonstrated, and later pathologically proven, in the small group reported here. In both cases (pts 5 and 6) the uptake of 18F-FDG in the stoma was intense, yet could have been confused with inflammation. In these cases the additional anatomic information provided by the CT served to differentiate the findings from physiological uptake and inflammatory changes. The intensely increased 18F-FDG uptake, which was localized to ill-defined soft tissue masses in the abdominal wall surrounding the stoma, allowed for the diagnosis of viable tumour. This is not a report on all CRC patients referred for PET/CT, but rather a retrospective evaluation of 12 patients with abdominal wall lesions. Only two patients presented with stoma lesions. No remarkable findings were noted in stoma sites in the remainder of the group (data not presented). As this is a report on the 12 patients with positive scans, there is also no information regarding the incidence of inflammatory uptake in stoma in the overall CRC population evaluated. This report does not attempt to compare the contribution of PET/CT with that of CT in the context of abdominal wall lesions. Certainly, if abdominal wall lesions are large enough to be detected on conventional CT, and considering the high prevalence of concomitant disease in the presence of abdominal wall metastases, PET/CT may even be non-contributory and not costeffective. Certainly, if the lesions, particularly intraabdominal findings, are discovered (and duly reported) from CT, the PET may be unnecessary as curative options are limited. Our population was comprised of patients who presented at follow-up with clinical, radiological or laboratory findings suggestive of recurrent disease. Four patients were referred for PET/CT for evaluation of palpable masses, one had intestinal obstruction, one had equivocal CT findings suggestive of recurrence. The remaining six were referred for PET/ CT due to marker elevation in the presence of ‘‘normal’’ CT. 575

E Goshen, T Davidson, D Aderka and S T Zwas

Interestingly, while all abdominal wall lesions were at least 1 cm in size, and could have been identified on the concurrent contrast-enhanced CT, this was not the case. Unfortunately, missed diagnoses occurred. The PET consistently elucidated equivocal findings, and yielded ‘‘hot-spots’’ which easily directed attention to underlying lesions. The point which cannot be overstressed is that PET ‘‘shines the light’’ on small, sometimes overlooked, findings. Port sites and open wounds are at equal risk of tumour implantation, with wound implantation affected mostly by the stage of the tumour and the operative techniques [10]. While our group is of limited size, and a complete demographic evaluation of the entire population (120 patients with suspected recurrent CRC) was not performed, a patient profile can be outlined. In general, the patients who developed abdominal wall recurrences were mostly young, and presented initially with advanced disease of moderate–poor differentiation. The general population (120 patients) from which this subgroup was derived includes patients in all age groups referred for evaluation, regardless of the stage of CRC at presentation. As such, this subgroup, with its relative homogeneity, is not representative of the entire population. While none were reported to have intra-abdominal dissemination at surgery, local spread of microscopic disease could not be excluded, particularly in two patients (pts #1 and 11, who presented initially with abscess, and peritonitis, respectively). As expected, the abdominal wall lesions were indistinguishable clinically whether they resulted in patients after open surgery or developed in port sites after laparoscope-assisted procedures. The average time interval between resection of primary tumour and PET/CT diagnosis of recurrence was 15 months in this group (ranging 4–37 months). This is in keeping with anecdotal reports of abdominal wall recurrences. A case of a port site recurrence of colonic adenocarcinoma with diffuse peritoneal carcinomatosis was reported 1 month after laparoscopic-assisted right hemicolectomy, suggesting that intraperitoneal dissemination and tumour implantation on surgical wounds may have been the principal mechanism of recurrence after laparoscopic surgery [9]. In another case reported, a drain-site tumour recurrence was diagnosed 2 years after right colon resection for adenocarcinoma of the ascending colon [11]. Itano et al [12] reported on a patient who developed two subcutaneous metastases at two trocar sites. Multiple lesions in the abdominal wall, as observed in three of our patients (pt # 6, #11 and #4) are, therefore, also in keeping with the literature. While the patient group is too small to allow statistical evaluation, the fact that three out of 12 patients had multiple abdominal wall lesions is an interesting observation which may warrant further investigation and should, perhaps, be taken into consideration during follow-up of patients undergoing such procedures. Of special interest is the lesion detected in the insertion site of a hepatic drain (pt #12). While a similar finding has been previously described, the occurrence is rare. A previous case has been reported in the literature of a colon carcinoma metastasizing to a port site following laparoscopic cholecystectomy has been reported. That 576

case too was of an advanced tumour at the time of diagnosis, in an immunosuppressed patient, who developed clinical recurrence within 3 months of her procedure [13]. Previous reports of abdominal wall metastases after laparoscopic procedures for colorectal cancer have suggested the finding to be anecdotal [12–14]. In our group, however, 12 of 120 (10%) of those initially referred for investigation of suspected recurrent CRC had PET/CT findings consistent with abdominal wall lesions. Excluding the two lesions localized to stomas, of the remaining lesions half (7/16) were in incision sites from open surgery and half (7/16) in trocar and laparoscope insertion sites. This appears to be in keeping with large-scale studies that have shown the actual rate of port site metastasis to be similar to that observed in open surgery [12]. Furthermore, PET/CT disclosed concomitant findings in the vast majority (75%) of our, albeit small, group. PET/CT is frequently indicated for the restaging of disease in CRC patients prior to metastasectomy. As limited disease is considered operable with curative intent, the scan is performed with hopes of demonstrating only limited disease, thereby allowing the planned surgery to proceed. While the presence of additional findings may contraindicate surgery, abdominal wall and port site metastases per se should not serve as contraindications, but, rather, should be considered for resection as well. The combined findings should be taken into consideration when planning treatment and the presence of a resectable abdominal wall lesion should not exclude surgery as an option.

Conclusion Although preventive measures have been proposed [9], it seems likely that abdominal wall lesions, including port site metastases, will continue to present diagnostic challenges. A well-documented complication after laparoscopic resection of intra-abdominal malignancies, port site metastases have been reported in various other clinical settings including breast cancer [15], gallbladder and bile duct cancer [16, 17], renal cell carcinoma [18], pancreatic cancer [19], upper gastrointestinal tract cancers [20] and gynaecological tumours including ovarian, cervical and endometrial cancer [21–23]. The metabolic images provided by PET scanning appear to have a high positive predictive value regarding the detection of abdominal wall metastases from colorectal cancer. The implementation of PET/CT will likely allow detection, characterization and anatomic localization of such occult lesions, as well as the diagnosis of additional concomitant lesions, in other clinical settings as well.

References 1. Koea JB, Lanouette N, Paty PB, Guillem JG, Cohen AM. Abdominal wall recurrence after colorectal resection for cancer. Dis Colon Rectum 2000;43:628–32. 2. Beyer T, Townsend DW, Brun T, Kinahan PE, Charron M, Roddy R, et al. A combined PET/CT scanner for clinical oncology. J Nucl Med 2000;41:1369–79.


PET/CT detects abdominal wall and port site metastases of CRC 3. Schoder H, Erdi YE, Larson SM, Yeung HW. PET/CT: a new imaging technology in nuclear medicine. Eur J Nucl Med Mol Imaging 2003;30:1419–37. 4. Martinelli M, Townsend D, Meltzer C, Villemagne VV. Survey of results of whole body imaging using the PET/CT at the University of Pittsburgh Medical Center PET Facility. Clin Positron Imaging 2000;3:161. 5. Cohade C, Osman M, Leal J, Wahl RL. Direct comparison of (18)F-FDG PET and PET/CT in patients with colorectal carcinoma. J Nucl Med 2003;44:1797–803. 6. Hustinx R. PET imaging in assessing gastrointestinal tumor. Radiol Clin North Am 2004;42:1123–39. 7. Silecchia G, Perrotta N, Giraudo G, Salval M, Parini U, Feliciotti F, et al; For the Italian Registry of Laparoscopic Colorectal Surgery. Abdominal wall recurrences after colorectal resection for cancer: results of the Italian registry of laparoscopic colorectal surgery. Dis Colon Rectum 2002;45:1172–7. 8. Jacquet P, Averbach AM, Jacquet N. Abdominal wall metastasis and peritoneal carcinomatosis after laparoscopic-assisted colectomy for colon cancer. Eur J Surg Oncol 1995;21:568–70. 9. Curet MJ. Port site metastases. Am J Surg 2004;187:705–12. 10. Allardyce RA. Is the port site really at risk? Biology, mechanisms and prevention: a critical view. Aust N Z J Surg 1999;69:479–85. 11. Torzilli G, Cremascoli G, Cattaneo S, Stefanini P, Olivari N. Drain-site tumour recurrence after laparotomy resection for colorectal cancer. Eur J Surg Oncol 1999;25:546–7. 12. Itano O, Watanabe T, Jinno H, Suzuki F, Baba H, Otaka H. Port site metastasis of sigmoid colon cancer after a laparoscopic sigmoidectomy: report of a case. Surg Today 2003;33:379–82. 13. Neuhaus S, Hewett P, Disney A. An unusual case of port site seeding. Surg Endosc 2001;15:896. 14. Lauter DM, Froines EJ. Initial experience with 150 cases of laparoscopic assisted colectomy. Am J Surg 2001;181:398–403.


15. Mylonas I, Janni W, Friese K, Gerber B. Unexpected metastatic lobular carcinoma of the breast with intraabdominal spread and subsequent port-site metastasis after diagnostic laparoscopy for exclusion of ovarian cancer. Gynecol Oncol 2004;95:405–8. 16. Nakagawa S, Tada T, Furukawa H, Abe M, Hatakeyama K. Late-type recurrence at the port site of unexpected gallbladder carcinoma after a laparoscopic cholecystectomy: report of a case. Surg Today 2000;30:853–5. 17. Sakata N, Suzuki M, Shibuya K, Takeda K, Matsuno S. Unexpected bile duct carcinoma presenting with port-site metastasis after laparoscopic cholecystectomy for cholecystolithiasis. Hepatobiliary Pancreat Surg 2002;9:511–4. 18. Iwamura M, Tsumura H, Matsuda D, Kurosaka S, Yoshida K, Baba S. Port site recurrence of renal cell carcinoma following retroperitoneoscopic radical nephrectomy with manual extraction without using entrapment sac or wound protector. J Urol 2004;171:1234–5. 19. Jorgensen JO, McCall JL, Morris DL. Port site seeding after laparoscopic ultrasonographic staging of pancreatic carcinoma. Surgery 1995;117:118–9. 20. Shoup M, Brennan MF, Karpeh MS, Gillern SM, McMahon RL, Conlon KC. Port site metastasis after diagnostic laparoscopy for upper gastrointestinal tract malignancies: an uncommon entity. Ann Surg Oncol 2002;9:632–6. 21. Viala J, Morice P, Pautier P, Castaigne D, Vanel D. CT findings in two cases of port-site metastasis after laparoscopy for ovarian cancer. Eur J Gynaecol Oncol 2002;23: 293–4. 22. Gregor H, Sam CE, Reinthaller A, Joura EA. Port site metastases after laparoscopic lymph node staging of cervical carcinoma. J Am Assoc Gynecol Laparosc 2001;8: 591–3. 23. Muntz HG, Goff BA, Madsen BL, Yon JL. Port-site recurrence after laparoscopic surgery for endometrial carcinoma. Obstet Gynecol 1999;93(5 Pt 2):807–9.

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Quantitative colorectal cancer perfusion measurement by multidetector-row CT: does greater tumour coverage improve measurement reproducibility? 1,2

V GOH, MA, MRCP, FRCR, 1S HALLIGAN, C I BARTRAM, FRCP, FRCS, FRCR

MD, FRCP, FRCR,

1

L GARTNER,

MRCP,

1

P BASSETT,

MSc

and

1

1

Intestinal Imaging Centre, St Mark’s Hospital, Harrow, and 2Paul Strickland Scanner Centre, Mount Vernon Hospital, Northwood, UK

ABSTRACT. The purpose of this study was to determine if greater z-axis tumour coverage improves the reproducibility of quantitative colorectal cancer perfusion measurements using CT. A 65 s perfusion study was acquired following intravenous contrast administration in 10 patients with proven colorectal cancer using a fourdetector row scanner. This was repeated within 48 h using identical technical parameters to allow reproducibility assessment. Quantitative tumour blood volume, blood flow, mean transit time and permeability measurements were determined using commercially available software (Perfusion 3.0; GE Healthcare, Waukesha, WI) for data obtained from a 5 mm z-axis tumour coverage, and from a 20 mm z-axis tumour coverage. Measurement reproducibility was assessed using Bland-Altman statistics, for a 5 mm z-axis tumour coverage, and 20 mm z-axis tumour coverage, respectively. The mean difference (95% limits of agreement) for blood volume, blood flow, mean transit time and permeability were 0.04 (22.50 to +2.43) ml/100 g tissue; +8.80 (250.5 to +68.0) ml/100 g tissue/min; –0.99 (28.19 to +6.20) seconds; and +1.20 (25.42 to +7.83) ml/100 g tissue/min, respectively, for a 5 mm coverage, and 20.04 (22.61 to +2.53) ml/100 g tissue; +7.40 (250.3 to +65.0) ml/100 g tissue/min; 22.46 (212.61 to +7.69) seconds; and 20.23 (28.31 to +7.85) ml/100 g tissue/min, respectively, for a 20 mm coverage, indicating similar levels of agreement. In conclusion, increasing z-axis coverage does not improve reproducibility of quantitative colorectal cancer perfusion measurements.

Perfusion imaging techniques have been advocated increasingly for assessment of tumour response to therapy because antiangiogenic and vascular targeting drugs may not cause tumour shrinkage, despite being therapeutically effective. As a result, conventional response assessment based on changes in tumour size, such as RECIST or WHO criteria [1, 2], may underestimate the effect of these agents [3]. For example, Hurwitz et al reported a 5 month improvement in overall survival in patients with advanced colorectal cancer who had been treated with bevacizumab (Avastin; Genentech, CA), an antivascular endothelial growth factor agent: this was accompanied by a response rate of only 10% using conventional size-based criteria [4]. Tumours are spatially and temporally heterogeneous; at any time point they may recruit 20–85% of the vasculature available to them [5]. Thus functional assessment derived from a single tumour level (or time) This research was supported by a grant from the Royal College of Radiologists, London, UK. Address correspondence to: Professor Steve Halligan, Intestinal Imaging Centre, St Mark’s Hospital, Watford Road, Harrow, Middlesex HA1 3UJ, UK. Current address: Specialist Radiology, Level 2 Podium, University College Hospital, 235 Euston Road, London NW1 2BU, UK.

578

Received 24 August 2005 Revised 7 November 2005 Accepted 24 November 2005 DOI: 10.1259/bjr/18842556 ’ 2006 The British Institute of Radiology

may not reflect perfusion in the tumour as a whole. Greater anatomical coverage is now possible with multidetector row CT scanners and this is likely to increase in the future as ever more detector rows become possible. To date, there has been no attempt to determine if increased tumour coverage results in decreased measurement variability for colorectal cancer. We obtained colorectal cancer perfusion measurements from a single 5 mm axial slice, and from 4 contiguous 5 mm slices, producing a z-axis coverage of 20 mm, to determine if reproducibility is enhanced when a greater volume of data is analysed.

Materials and methods Subjects This prospective study received local ethics committee approval, and written informed consent was obtained from all patients. 10 patients (mean age 67 years, range 39–84 years; 6 male, 4 female) with histologically proven colorectal adenocarcinoma, and attending for a CT staging examination, were recruited prospectively. Of the 10 tumours examined, 5 were located in the rectum The British Journal of Radiology, July 2006

Colorectal perfusion measurement by multidetector-row CT

and 5 in the sigmoid colon. Of these, 2 were T2, 5 were T3, and 3 were T4 tumours (mean size 6.2 cm, range 5– 11.5 cm).

CT scanning Following a 4 h fast, 1000 ml of water-soluble contrast, 2–4% meglumine and sodium diatrizoate (Gastrografin; Bracco, Milan, Italy), was ingested orally 30 min prior to CT scanning in order to opacify the small bowel, as per normal practice in our institution. With the patient lying supine on the scanner table, an 18 G venous cannula was sited in the antecubital fossa. 20 mg of the spasmolytic hyoscine N butylbromide (Buscopan; Boehringer Ingelheim, Ingelheim am Rheim, Germany) was administered intravenously to reduce bowel peristalsis. Patient movement was minimized by placing a restraining band around the abdomen. All patients were scanned using a four-detector row CT scanner (Lightspeed Plus; GE Healthcare Technologies, Waukesha, WI). A non-contrast abdominal-pelvic study was performed initially in order to confirm the location of the known colorectal tumour, using the following parameters: slice thickness/interval 10 mm/5 mm, mode HS/speed 30 mm s21 (pitch 1.5), 120 kV, 180 mA, 0.6 s rotation speed, scan field of view (SFOV) 50 cm, matrix 512 mm6512 mm. The images were then inspected on the CT console by the supervising radiologist, the mid-tumour level identified, the scan location noted and thus used to plan the subsequent dynamic study. The dynamic studies were performed using contrast and image acquisition protocols in accordance with manufacturer’s recommendations. A pump injector (Percupump Touchscreen; EZ-EM, Westbury, NY) was used to inject 100 ml of iopamidol 340 (Niopam 340; Bracco) intravenously at a rate of 5 ml s21. Four contiguous slices collimated to 5 mm each were obtained at 1 s intervals through the mid-point of the tumour using a ‘‘cine mode’’ (120 kV, 60 mA, SFOV 50 cm, matrix 512 mm6512 mm). Data acquisition commenced 5 s following the start of intravenous injection, to obtain baseline non-contrasted images, and lasted for a total duration of 65 s. All patients returned within 48 h of the initial study for a second dynamic study specifically to assess measurement reproducibility. Scans were acquired in an identical fashion to the initial study. In particular, the second non-contrast planning scan was compared with that used for the initial study so that the tumour level examined for each study could be matched. Intravenous spasmolytic and contrast were administered exactly as previously and data was acquired using technical parameters identical to the initial dynamic study.

Image analysis Image analysis was performed by a single radiologist experienced in CT perfusion analysis. All 20 dynamic studies (10 patients; 2 studies each) were analysed on a stand-alone workstation (Advantage 4.1; GE Healthcare Technologies) using commercial software based on The British Journal of Radiology, July 2006

deconvolution analysis (Perfusion 3.0; GE Healthcare Technologies). The initial 65 s dynamic study for each patient was loaded into the software (Body tumour, Perfusion 3.0) and a single 5 mm axial slice that best visualized the tumour was chosen from the four axial slices available. A processing threshold of 0–120 Hounsfield Units was selected so that the subsequent analysis appropriately included soft tissue, both unenhanced and enhanced. The arterial input was determined by manually selecting a circular region of interest (ROI) from the control panel and placing this, using a mouse, within either the iliac or femoral arteries, whichever was best visualized in the imaging plane. Arterial attenuation change was determined over the 65 s acquisition by the software. A time-attenuation curve was generated automatically, and from this the timing of the end of the first pass of contrast could be estimated by visual inspection of the shape of the time-attenuation curve. Subsequent selection of this time point separating the two phases of contrast enhancement, as necessitated by the software program, then permitted generation of the four perfusion parametric maps (blood volume, blood flow, mean transit time and permeability) for all of the tissues within the imaging plane, within the processing threshold selected. This arterial ROI was saved using the software so that the exact same sized ROI could be automatically placed in the same location in subsequent analysis in an attempt to minimize measurement variability due to this confounder. A ROI was then drawn freehand around the peripheral margin of the tumour using an electronic cursor and mouse. Care was taken to exclude perirectal or pericolonic fat and also intraluminal gas, a process that was facilitated by viewing a cine-loop of the perfusion acquisition in order to gauge the degree and margins of patient movement during acquisition. A time-attenuation curve for the selected tumour tissue and the four perfusion parameters within the tumour ROI were then generated. Mean values for the four tumour perfusion parameters (blood volume, blood flow, mean transit time and permeability) from this single 5 mm axial slice were recorded for each individual patient. Image analysis was repeated for the remaining three contiguous axial slices in the same manner, recalling the arterial ROI to minimize variability due to this. The tumour ROI was drawn freehand due to morphological differences in each slice. Mean values for all four perfusion parameters were recorded for each of the three contiguous slices. Then, in order to obtain overall mean values for all four perfusion parameters for an equivalent z-axis tumour coverage of 20 mm, values were averaged from all four 5 mm axial slices, and recorded for each individual patient. Image analysis was performed by the same investigator for the second set of dynamic scans obtained from each of the 10 patients. Analysis was performed exactly as previously, for a single 5 mm axial slice, most similar to the initial analysis, and for all four slices. Mean values for the four tumour perfusion parameters (blood volume, blood flow, mean transit time and permeability) for a zaxis tumour coverage of 5 mm, and for a z-axis tumour coverage of 20 mm were recorded for each individual patient. 579

V Goh, S Halligan, L Gartner et al

Statistical analysis The mean (standard deviation (SD)) of tumour blood volume, blood flow, mean transit time and permeability measurements from a 5 mm and 20 mm z-axis tumour coverage were determined. Measurement reproducibility was assessed using the Bland Altman test statistic [6, 7]: the mean difference, standard deviation (SD) of the differences, and 95% limits of agreement were calculated for each of the four perfusion parameters (blood volume, blood flow, mean transit time and permeability) for a 5 mm and 20 mm z-axis tumour coverage.

Results The mean value (standard deviation, SD) for each perfusion parameter and the mean difference, SD of the differences, and 95% limits of agreement are summarized in Table 1, with corresponding scatter and BlandAltman agreement plots for blood volume and flow shown in Figures 1 and 2. Similar limits of agreement were obtained between measurements from a 5 mm and 20 mm z-axis coverage, indicating no improvement in measurement reproducibility.

Discussion CT perfusion has been advocated to monitor the effects of antivascular and antiangiogenic drugs because this quantitative technique allows in vivo assessment of tumour vascularity [8], and is more robust than techniques such as dynamic contrast enhanced MRI [9]. In vivo functional assessment is advantageous because histological methods that measure angiogenesis, such as microvessel density counting, may not accurately reflect functional tumour vascularity [5]. However, whether assessment of tumour perfusion at a single tumour level, rather than of the whole tumour, is representative and adequate is questionable, particularly in the context of therapeutic assessment. Multidetector row CT permits a number of contiguous axial images to be acquired at a given tumour level, and z-axis coverage is contingent on the number of detector rows. For example, the maximum tumour coverage

achievable for a perfusion study on the four-detector row scanner is 20 mm in the z-axis dimension, consisting of 4 contiguous 5 mm axial slices in cine mode. However, the rapid pace of technological advancement has meant that a 40 mm acquisition in the z-axis dimension, consisting of 8 contiguous 5 mm axial slices, will soon be possible for a perfusion study with the latest 64 detector row scanners. It seems logical to assume that by increasing tumour coverage, a more representative assessment of global tumour perfusion would be obtained, and that this would be less prone to measurement error. However, this has not been proven to date with colorectal cancer. Reproducibility assessment encompasses quantification of all the intrinsic and extrinsic factors that contribute to measurement variability. These include tumour heterogeneity, CT technique, software variability, and observer variability [10]. While it is not possible to separate these factors, it is possible to minimize effects of certain factors to permit evaluation the effect of the volume of data surveyed on reproducibility. For example, it is possible to minimize the effects of observer variability by using a single experienced observer to obtain all measurements [11], and to minimize effects from the technique and software evaluation by keeping technical factors constant between studies. Thus, we compared the reproducibility of measurements from a 5 mm z-axis tumour coverage with that from 4 contiguous 5 mm slices, the results from which had been averaged to provide information equivalent to from a 20 mm z-axis tumour coverage to identify if any improvement in reproducibility was achieved. Overall, reproducibility was acceptable for all four perfusion measurements obtained from both a 5 mm and 20 mm tumour coverage, and comparable with previously cited reproducibility levels using CT within the intracranial and extracranial circulation in animal and human studies [12–15]. For example, a variability of 13% and 7%, assessed using analysis of variance, has been quoted for repeated cerebral blood flow and blood volume measurements, respectively, in rabbits [13], while a variability of 14%, 20% and 18% has been quoted for repeated blood flow, blood volume and permeability measurements in the rabbit VX2 tumour [15]. Analysing our data in the same manner using analysis of variance

Table 1. 95% limits of agreement analysis showing the mean difference, standard deviation (SD) of differences and 95% limits of agreement for repeated measurements of tumour blood volume, blood flow, mean transit time and permeability for a 5 mm and 20 mm z-axis tumour coverage. The overall mean value (SD) for each perfusion parameter is also shown Tumour perfusion measurement

Blood volume (ml/100 g tissue) Blood flow (ml/100 g tissue/min) Mean transit time (seconds) Permeability (ml/100 g tissue/min)

580

z-axis coverage (mm)

Mean (SD)

Mean difference

SD

95% limits of agreement

5

6.1 (1.3)

–0.04

1.26

–2.50 to +2.42

20 5

5.9 (1.3) 91.1 (34.4)

–0.04 +8.8

1.31 30.2

–2.61 to +2.53 –50.5 to +68.0

20 5

81.1 (28.9) 7.3 (3.3)

+7.4 –0.99

29.4 3.67

–50.3 to +65.0 –8.19 to +6.20

20 5

8.5 (3.9) 14.1 (3.6)

–2.46 +1.20

5.18 3.38

–12.61 to +7.69 –5.42 to +7.83

20

13.6 (2.9)

–0.23

4.12

–8.31 to +7.85



Figure 1. Blood volume: (a) scatter plot showing the distribution of values with a line of perfect agreement shown; and (b,c) Bland-Altman agreement plots with the mean difference and 95% limits of agreement for a 5 mm (D) and 20 mm (O) z-axis tumour coverage.

would have resulted in a variability of 23%, 14% and 17% for colorectal blood flow, blood volume and permeability, respectively. Reproducibilty of permeability measurements is also comparable with that reported for dynamic contrast enhanced MRI for a variety of extracranial tumours, where a mean difference of 0.03, 95% confidence interval of 20.04 to +0.06, and coefficient of variation of 29% was reported for log transformed values of ktrans [16]. With reference to published data, the differences between repeated blood flow measurements that we observed may be sufficiently small to be overwhelmed by the change in perfusion induced therapeutically by antivascular targeting and antiangiogenic agents being The British Journal of Radiology, July 2006

evaluated currently. For example, dynamic contrast enhanced (DCE)-MRI data from a Phase I study of combretastatin, a vascular targeting agent, have shown a significant group mean measurement change of 37% after drug administration [17]. Likewise, perfusion CT data from a Phase I study of bevacizumab have shown mean blood flow changes of the order of 40% [18], while data from a study of the effects of chemoradiation have shown a mean blood flow changes of 62% [19]. However, for treatments that produce a smaller effect on tumour vascularity, some caution must be applied to the in the interpretation of results, as measurement change may remain within the range of measurement variability. On an individual patient basis, there may be a role for 581

V Goh, S Halligan, L Gartner et al

Figure 2. Blood flow: (a) scatter plot showing the distribution of values with a line of perfect agreement shown; and (b,c) Bland-Altman agreement plots with the mean difference and 95% limits of agreement for a 5 mm (D) and 20 mm (O) z-axis tumour coverage.

assessing individual intrapatient measurement variability, so that any response may be interpreted in the light of this. An alternative would be to use surrogates that show better reproducibility, for example semiquantitative measurements such as peak enhancement, or standardized perfusion value. Further evaluation of these strategies is required. Reproducibility did not improve with increasing tumour z-axis coverage. There may be several explanations why data acquisition from a larger tumour volume did not improve reproducibility. Most obviously, it may be the case that colorectal spatial heterogeneity is insufficient to influence data acquisition once the tumour 582

volume has achieved 5 mm in its z-axis dimension. This is in contrast to preliminary data from lung tumours that suggests improved reproducibility with increasing z-axis acquisition [20]. However, the lung tumours studied were large heterogeneous tumours with a necrotic centre, unlike the colorectal tumours within this study, and thus spatial heterogeneity may have been a major factor contributing to measurement variability in these tumours. It is also possible that the variation in tumour morphology over time encountered with colorectal cancer may be partly responsible for the lack of improvement observed. For example, bowel tumours The British Journal of Radiology, July 2006


might change shape and position as a result of peristalsis. Assessment of reproducibility from a single 5 mm-tumour level is straightforward, as it is usually possible to ensure comparable tumour levels on both sets of scans when choosing from the 4 slices available. However, we found that ensuring the same tumour level was assessed on all 4 slices was more difficult, because there was no leeway in post-acquisition slice selection. Tumour shape and position were unlikely to be identical on both sets of scans, despite care in scan acquisition, including administration of an antiperistaltic agent. It is possible that the quantitative software analysis method we used may have contributed to the lack of improvement found with increased tumour coverage. While the ROI used to define the arterial input may be saved using the software for future use, to minimize variability from ROI placement in subsequent analyses, due to variation in tumour shape between studies, a ROI has to be drawn freehand around the tumour on the parametric perfusion maps, which are generated automatically by the software to obtain the quantitative parameters. Each colour pixel on these maps represents an individual quantitative value of the parameter measured (blood flow, blood volume, mean transit time or permeability) and a mean parameter value is generated for the ROI drawn. It is inevitable that some variability will be introduced when drawing a ROI (though this can be reduced by using the same unblinded observer to create each ROI). Finally, our patient numbers were small and inevitably reduced the power of our experiment, but this was a pragmatic consideration based on a natural reluctance to unnecessarily expose patients to ionizing radiation on two separate occasions. In summary, we found that increasing z-axis tumour coverage did not improve reproducibility of measurements of colorectal tumour perfusion. Measurement reproducibility remains clinically acceptable with single level measurements.

6.

7.

8. 9.

10. 11.

12.

13.

14.

15.

16.

17.

References 1. Therasse P, Arbuck SG, Eisenhauer EA, Wanders J, Kaplan RS, Rubenstein L, et al. New guidelines to evaluate the response to treatment in solid tumors. J Natl Cancer Inst 2000;92:205–16. 2. World Health Organisation. WHO hand book for reporting results of cancer treatment. WHO offset publication 1979;no 48. Geneva, Switzerland. 3. Tozer GM. Measuring tumour vascular response to antivascular and antiangiogenic drugs. Br J Radiol 2003;76: S23–S35. 4. Hurwitz H, Fehrenbacher L, Novotny W, Cartwright T, Hainsworth J, Heim W, et al. Bevacizumab plus irinotecan, fluorouracil, and leucovorin for metastatic colorectal cancer. N Engl J Med 2004;350:2335–42. 5. Endrich B, Vaupel P. The role of the microcirculation in the treatment of malignant tumors: facts and fiction. In: Molls


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M, Vaupel P, editors. Blood perfusion and microenvironment of human tumors: implications for clinical radiooncology. Berlin, Germany: Springer-Verlag, 2001;19–39. Bland JM, Altman DG. Statistical methods for assessing agreement between two methods of clinical measurement. Lancet 1986;1:307–10. Bland JM, Altman DG. Applying the right statistics: analyses of measurement studies. Ultrasound Obstet Gynecol 2003;22:85–93. Miles KA, Griffiths MR. Perfusion CT: a worthwhile enhancement? Br J Radiol 2003;76:220–31. Goh V, Halligan S, Balmer JA, Bartram CI. Colorectal cancer perfusion measurement with multi-detector computed tomography: measurement reproducibility and clinical implications. In: Radiological Society of North America 90th Scientific Assembly and Annual Meeting Program; 2004. Chicago, USA: RSNA, 2004. Bland M. An introduction to medical statistics, 3rd edn. Oxford, UK: Oxford University Press, 2003. Goh V, Halligan S, Hugill JA, Bassett P, Bartram CI. Quantitative assessment of colorectal cancer perfusion using MDCT: inter and intra-observer agreement. AJR Am J Roentgenol 2005;185:225–31. Nabavi DG, Cenic A, Dool J, Smith RM, Espinosa F, Craen R, et al. Quantitative assessment of cerebral hemodynamics using CT: stability, accuracy, and precision studies in dogs. J Comput Assist Tomogr 1999;23:506–15. Cenic A, Nabavi DG, Craen RA, Gelb AW, Lee TY. A CT method to measure hemodynamics in brain tumors: validation and application of cerebral flow maps. AJNR Am J Neuroradiol 2000;21:462–70. Gillard JH, Antoun NM, Burnet NG, Pickard JD. Reproducibility of quantitative CT perfusion imaging. Br J Radiol 2001;74:552–5. Purdie TG, Henderson E, Lee TY. Functional CT imaging of angiogenesis in rabbit VX2 soft tissue tumour. Phys Med Biol 2001;46:3161–75. Galbraith SM, Lodge MA, Taylor NJ, Rustin GJ, Bentzen S, Stirling JJ, et al. Reproducibility of dynamic contrast enhanced MRI in human muscle and tumours: comparison of quantitative and semi-quantitative analysis. NMR Biomed 2002;15:132–42. Galbraith SM, Maxwell RJ, Lodge MA, Tozer GM, Wilson J, Taylor NJ, et al. Combretastatin A4 Phosphate has tumor antivascular activity in rat and man as demonstrated by dynamic magnetic resonance imaging. J Clin Oncol 2003;21:2831–42. Willett CG, Boucher Y, di Tomaso E, Duda DG, Munn LL, Tong RT, et al. Direct evidence that the VEGF-specific antibody bevacizumab has antivascular effects in human rectal cancer. Nature Med 2004;10:145–7. Sahani DV, Kalva SP, Hamberg LM, Hahn PF, Willett CG, Saini S, et al. Assessing tumor perfusion and treatment response in rectal cancer with multisection CT: initial observations. Radiology 2005;234:785–92. Ng QS, Goh V, Klotz E, Fichte H, Carnell D, Saunders M, et al. Single versus four-level tumour perfusion measurement using helical dynamic contrast enhanced MDCT: effect on measurement agreement and potential for whole tumour assessment. In: Proceedings for European Congress of Radiology 2005; Vienna, Austria. Eur Radiol 2005; 15(Suppl. 1):414.

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Prophylactic implantation of inferior vena cava filter during interventional radiological treatment for deep venous thrombosis of the lower extremity T YAMAGAMI, MD, PhD, T KATO, T NISHIMURA, MD, PhD

MD,

T HIROTA,

MD,

R YOSHIMATSU,

MD,

T MATSUMOTO,

MD

and

Department of Radiology, Graduate School of Medical Science, Kyoto Prefectural University of Medicine, 465 Kajii-chyo, Kawaramachi-Hirokoji, Kamigyo, Kyoto, 602-8566, Japan ABSTRACT. The purpose of this study was to evaluate the filtering effect of the Gunther tulip retrievable vena cava filter (GTF) during treatment of deep venous thrombosis (DVT) in the lower extremity using various interventional radiological procedures. Subjects of the study were all 17 consecutive patients (8 women, 9 men; age range 18– 87 years; mean age 55.9 years) with symptomatic lower limb DVT referred for interventional radiological treatment between February 2001 and September 2004. In all of these patients, the GTF was implanted during interventional radiological treatment. Trapped thrombus in the filter was evaluated with venocavography performed repeatedly during the treatment for DVT. Implantation of a total of 29 GTFs was successfully performed in the 17 patients. In 10 (58.8%), more than 2 filters were subsequently implanted to prolong implantation time. Also in 10 patients, the DVT resolved after therapy and retrieval of the final GTF was successful with one exception. Worsening of or new formation of pulmonary embolism was avoided in all patients. In 8 (47.1%) of the 17 patients, a trapped thrombus in the GTF was observed during treatment for DVT. In six patients the trapped thrombus was large, filling more than half the height of the filter. In conclusion, we found that the GTF is effective in filtering the relieved thrombus from DVT in the lower extremity and in protecting against movement of the thrombus to the pulmonary artery during therapies with interventional radiological procedures.

The initial symptoms of deep venous thrombosis (DVT) in the lower extremities may be quite severe [1, 2] and, additionally, there is substantial risk of pulmonary embolism [1, 3]. Moreover, in DVTs that remain, delayed complications are possible, including the spectrum of debilitating symptoms referred to as post-thrombotic syndrome [4]. Therefore, treatment of DVT is necessary. Traditionally, therapy for DVT has consisted of systemic intravenous administration of heparin [5]. Other options have been systemic intravenous administration of fibrinolytic agents such as urokinase or streptokinase [5–7]. Recombinant tissue plasminogen activator is gaining attention as an effective fibrinolytic agent [8, 9]. Systemic thrombolysis is more effective than heparinization, but less effective than catheter-directed thrombolysis, which is an interventional radiological treatment that has recently become widely accepted as useful in treating DVT [4, 9– 11]. Most recently, other interventional radiological procedures such as manual aspiration of thrombus [12], mechanical thrombectomy [4, 13, 14], percutaneous transluminal angioplasty [13, 15] and self-expandable metallic stent placement [4, 13, 15, 16] have sometimes been combined with catheter-directed thrombolysis to improve the quality of treatment for DVT. Theoretically, a thrombus released from the DVT during catheter-directed thrombolysis could move to the pulmonary artery, which may cause pulmonary 584

Received 30 August 2005 Revised 13 November 2005 Accepted 10 January 2006 DOI: 10.1259/bjr/54853087 ’ 2006 The British Institute of Radiology

embolism. Thus, some physicians implant inferior vena cava filters prophylactically to prevent such an occurrence [17–20]. However, the prophylactic use of an inferior vena cava filter is considered by some to be unnecessary [1, 10, 11], with the result that this procedure is currently very controversial. Nevertheless, in our daily clinical work, we have experienced many situations where, if a filter had not been implanted during catheter-directed thrombolysis, severe pulmonary embolism might have occurred. Many interventional radiological procedures for DVT other than catheterdirected thrombolysis have been developed [12–16], but there is little information on the prophylactic use of filters in these procedures [21]. When such prophylactic use of filters is required, ideally, a permanent inferior vena cava filter would not be placed, considering the long life expectancy of such patients [22, 23]. Hence, temporary vena cava filters have been widely used [17, 24]. However, paralleling the increase in the use of temporary vena cava filters have been reports describing complications related to their insertion, mainly associated with their structure, where part of the device projects from the insertion site [24–26]. Because of such problems, it would be advantageous to use a retrievable vena cava filter, that is, a permanent filter that can be retrieved when necessary. The Gunther tulip retrievable vena cava filter (GTF) (Cook, The British Journal of Radiology, July 2006

Inferior vena cava filter for DVT

Bjaeverskov, Denmark) is among the commercially available retrievable vena cava filters. The present study evaluated whether an inferior vena cava filter could have a role in filtering a thrombus released from the lower limb during treatment of lower extremity DVT using various interventional radiological procedures. Also examined were the feasibility and safety of a GTF used for that purpose.

Subjects and methods Patients The cohort of this study was 17 consecutive patients (8 women, 9 men; age range 18–87 years; mean age 55.9 years) who were treated for symptomatic DVT of the lower extremity, utilizing various interventional radiological procedures between February 2001 and September 2004, at our institution. In all of these patients, traditional intravenously administered systemic anticoagulation and/or thrombolysis had been performed first as a treatment for DVT. However, despite such treatment, DVT developed. A GTF was implanted to prevent the complication of pulmonary embolism during the entire period of interventional radiological treatment of DVT (Figure 1). Interval from initial symptoms to start of therapy for DVT was within 1 week in 11 patients, between 1 week and 4 weeks in 5 patients and after 6 weeks in 1 patient. In all cases, the existence of lower extremity DVT was diagnosed by ultrasound. When extension of the thrombus from the lower extremity to the inferior vena cava was suspected, enhanced abdominal CT was performed. Intravenous ascending venography with contrast injection into the vein of the foot was also performed as an interventional radiological diagnostic procedure. Before starting therapies for DVT, enhanced chest CT was performed in all 17 patients to evaluate the existence of pulmonary embolism. Pulmonary embolism was found in two patients, who received intravenously administered anticoagulation and thrombolytic therapies for pulmonary embolism, in addition to therapies for DVT. Enhanced chest CT was also obtained after therapies for DVT in all 17 patients.

Procedures All interventional radiological procedures performed as therapy for DVT and GTF placement and retrieval were performed by one of three experienced interventional radiologists in our institution after written informed consent was obtained from each patient. The consent included use of records, images, data, etc. for research purposes. Our institution does not require institutional review board approval for this type of report. Principles of the Declaration of Helsinki were followed. The following are interventional radiological therapies for DVT that were performed in our institution and that were reviewed retrospectively in all 17 patients who received such therapy. The British Journal of Radiology, July 2006

In the angiography suite, catheter-directed thrombolysis was performed after a 0.035-inch guide-wire was advanced via a 5-French catheter over the thrombosed segment and this catheter was exchanged with a commercially available multiside-hole catheter (CraggMcNamara Valved Infusion Catheter; Micro Therapeutics, Irvine, CA). Venous access was through the femoral, jugular, or popliteal vein. After positioning the tip of the multiside-hole catheter in the thrombosed segment, a mixture of 120 000–240 000 IU urokinase and 1000–2000 IU heparin diluted by 20–40 ml of physiological saline was manually administered over a period of approximately 10 min. In the majority of patients (n516), because the thrombus remained, 120 000–240 000 IU urokinase per day was further administered persistently via the catheter which remained positioned in the thrombosed segment over a prolonged period, even after the patient returned to the ward from the angiography suite. Heparin was administered systemically during the lytic procedure and after lysis to keep the partial thromboplastin time at 75–100 s. When DVT was shown to be completely or almost completely resolved on venography via the catheter inserted for catheter-directed thrombolysis, lysis was stopped. But if much of the thrombus remained as shown on this venography, manual aspiration of the venous thrombus and/or mechanical thrombectomy were often combined concurrently with or after catheter-directed thrombolysis. For manual aspiration, after a 10-French or 11-French with 11 cm long sheath introducer was inserted from the femoral vein at the site of the thrombosis, an 8-French with 25 cm long sheath was advanced through the wider sheath into the segment occupied by the thrombus. The 8-French sheath was fitted with a syringe by which the clot was manually aspirated. After the aspiration, the 8-French sheath was withdrawn from the 10-French or 11-French sheath. If clots of thrombus were observed, the 8-French sheath was reinserted and the thrombus was aspirated. These steps were repeated until the venous thrombus was no longer observed by aspiration. Alternatively, in the more recent cases, after a 6-French sheath introducer was inserted from the femoral vein at the site of DVT, the venous thrombus was aspirated utilizing a 6-French catheter (Thrombuster; Kaneka, Tokyo, Japan). This catheter is commercially available and was developed exclusively for aspiration of a venous thrombus. Mechanical thrombectomy was performed using a commercially available thrombectomy device (Hydrolyser; Cordis or Oasis; Boston Scientific) with access from the femoral vein. When an underlying stenosis or obstruction due to a non-fresh thrombus remained, percutaneous transluminal angioplasty for the segment that was narrowed due to thrombus was occasionally performed. If necessary, a selfexpandable metallic stent was placed to obtain sufficient relief of the lesion, particularly at the level of the iliac vein. With the femoral or jugular venous approach, percutaneous transluminal angioplasty was performed with a commercially available angioplasty balloon catheter after advancing a 0.035-inch guidewire over the narrowed lesion. The angioplasty balloon catheter was positioned in the narrow segment, after which percutaneous transluminal angioplasty was performed for the entire narrow 585

T Yamagami, T Kato, T Hirota et al

586

(a)

(b)

(c)

(d)


Inferior vena cava filter for DVT Figure 1. A 19-year-old man (patient no. 6). (a) Venography shows venous thrombus distributing from the left common iliac vein to the left femoral vein (arrows). Note that Gunther tulip retrievable vena cava filter (GTF) was implanted at the inferior vena cava with an approach from the right jugular vein. (b) Venography obtained immediately after catheter-directed thrombolysis still shows that much of the thrombus remained. Note that multiple side holes of the catheter for catheter-directed thrombolysis were positioned in the segment of the thrombus (arrows). (c) Venocavography obtained from the right femoral vein shows that thrombus filling greater than the height of filter was captured in the filter (arrows). This image was obtained just after manual aspiration, mechanical thrombectomy and percutaneous transluminal angioplasty were performed. These interventional radiological procedures were performed after catheter-directed thrombolysis, as shown in Figure 1a,b, and prolonged thrombolysis performed subsequently in the ward through the catheter with its tip positioned at the thrombosed segment. (d) Roentgenogram shows a temporary filter (arrow) (Neuhaus Protect, Toray) that was inserted at the cephalad level of GTF before GTF with thrombus was retrieved to be exchanged with a new GTF. (Continued)

(e) Figure 1. (Cont.) (e) Venography through the right femoral vein, obtained after GTF was exchanged for a new one (arrow), shows no thrombus remaining in the inferior vena cava. The new GTF was implanted to continue treatment for the DVT, further avoiding the complication of pulmonary embolism. 7 days after this, metallic stent placement was performed, then venous thrombus resolved and rapid blood flow in the previously obstructed segment due to thrombus in the left common and external iliac and femoral vein was obtained.

segment. Either the Ultra-thin diamond balloon catheter (Boston Scientific, Watertown, MA) or the Power flex plus balloon catheter (Cordis, Miami, FL) was used. For selfexpandable metallic stent placement, an Easywall stent (Boston Scientific) or spiral Gianturco Z stent (Cook) was implanted with the femoral venous approach. Placement over the area of joint movement (i.e. hip joint) was avoided. The British Journal of Radiology, July 2006

During the entire period of therapy by the abovedescribed interventional radiological procedures, the GTF remained in place. When the period of implantation greatly exceeded 10–14 days, the period cited in many reports within which a filter could be withdrawn safely [27–30], the first implanted GTF was retrieved and another inserted near at a different site from the former GTF. When the venous thrombus in the lower extremities resolved after therapy, as shown on intravenous venography of the lower extremity, the GTF was retrieved and reinsertion was not required. However, when it remained in spite of various therapies, the last GTF inserted during therapy was left in the vena cava without retrieval to play a role thereafter as a permanent inferior vena cava filter. After a venacavogram was performed to plan the position of filter placement, the GTF was introduced through the right internal jugular vein and was positioned at the cephalad side of the top of the venous thrombus. In cases when insertion from the right internal jugular vein was difficult, the right subclavian vein was used to introduce the GTF. In principle, the GTF was positioned at the infrarenal inferior vena cava. The GTF was placed through the sheath introducer according to the manufacturer’s instructions. Retrieval was performed using the GTF retrieval set supplied by the manufacturer (Cook). Details of techniques to place and retrieve the GTF are described elsewhere [29, 31]. Venocavography was performed repeatedly via the inserted catheter positioned in the thrombosed segment to perform catheter-directed thrombolysis continuously over a prolonged period in the ward. After retrieving the catheter for thrombolysis, venocavography was obtained through an 18-gauge intravenous catheter (Surflo; Terumo, Tokyo, Japan) or a 4-French catheter, which was inserted from the femoral vein in which the thrombus did not exist. In cases when a GTF was exchanged or retrieved, just prior to retrieval, venocavography was performed through the sheath. Trapped thrombus in the filter was evaluated from all these venocavograms obtained by various approaches. Also, immediately after retrieving the GTF, venography was performed to confirm that no thrombus remained where the retrieved GTF had been implanted.

Investigated parameters We retrospectively investigated the following: selected interventional radiological treatments for DVT, state of DVT after such therapies, number of GTFs implanted and retrieved, rate of success in preventing through GTF 587


placement the worsening or development of pulmonary embolism and degree of trapped thrombus in the filter.

Results As shown in Tables 1 and 2, implantation in the planned position of a total of 29 GTFs was successful in 17 patients with DVT in the lower extremity for which various therapies utilizing interventional radiological procedures were performed. Details of each GTF implanted are shown in Table 3. In four patients, the GTF was placed at the suprarenal inferior vena cava for the following reasons. In two of those four patients, a venous thrombus existed in the inferior vena cava just below the level of entry of the renal vein, leaving insufficient space for implantation in the infrarenal inferior vena cava. The third patient had a double inferior vena cava with the thrombus in both. In the fourth patient, placement of an additional GTF at the cephalad level of the normally placed GTF was necessary. Retrieval of the GTF was successful in all 22 attempts, with one exception (Table 3). The implantation period per filter before retrieval was 10.4¡5.0 days (mean¡SD; range 0–18 days; median 13 days). The period of treatment with GTFs in place was 19.1¡8.2 days (mean¡SD; range 5–37 days; median 20 days). In 10 (58.8%) of the 17 patients, more than 2 filters were subsequently implanted to prolong implantation time. No complications were encountered relating to filter insertion and retrieval. In both patients having a pulmonary embolism at the beginning of interventional radiological treatments for DVT as shown in Table 1, the pulmonary embolism had

resolved at the end of therapy for DVT. Pulmonary embolism did not occur in any of the 15 patients without pre-existing pulmonary embolism. On venocavography performed repeatedly during the entire period of therapies for DVT with interventional radiological techniques, in 8 (47.1%) of the 17 patients, a trapped clot of thrombus was observed. Trapped thrombus filled , 1/4 the height of the filter in 1, from >1/4 to , 1/2 in 1, and >1/2 but within the height of the filter in 4. In 2, the trapped thrombus occupied greater than the height of the filter. Among those eight patients, in only one patient was catheter-directed thrombolysis the only interventional radiological procedure performed. In the remaining seven patients, other interventional radiological procedures were added. In other words, a trapped thrombus in the filter was confirmed in one (25.0%) of four patients who received only catheter-directed thrombolysis. However, a trapped thrombus was observed in seven (53.8%) of 13 patients in whom other interventional radiological treatments were combined with catheterdirected thrombolysis. The difference was not significant (p50.34, Fisher’s exact probability test). All filters containing a trapped thrombus as revealed on venocavogram were retrieved without any complications, including pulmonary embolism. Attempts were made to decrease all trapped thrombi in the filter by catheterdirected thrombolysis, whereby the catheter tip that had been in the DVT lesion was repositioned to the filter. In two cases, both of which had a trapped thrombus that occupied greater than the height of the filter, manual aspiration of the thrombus in the filter was additionally performed using the sheath introducer inserted from the

Table 1. Characteristics of patients with deep venous thrombosis of the lower extremity Patient no./sex/ age (years)

Site of DVT

Location of venous thrombus

Interventional radiological procedures performed as therapy for DVT

Existence of PE

Final status of DVT

1/m/87

bilateral

IVC, CI, EI, F

CDT, p-CDT, A, PTA, EMS

absent

2/f/62 3/f/18

left right

CI, EI, F, P F

CDT, p-CDT A, MT, PTA, EMS CDT, p-CDT, A, PTA

absent absent

4/f/83

right

IVC, CI, EI, F

CDT, A

absent

5/f/52 6/m/19 7/f/41 8/m/54

bilateral left bilateral right

F, P CI, EI, F EI, F F

CDT, CDT, CDT, CDT,

present absent absent absent

9/m/76 10/m/21

right left

IVCa, CI, EI, F CI

CDT, p-CDT CDT, p-CDT, A

absent absent

11/f/80 12/m/63

bilateral left

IVC, CI, EI, F CI, EI, F

CDT, p-CDT, A, PTA, EMS CDT, p-CDT, A

absent present

13/f/75 14/m/27

left left

IVC, CI, EI CI, EI, F, P

CDT, p-CDT, A CDT, p-CDT, A, PTA

absent absent

15/m/77 16/f/61 17/m/55

right left left

CI, EI EI, F EI, F

CDT, p-CDT, A, PTA CDT, p-CDT, A, PTA CDT, p-CDT, A

absent absent absent

resolved decreased present resolved decreased present decreased present resolved resolved resolved decreased present resolved decreased present resolved decreased present resolved decreased present resolved resolved

p-CDT p-CDT, A, MT, PTA, EMS p-CDT p-CDT

but still

but still but still

but still

but still

but still

but still

IVC, inferior vena cava; CI, common iliac vein; EI, external iliac vein; F, femoral vein; P, popliteal vein; CDT, catheter-directed thrombolysis; p-CDT, prolonged thrombolysis performed from the catheter directly inserted in the thrombosed segment; A, aspiration of venous thrombus; MT, mechanical thrombectomy; PTA, percutaneous transluminal angioplasty; EMS, placement of self-expandable metallic stent; DVT, deep venous thrombus; PE, pulmonary embolism. a Case with double inferior vena cava.

588


Inferior vena cava filter for DVT Table 2. Usage of Gunther retrievable vena cava filter Patient no.

No. filters used

Final use of GTF as a permanent vena cava filter

Duration of therapy under protection from PE with GTF prior to exclusive use of conventional therapy (days)

Trapped thrombus in GTF as revealed

1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17

2 2 2 1a 1 2 2 3b 1 3 1 1 2 2 1 2 1

No Yes No Yes Yes Yesc No No Yes No Yes No Yes No Yes No No

8 24 23 7 5 21 28 21 15 37 19 18 14 28 23 20 14

Grade Grade Grade Grade Grade Grade Grade Grade Grade Grade Grade Grade Grade Grade Grade Grade Grade

1 3 3 0 0 4 0 4 0 0 0 2 3 0 0 3 0

a

Cases requiring correction of filter placement. One filter being placed during single interventional radiological procedure. c Case of failure to retrieve filter. PE, pulmonary embolism; GTF, Gunther tulip retrievable vena cava filter. Grade 0: No trapped thrombus is seen. Grade 1: Trapped thrombus filled >1/4 the height of the filter. Grade 2: Trapped thrombus filled from >1/4 to >1/2 the height of the filter. Grade 3: Trapped thrombus filled >1/2 but within the height of the filter. Grade 4: Trapped thrombus filled greater than the height of the filter. b

Table 3. Characteristics of each Gunther retrievable vena cava filter GTF implanted

Number

Approach for implantation of GTF right jugular vein right subclavian vein Level of IVC where GTF was placed infrarenal suprarenal Final status of each filter Left in IVC to use as a permanent filter thereafter Failed attempt at retrieval after resolution of DVTa Successfully retrieved exchanged with another GTF to prolong the implantation period temporarily placed during an IR procedure and retrieved immediately after that procedure retrieved and further implantation was not performed as the DVT had resolved

28 1 25 4 7 1 21 11 1 9

GTF, Gunther tulip retrievable vena cava filter; IVC, inferior vena cava; DVT, deep venous thrombus; IR, interventional radiology. Hook part of the filter had become attached to the wall of the inferior vena cava resulting in failure of retrieval.

a

femoral vein into the inside of the filter. As a result, two GTFs were retrieved after the thrombus resolved following attempts to decrease the size of the captured thrombus by means such as catheter-directed thrombolysis. Regarding the remaining six cases, the size of the thrombus in the filter either did not change or had increased, so the filter was replaced with a new one. In exchanging the filter, a temporary filter (Antheor; Boston Scientific, Watertown, MA, n54, or Neuhaus Protect, Toray Medical, Tokyo, Japan, n51) or another GTF (n51) was temporarily placed at the cephalad level of the GTF, trapping the thrombus exclusively during the GTF retrieval procedure. Thus, the thrombus dropped from the first GTF was captured during GTF retrieval. In the case of another GTF used The British Journal of Radiology, July 2006

temporarily, displacement of that freshly placed GTF avoided by taking meticulous care in using the retrieval kit during the procedure to remove the originally implanted GTF. Venocavography after retrieval revealed no thrombus remaining where the first GTF had been placed in any of the eight patients. Figure 1 shows one such case.

Discussion There has been much controversy regarding the prophylactic use of inferior vena cava filters to prevent the occurrence of pulmonary embolism during catheterdirected thrombolysis of lower extremity DVT [1, 9–11, 589


17–21]. Some researchers have advocated that such filter placement is necessary during catheter-directed thrombolysis only in patients who have large, mobile, freefloating thrombi within the inferior vena [1]. On the other hand, there has been little discussion on the necessity of filter placement during therapies for DVT with interventional radiological procedures other than catheter-directed thrombolysis [21]. Pulmonary embolism occurred in only two (0.9%) of 214 patients who received catheter-directed thrombolysis for DVT in the lower extremity without inferior vena cava filter placement, according to Bjarnason et al [1], and in only 6 (1.3%) of 473 patients, according to Mewissen et al [11]. However, one of the six cases was a fatality. Some investigators [1, 9–11] insist that prophylactic use of an inferior vena cava filter during catheter-directed thrombolysis is unnecessary because of such a low rate of pulmonary embolism. In a summary of their experience with implantation of temporary vena cava filters in 132 patients with DVT who were receiving thrombolytic therapy, Thery et al [32] observed a clot lysed by thrombolytic therapy in the filter in 41 (31%) of the 132 patients. None of the 132 patients experienced pulmonary embolism. This suggests that the filter might prevent pulmonary embolism during thrombolytic treatment in at least 41 of these patients, if not in all 132 patients. Some researchers [17–21, 30, 33, 34] propose that the prophylactic use of an inferior vena cava filter is necessary in view of the high rate of trapped thrombus in these filters during therapies for DVT, as revealed by Thery et al [32], and attach great importance to the potential mortality in such cases, although rare [11]. Based on these opinions, we perform therapy for DVT under the prophylactic protection against pulmonary embolism by using the GTF. Ease and safety of insertion of GTF is well known [21, 27–31, 33, 35, 36], with some of these reports citing its use as a non-permanent filter [21, 27, 29, 30, 33–36]. Millward et al [33] reported that attempts at retrieval were successful in 98% of 53 GTFs. According to some previous reports [27–30], the maximal period of implantation before successful retrieval was recommended as no longer than 10–14 days. Although we had roughly followed this recommended period to implant the GTF in the patients presented here, successful retrieval of a GTF after implantation longer than 14 days has been reported, as described in some recently published studies. Millward et al [33] noted that the maximal period of implantation before successful retrieval was 25 days, and very recently Terhaar et al [37] reported a period of 126 days. Also, we would like to mention that other retrievable filters are now available that can be left in place for much longer [38]. Few complications related to GTF placement have been reported [33, 39]. Millward et al [33] documented filter occlusion in two (5%) of 90 patients with an implanted GTF; no serious complications were documented. However, another case report [39] described right atrial migration. Our study also showed good results in insertion and retrieval of GTFs. Insertion of all 29 GTFs and retrieval of 95.5% (21) of 22 GTFs were successfully accomplished. The maximal period of implantation before successful retrieval was 18 days. No complications were 590

encountered in filter placement and retrieval. In seven patients in whom the lower extremity DVT remained in spite of various treatments and in who the possibility of pulmonary embolism remained, the GTF was used as a permanent filter by simply leaving it in the inferior vena cava. On the other hand, the GTF was retrieved in 9 of the 10 patients as it was no longer needed after successful resolution of the deep venous thrombus. In the present study, in which a retrievable inferior vena cava filter was placed in all 17 patients with DVT who underwent various interventional radiological treatments such as catheter-directed thrombolysis, no worsening or development of pulmonary embolism in any of the 17 patients was seen. However, in eight (47.1%) patients a trapped thrombus in the filter was confirmed during interventional radiological treatments. This was despite adequate anticoagulation therapy. In these eight patients, especially in the six in whom the thrombus captured by the GTF filled greater than half the height of the filter, it is possible that if the filters had not been implanted, the thrombus might have moved to the pulmonary artery causing pulmonary embolism. Furthermore, we would like to note that the rate of occurrence of trapped thrombus in the filter was higher in cases that were treated not only with catheter-directed thrombolysis, but additionally with other interventional radiological procedures than in those solely treated by catheter-directed thrombolysis (53.8% versus 25.0%). The difference, however, was not statistically significant. In addition, our results showed that the DVT did not extend to the inferior vena cava in six (75.0%) of our eight cases with a captured thrombus in the filter. This might suggest that limiting the indication for a filter to large, mobile, free-floating thrombi within the inferior vena cava is insufficient to avoid a pulmonary embolism in such situations. We want to reiterate that, from the standpoint of preventing pulmonary embolism during therapies for DVT in the lower extremity with various interventional radiological procedures, the inferior vena cava filter played an important role in filtering a thrombus released from the lower limb. However, the use of permanent or temporary filters had some drawbacks, as previous studies have suggested [22–26]. Hence, when interventional radiological procedures are performed as therapy for DVT, prophylactic placement of a retrievable filter, such as GTF, should be employed, especially when another interventional radiological procedure is added to the catheter-directed thrombolysis, which is the procedure most commonly performed.

References 1. Bjarnason H, Kruse JR, Asinger DA, Nazarian GK, Dietz CA Jr, Caldwell MD, et al. Iliofemoral deep venous thrombosis: safety and efficacy outcome during 5 years of catheterdirected thrombolytic therapy. J Vasc Interv Radiol 1997;8:405–18. 2. Strandness DE Jr, Langlois Y, Cramer M, Randlett A, Thiele BL. Long-term sequelae of acute venous thrombosis. JAMA 1983;250:1289–92. 3. Hirsch J, Hoak J. Management of deep vein thrombosis and pulmonary embolism. Circulation 1996;93:2212–45. 4. Sharafuddin MJ, Sun S, Hoballah JJ, Youness FM, Sharp WJ, Roh BS. Endovascular management of venous thrombotic


Inferior vena cava filter for DVT

5.

6.

7.

8.

9.

10.

11.

12.

13. 14.

15.

16.

17.

18. 19.

20.

21.

22.

and occlusive diseases of the lower extremities. J Vasc Interv Radiol 2003;14:405–23. Elliot MS, Immelman EJ, Jeffery P, Benatar SR, Funston MR, Smith JA, et al. A comparative randomized trial of heparin versus streptokinase in the treatment of acute proximal venous thrombosis: an interim report of a prospective trial. Br J Surg 1979;66:838–43. Arnesen H, Hoiseth A, Ly B. Streptokinase or heparin in the treatment of deep vein thrombosis. Acta Med Scand 1982;211:65–8. Robertson BR, Nilsson JM, Nylander G. Value of streptokinase and heparin in treatment of acute deep venous thrombosis. Acta Chir Scand 1968;134:203–8. Sugimoto K, Hofmann LV, Razavi MK, Kee ST, Sze DY, Dake MD, et al. The safety, efficacy, and pharmacoeconomics of low-dose alteplase compared with urokinase for catheter-directed thrombolysis of arterial and venous occlusions. J Vasc Surg 2003;37:512–7. Grossman C, McPherson S. Safety and efficacy of catheterdirected thrombolysis for iliofemoral venous thrombosis. AJR Am J Roentgenol 1998;172:667–72. Semba CP, Dake MD. Iliofemoral deep venous thrombosis: aggressive therapy with catheter-directed thrombolysis. Radiology 1994;191:487–94. Mewissen MW, Seabrook GR, Meissner MH, Cynamon J, Labropoulos N, Haughton SH. Catheter-directed thrombolysis for lower extremity deep venous thrombosis: report of a national multicenter registry. Radiology 1999;211:39–49. Roy S, Laerum F. Transcatheter aspiration: the key to successful percutaneous treatment of deep venous thrombosis? Acad Radiol 1999;6:730–5. Bravo SM, Reinhart RD, Meyerovitz MF. Percutaneous venous interventions. Vascular Med 1998;3:61–6. Delomez M, Beregi JP, Willoteaux S, Bauchart JJ, d’Othee BJ, Asseman P, et al. Mechanical thrombectomy in patients with deep venous thrombosis. Cardiovasc Interv Radiol 2001;24:42–8. AbuRahma AF, Perkins SE, Wulu JT, Ng HK. Iliofemoral deep vein thrombosis: conventional therapy versus lysis and percutaneous transluminal angioplasty and stenting. Ann Surg 2001;233:752–60. Nazarian GK, Bjarnason H, Dietz CA Jr, Bernadas CA, Hunter DW. Iliofemoral venous stenoses: effectiveness of treatment with metallic endovascular stents. Radiology 1996;200:193–9. Neuerburg J, Gu¨nther RW. Developments in inferior vena cava filters: a European viewpoint. Semin Intervent Radiol 1994;11:349–57. Millward S. Temporary and retrievable inferior vena cava filters: current status. J Vasc Interv Radiol 1998;9:381–7. Emanuelli G, Segramora V, Frigerio C. Selected strategies in venous thromboembolism: local thrombolytic treatment and caval filters. Haematologica 1995;80:84–6. Palombo D, Porta C, Brustia P, Peinetti F, Udini M, Antico A, et al. Loco-regional thrombolysis in deep venous thrombosis. Phlebologie 1993;46:293–302. Yamagami T, Kato T, Iida S, Tanaka O, Nishimura T. Retrievable vena cava filter placement during treatment for deep venous thrombosis. Br J Radiol 2003;76:712–8. Ferris E, McCowan TC, Carver DK, McFarland DR. Percutaneous inferior vena cava filters: follow-up of seven designs in 320 patients. Radiology 1993;188:851–6.


23. Decousus H, Leizorovics A, Parent F, et al. A clinical trial of vena caval filters in the prevention of pulmonary embolism in patients with proximal deep-vein thrombosis. N Engl J Med 1998;338:409–15. 24. Lorch H, Welger D, Wagner V, Hillner B, Strecker EP, Herrmann H, et al. Current practice of temporary vena cava filter insertion: a multicenter registry. J Vasc Interv Radiol 2000;11:83–8. 25. Carcone B, Pernes JM, Carcopino JM, Chollet D, Pouillard F. Worsening of proximal thrombosis after insertion of a temporary caval filter (3 cases). Rev Med Interne 1995;16: 351–3. 26. Stosslein F, Altmann E. A rare complication with an Antheor vena cava filter. Cardiovasc Intervent Radiol 1998;21:165–7. 27. Tay KH, Martin ML, Fry PD, Webb JG, Machan LS. Repeated Gunther tulip inferior vena cava filter repositioning to prolong implantation time. J Vasc Interv Radiol 2002;13:509–12. 28. de Gregorio MA, Gamboa P, Gimeno MJ, Madariaga B, Tobio R, Herrera M, et al. The Gunther Tulip retrievable filter: prolonged temporary filtration by repositioning within the inferior vena cava. J Vasc Interv Radiol 2003;14:1259–65. 29. Neuerburg JM, Gunther RW, Vorwerk D, Dondelinger RF, Jager H, Lackner KJ, et al. Results of a multicenter study of the retrievable tulip vena cava filter: early clinical experience. Cardiovasc Interv Radiol 1997;20:10–6. 30. Millward SF, Bhargava A, Aquino J Jr, Peterson RA, Veinot JP, Bormanis J, et al. Gunther tulip filter: preliminary clinical experience with retrieval. J Vasc Interv Radiol 2000;11:75–82. 31. Millward SF. Gunther tulip retrievable filter: why, when and how? Can Assoc Radiol 2001;52:188–92. 32. Thery C, Bauchart JJ, Lesenne M, Asseman P, Flajollet JG, Legghe R, et al. Predictive factors of effectiveness of streptokinase in deep venous thrombosis. Am J Cardiol 1992;69:117–22. 33. Millward SF, Oliva VL, Bell SD, Valenti DA, Rasuli P, Asch M, et al. Gunther tulip retrievable vena cava filter: results from the registry of the Canadian interventional radiology association. J Vasc Interv Radiol 2001;12:1053–8. 34. Yamagami T, Kato T, Iida S, Hirota T, Nishimura T. Gunther tulip inferior vena cava filter placement during treatment for deep venous thrombosis of the lower extremity. Cardiovasc Interv Radiol 2005;28:442–53. 35. Owen RJT, Krarup KC. The successful use and removal of the Gunther tulip inferior vena caval filter in pregnancy. Clin Radiol 1997;52:241–3. 36. Lin M, Soo TB, Horn LC. Successful retrieval of infected Gunther tulip IVC filter. J Vasc Interv Radiol 2000;11: 1341–3. 37. Terhaar OA, Lyon SM, Given MF, Foster AE, McGrath F, Lee MJ. Extended interval for retrieval of Gunther tulip filters. J Vasc Interv Radiol 2004;15:1257–62. 38. Grande WJ, Trerotola SO, Reilly PM, Clark TWI, Soulen MC, Patel A, et al. Experience with the recovery filter as a retrievable inferior vena cava filter. J Vasc Interv Radiol 2005;16:1189–93. 39. Bochenek KM, Aruny JE, Tal MG. Right atrial migration and percutaneous retrieval of Gu¨nther tulip inferior vena cava filter. J Vasc Interv Radiol 2003;14:1207–9.

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A survey of MRI quality assurance programmes C J KOLLER,

MSc,

J P EATOUGH,

PhD,

P J MOUNTFORD,

PhD

and G FRAIN,

MMath

Medical Physics Directorate, University Hospital of North Staffordshire, Princes Road, Hartshill, Stoke on Trent ST4 7LN, UK ABSTRACT. There are currently no national guidelines on appropriate quality assurance (QA) test frequencies for MRI equipment in clinical use. From a random selection of 45 hospitals in England, who were contacted by phone, 35 hospitals agreed to participate in a survey of MRI QA and were sent a questionnaire requesting information on the range and frequency of QA tests, as well as the staff groups who conduct these tests. Twenty-four completed replies were received, representing a 68% response rate from the distributed questionnaires. Of these, 79% undertook some form of QA, typically conducted by the radiographic staff. Tests were most often undertaken on the head coil, but there was a considerable variation in the frequency and range of tests undertaken at different hospitals. For example, exactly half of the respondents conducted signal to noise ratio (SNR) tests on both head and body coils, but only 13% of centres extended this test to other coils. Results from this survey should inform radiology departments regarding practice at other hospitals and should assist in formulating the frequency and scope of appropriate MRI QA programmes.

The use of MRI has more than doubled over the last decade and, on average, 3000 scans per day were undertaken in England in 2004 [1]. There will be a variation in the quality of images produced in different centres, due to differences in equipment and scan parameters. Within a given centre, it is important to ensure that all medical images produced are consistent and of sufficient quality to answer the diagnostic question. A change in sensitivity of the system or an increase in the noise present in an image may cause fine detail to be lost. Image artefacts may affect the region of interest. These effects may be subtle, leaving the observer unaware that they are happening, yet can result in loss of confidence in the diagnosis or even misdiagnosis. It may be argued that the skilled observer would detect any decrease in quality of the radiological image. However, a case was reported in the national press several years ago where more than 1000 patients had to be recalled as a result of undetected image quality problems with an MRI scanner (‘‘Hospital’s brain scanner fails’’ Daily Telegraph, 6th June 2002). Quality assurance (QA) is a process to ensure that any product or service meets a required standard. This is particularly important for diagnostic imaging equipment, as it may not be immediately obvious that there have been any changes in the performance of the equipment. A QA programme for a MRI scanner must be able to detect changes in system performance, allowing equipment faults to be identified and rectified before they become clinically significant. This goal can only be achieved if the tests are of an appropriate type and range, and they are carried out at an appropriate frequency. Data from QA tests can be used to identify trends and hence anticipate deterioration in performance, and in some cases provide evidence to support equipment replacement business cases. The results also 592

Received 13 July 2005 Revised 7 October 2005 Accepted 24 October 2005 DOI: 10.1259/bjr/67655734 ’ 2006 The British Institute of Radiology

enable system drift to be monitored and quantified. This is particularly important when comparing follow-up scans on a patient, as it is necessary to ensure that any change in the appearance of the image is not due to changes in the equipment performance. The importance of undertaking appropriate measurements on new systems should also not be overlooked. Acceptance and commissioning tests allow a judgement to be made on whether the equipment meets the manufacturer’s specifications, as well as providing baseline performance data for future QA testing. McRobbie et al [2] found considerable value in undertaking acceptance tests on new MR installations, with signal to noise ratio and geometric linearity being the most common parameters to fail acceptance tests, and these were successfully corrected and improved in the majority of instances. For radiological equipment where ionizing radiation is used, not only are there national standards for QA tests to be undertaken [3, 4], but also the requirement to carry out these tests has been enshrined in legislation [5]. For MRI the situation is somewhat different. Comprehensive guidelines exist that detail a range of appropriate QA tests and their methodology [6]. However, there appears to be no consensus as to how frequently these tests should be undertaken in clinical practice, and currently there are no statutory requirements for these tests to be carried out. The purpose of this survey was to identify which MRI QA tests are currently undertaken in England, who undertakes the tests, and the frequency at which these tests are carried out. The results of this survey should assist radiology departments in the formulation of their own appropriate QA programme and contribute towards establishing which tests to undertake and their appropriate frequency. The British Journal of Radiology, July 2006

Survey of QA programmes

Method Within the National Health Service (NHS) there are over 200 MRI scanners in England, located in around 180 hospitals. From an alphabetical list of hospital trusts with MR scanners, 45 were selected by contacting every fourth hospital, enquiring if they would participate in this study. Of these, 35 hospitals agreed to take part in this survey, and a questionnaire was sent to each. Where hospitals had more than one MRI scanner, only one questionnaire was returned which reflected their typical QA programme. However, in practice this was similar for all the scanners within a particular hospital. For those hospitals that received medical physics support, a copy of the questionnaire was also sent to the relevant medical physics department. Follow-up calls were made to those hospitals that had not returned the survey to maximize the return rate. Information was requested on the nature and frequency of QA tests undertaken on the MRI equipment, as well as the type, field strength and age of the MRI equipment and also the staff group (radiographers, medical physicists, clinical engineering or service engineers) who undertook these tests.

Results From 35 hospitals who received the questionnaire, completed replies were received from 24, corresponding to a response of just over 68%. Responses were received in two further cases where the MRI scanner was undergoing replacement and, as routine QA was under review, the users felt unable to complete the questionnaire. Overall, this corresponded to a 74% response rate. The responses covered a broad range of equipment manufacturers, field strengths and ages (as shown in Figure 1) and all four major MRI manufacturers were represented to varying degrees. Out of the 24 hospitals, 79% undertook some form of QA (Figure 2a). Most routine QA was undertaken by the radiographers who operated the equipment on a regular basis. Service engineers have been excluded from this graph as it is understood that they will invariably undertake some QA during each service visit. Figure 2b compares the different tests performed by radiographers and medical physicists. Both staff groups performed a range of tests, with signal to noise ratio (SNR) measurements most commonly undertaken by radiographic staff. However, there was a considerable difference in the range of tests undertaken in different hospitals with only four hospitals carrying out all eight of the QA tests listed. Test frequencies for head coils and for other coils are compared between radiographers and medical physicists in Figure 2c. The head coil was the most frequently tested of all the coils, typically tested on either a daily or weekly basis by radiographic staff. The frequency of testing of the other coils was considerably less and this was further highlighted in Figure 2d which shows that where SNR measurements were performed, they were always undertaken on the head coil. Half of the hospitals extended this test to the body coil, but less than 13% included any further coils. The British Journal of Radiology, July 2006

Figure 1. Number of hospitals responding to the survey by (a) manufacturer, (b) field strength and (c) equipment age.

Figure 2e,f detail how often the other QA tests (i.e. non-SNR) are performed by radiographers and medical physicists, respectively. Image uniformity testing was carried out most often by both groups of staff. Slice position was the least frequent test to be carried out. The time interval between tests undertaken by the two groups of staff generally mirrored that of the SNR tests, with medical physicists tending to carry out a wider range of tests less often. Once again, there was a marked variation in practice between hospitals. The time interval between service visits varied between 1 month and 6 months, with 3 months being the most common (Figure 3a). At each service visit some 593

C J Koller, J P Eatough, P J Mountford and G Frain 80 Percentage of hospitals

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Figure 2. The percentage of hospitals that: (a) undertake quality assurance (QA), by staff group; (b) undertake QA by type of test, and by staff group; radiographers &medical physicists; (c) undertake signal to noise ratio (SNR) tests by frequency, by coil and by staff group; head coil-radiographers & head coil-medical physicists % other coils-radiographers other coils-medical physicists; (d) undertake SNR tests, by coil type; (e) undertake other QA tests, by type and by frequency – radiographers. Image Uniformity % Geometric Distortion & Spatial Resolution Ghosting Slice Thickness Slice Position. (f) undertake QA tests by type and by frequency – medical physicists. Image Uniformity % Geometric Distortion & Spatial Resolution Ghosting Slice Thickness Slice Position.

QA was routinely undertaken on both the head and body coils. However, the testing of the other coils was not as regular (Figure 3b).

Discussion One goal of undertaking QA is to maintain confidence in the performance of imaging equipment and to ensure that it is operating within specification. To that end, the measurement of SNR for an MRI unit gives a good, all round indication of the performance of the whole system. The SNR is determined by many different parameters, but for a particular sequence and coil combination, should remain stable over the long term 594

[2, 7, 8]. Hence a significant change in SNR will indicate a potential problem, but more specific tests are required to locate the exact cause. It should be noted, however, that testing the SNR on one coil does not itself provide information on the performance of any of the other coils. The other tests undertaken regularly, such as image uniformity, spatial resolution and slice thickness examine, amongst other things, field gradient strength and selection. Since many MRI systems have integral gradient coils, this aspect of the system may be sufficiently tested using just one receiver coil. SNR was the test most commonly performed in a QA program, typically on a daily to weekly basis. This test was always performed on the head coil at hospitals where QA was undertaken. Only half the hospitals The British Journal of Radiology, July 2006

Survey of QA programmes 30


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(b) Figure 3. Percentage of hospitals where service engineers undertake quality assurance (QA) tests: (a) by frequency; (b) by coil tested. , Head coil; &, body coil; %, other coils.

extended this test to the body coil, and even fewer hospitals (, 13%) tested the other coils available. Hence these hospitals are unable to make any judgment of the long-term performance of these coils. Another goal of QA is to maintain equipment calibrations, to ensure that there is no image distortion, and that any numerical quantities measured are accurate. The importance of this will depend to a great extent on how the diagnostic images are to be interpreted, and the type of diagnostic information required. Hence certain specific QA tests will be more appropriate to some hospitals [9]. Due to the complexity of MRI equipment, all hospitals received regular service visits on a 1–6 monthly basis. During these service visits, the service engineer may undertake some form of QA tests, but this will depend upon the work required and the availability of the MRI equipment. The particular tests undertaken will vary between manufacturers, engineers and between different systems, and may well not be appropriate for the required diagnostic information. Of note is that 21% of hospitals did not perform any QA measurements themselves, but left these tests to the service engineers. These hospitals did not receive service visits with any greater frequency than any of the other hospitals. This approach should only be undertaken with due caution as the users and the service engineers may have different goals. The user may wish to observe and minimize any drift in equipment performance whereas the engineer is primarily concerned with ensuring that the equipment is just operating within specification. In The British Journal of Radiology, July 2006

addition, the responsibility for equipment performance and safety rests primarily with the user and not with the engineers. This survey was designed, in part, to capture a representative sample of practice of MRI QA across hospitals in England. It should, of course, be borne in mind that the final response of 24 completed questionnaires from the random selection of 45 trusts might allow the possibility of bias if the response, be it positive or negative, was related to the extent of QA undertaken. There was no evidence, however, that this was the case; and, indeed, it is possible to argue both for and against a bias towards hospitals that undertake greater QA. Moreover, the possibility of bias would not affect the major finding of this study, namely that the range and frequency of MRI QA varies substantially between different hospitals. Quality is a core principle of both clinical governance and the NHS Plan [10]. In terms of medical imaging, this means the assurance that equipment is always working optimally in order to enable the quality and confidence of diagnosis to be maintained. However effective, QA is also about maintaining a balance between undertaking sufficient meaningful tests to ensure the equipment is operating optimally, whilst ensuring that the resource is used effectively in scanning patients. In the current climate, patient throughput is essential in maintaining waiting lists. Hence too much QA may also be deleterious to the diagnostic service.

Conclusion This national survey indicated that 21% of hospitals did not undertake any form of in-house QA on their MR scanners, and as such may be out of step with national practice. However, there was a wide variation in the number and frequency of QA tests carried out, and even for the most common QA test (SNR measurements) there was a wide variation in the frequency and range of coils tested. A national standard would help users to achieve the balance between maintaining quality and effective patient throughput.

Acknowledgments The authors wish to thank the MR staff and the service engineers who took part in this survey.

References 1. Imaging and radiodiagnostics (KH12). http://www.performance.doh.gov.uk/hospitalactivity. Department of Health. October 2004. 2. McRobbie DW, Quest RA. Effectiveness and relevance of MR acceptance testing: results of an 8 year audit. Br J Radiol 2002;75:523–31. 3. The Institute of Physics and Engineering in Medicine. Recommended standards for the routine performance testing of diagnostic X-ray imaging systems. IPEM Report 91. York: IPEM, 2005. 4. The Institute of Physics and Engineering in Medicine. Quality control of gamma cameras and associated computer systems. IPEM Report 66. York: IPEM, 1997.

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C J Koller, J P Eatough, P J Mountford and G Frain 5. The Ionising Radiations Regulations 1999. London: The Stationery Office, 2000. 6. The Institute of Physics and Engineering in Medicine. Quality control in magnetic resonance imaging. IPEM Report 80. York: IPEM, 2000. 7. Colombo P, Baldassarri A, Del Corona M, Mascara L, Strocchi S. Multicenter trial for the set-up of a MRI quality assurance programme. Magn Reson Imaging 2004;22:93–101.

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8. Lerski RA, De Certaines JD. Performance assessment and quality control in MRI by eurospin test objects and protocols. Magn Reson Imaging 1993;11:817–33. 9. Barker GJ, Tofts PS. Semiautomated quality assurance for quantitative magnetic resonance imaging. Magn Reson Imaging 1992;10:585–95. 10. The NHS Plan: a plan for investment, a plan for reform. London: The Stationery Office, 2000.



Influence of menopausal status and use of hormone replacement therapy on radiation dose from mammography in routine breast screening C J WHITAKER,

BSc, PhD,

C M KELLY,

BA,

K FAULKNER,

FIPEM, FInstP, FSRP

and E C STAMP,

MMathStat

Quality Assurance Reference Centre, 9 Kingfisher Way, Silverlink Business Park, Wallsend NE28 9ND, UK ABSTRACT. Menopausal status and hormone replacement therapy (HRT) cause alterations in breast structure which can affect mammographic image quality. Here we present the results of a study to discover the effect of menopausal status and HRT use on breast dose. Women attending routine screening completed questionnaires which included questions regarding menopausal status and HRT use. Details of the radiographic technique factors were recorded, from which the mean glandular dose (MGD) per film for each woman was calculated. MGD values were analysed with regard to the woman’s menopausal status and HRT use. The data from 516 women were analysed. Among the women who had never used HRT, women who had not undergone the menopause had a mean MGD of 2.94 mGy per film, whereas postmenopausal women had a lower mean MGD of 2.52 mGy per film: a difference which was found to be highly significant (p50.0045). Post-menopausal women who had never used HRT and those who had previously used HRT, but had ceased using it, had identical mean MGDs (2.54 mGy per film), whereas current HRT users had a significantly greater mean MGD (2.89 mGy per film, p50.003). Women currently using HRT receive a statistically significantly larger radiation dose from routine breast screening than other women. However, this effect is small and only occurs during the period of HRT use. Women who have ceased using HRT show no difference in MGD compared with women who have never taken HRT.

In recent years, there has been increasing interest in the effects of hormone replacement therapy (HRT) on women’s health [1–6]. HRT use has been increasing rapidly since its widespread introduction in the early 1970s [7, 8], with an estimated 50% of women currently aged 50 years and over having taken HRT at some time [9]. Two of the largest studies encompassing research into the effects of HRT on health are the Million Women Study (MWS) (recruitment: 1996–2001) in the UK [4, 9] and the Women’s Health Initiative (WHI) (recruitment 1993–1998) in the USA [3, 5]. One of the main differences between these studies is that the WHI actively enrolled women into a clinical trial of HRT, whereas the MWS simply recorded the woman’s HRT use, as prescribed by her GP. The MWS was designed to investigate relationships between the patterns of use of HRT and certain medical conditions, including breast cancer, in women attending routine screening as part of the NHS Breast Screening Programme (NHSBSP) [4, 10]. Since the MWS began in 1996, it has recorded the personal details of over one Address correspondence to: Dr Keith Faulkner. This study was partially supported by the European Commission’s Radiation Protection Research Programme, project DIMOND III (Measures for Optimising Radiological Information and Dose in Digital Imaging and Interventional Radiology), contract number FIGM-CT-2000-00061.


Received 1 September 2005 Revised 6 October 2005 Accepted 25 October 2005 DOI: 10.1259/bjr/68819456 ’ 2006 The British Institute of Radiology

million women [9], including specific details about previous medical conditions, family history of cancer, menopausal status and HRT use, by use of a selfadministered questionnaire. Although data analysis for the MWS is still ongoing, the study has already described some of the effects of HRT on breast screening and breast cancer [9, 11, 12]. As part of the MWS, women were tracked through the NHSBSP to determine rates of breast cancer, which were then analysed with regard to their original questionnaire details. It was found that current users of HRT were more likely to develop breast cancer than women who had never taken HRT and women who were past users of HRT [9]. Many other studies have also found a direct link between HRT use and increased incidence of breast cancer [1, 5, 13–15]. It was this HRT-induced increase in breast cancer which prompted the premature end to the WHI trial of combined oestrogen and progesterone HRT in 2002 [5]. With specific regard to the population studied in the MWS, i.e. women attending for mammograms as part of the NHSBSP, the direct effects of HRT on mammography need to be considered. This is because the clarity of a mammogram is affected by the tissue composition of the breast [16]. Fatty tissue has low absorption of X-rays, whereas ductal and glandular tissues absorb radiation and are, therefore, radiologically dense [17]. It is this ratio between fatty and dense breast tissue which 597

C J Whitaker, C M Kelly, K Faulkner and E C Stamp

determines overall breast density and which can be strongly influenced by both menopausal status and HRT use [16–19]. Despite this, the impact of menopausal status and HRT use on the radiation dose received by women during routine breast screening has not been established. The objectives of this study were to: 1. Determine the radiation dose received by 500 women during routine breast screening across five Breast Screening Units; 2. Investigate the effect that a woman’s menopausal status has on MGD from mammography; 3. Establish whether taking HRT affects radiation dose and whether this effect continues after the cessation of HRT use; 4. Compare the main types of HRT preparation to see what influence, if any, current HRT type has on mammographic radiation dose.

Methods and materials Structure of the study Women were recruited from five Breast Screening Units across the North East, Yorkshire and the Humber Government Office Regions of the UK between November 2002 and September 2003, with Units contributing between 13.2% and 31.6% of the women in the study (Table 1). Ethical approval was obtained from the Northern and Yorkshire Multi-Centre Research Ethics Committee, as well as from each specific local Research Ethics Committee for each Unit.

Every woman who attended a particular screening session was given the opportunity to take part in the study. If the woman agreed, she was given an information sheet explaining the study and asked to complete a consent form and a modified version of the MWS questionnaire [20] whilst at the screening centre. In total, 520 women across the region gave their consent to the study and completed the questionnaire. The questionnaire was based on the MWS questionnaire [20] and included detailed questions on the woman’s menopausal status and HRT use. However, in order that additional data on the woman could be obtained, the woman’s individual NHS Screening Number (Sx number) was also recorded on each questionnaire. Following the woman’s routine mammogram, details of the radiographic technique used for each film (such as projection, tube potential, tube current exposure time product and breast thickness) were recorded. The radiographic data, along with the X-ray tube output measurements from the mammography machine, were inputted into the NHSBSP Breast Dose Calculator program [21], to estimate the mean glandular dose (MGD) for each film. However, since the women had varying numbers of films taken each due to various factors including the introduction of two-view screening at three out of five of the Units in this study, the average MGD for each woman was calculated on a ‘‘per film’’ basis, giving each woman a mean MGD per film (mean MGD). There were only four women whose radiographic data was not sufficiently complete to calculate their MGD, leaving 516 women for inclusion in the data analysis.

Definition of groups for data analysis Table 1. Characteristics of women in the study as determined from the questionnaire Criterion

Number of Percentage women

Number of women in studya 516 Age (years) ,55 215 55–59 160 60–64 84 64+ 57 Menopausal statusb Pre-menopausal 30 Peri-menopausal 30 Post-menopausal 454 Unknown 2 HRT use Never 272 Previous 126 Current 106 Unknown 12 HRT typec Combined 45 Oestrogen only 31 Other/Unknown 30 Unit 1 68 2 91 3 99 4 163 5 95 a

100.0 41.7 31.0 16.3 11.0 5.8 5.8 88.0 0.4 52.7 24.4 20.5 2.3 42.5 29.2 28.3 13.2 17.6 19.2 31.6 18.4

Only women with a completed questionnaire, consent form and full mammographic screening data were included in data analysis. b Definitions of menopausal status are included in the Methods and Materials. c Current hormone replacement therapy (HRT) users only.

598

The mean MGD value for each woman was linked to her questionnaire responses via her individual Sx number. This then allowed mean MGD values to be analysed with regard to the woman’s menopausal status and details of her HRT use. Initially, women were divided into three different menopausal groups, pre-, peri- and post-menopausal, based mainly on whether or not their periods had stopped at the time of screening. In addition, women aged 53 years and over who had had either a hysterectomy without oophorectomy (removal of both ovaries) or who had started taking HRT before their natural menopause were defined as post-menopausal, along with women of any age who had had a bilateral oophorectomy. This categorisation was in line with guidance from the Million Women Study [12]. Subsequently, the pre- and peri-menopausal women were combined for comparison against the mean MGD of post-menopausal women. All data were analysed using a one-way ANOVA to assess variation and to establish statistical significance. As part of the questionnaire, women were asked whether they had ever taken HRT and, if so, whether they were currently taking HRT and which particular type. Their answers led to the post-menopausal women being categorised as never having taken HRT (never), previously having taken HRT (previous) or currently taking HRT (current). Mean MGD values were compared between the three groups and also for a combination of The British Journal of Radiology, July 2006

Influence of menopausal status and HRT on breast dose

never and previous users, against current HRT users. Finally, current HRT users were divided according to the type of HRT they were taking; combined oestrogen and progesterone, oestrogen-only or another type of preparation. Again, the mean MGDs for the women in each group were compared using a one-way ANOVA.

Results The characteristics of the women who took part in the study are outlined in Table 1. The women ranged in age from 42 years to 82 years, with a mean age of 57.0 years, the majority being aged less than 60 years (72.7%). The vast majority of the women’s questionnaires indicated that they were post-menopausal (88.0%), with the remainder being equally spilt between pre- and perimenopausal (5.8% each). Only two women (0.4%) who completed the questionnaires failed to give details of their menopausal status. Just under half of the women had taken HRT (44.9%) at some time, with 20.5% indicating that they were current users. As with menopausal status, only a small proportion of women (2.3%) did not specify whether they had ever taken HRT. For current HRT users, combined oestrogen and progesterone was taken by 42.5%, with oestrogen-only HRT comprising 29.2%. However, there was a large percentage of women (28.3%) who either used other forms of HRT (such as progesterone-only) or did not state the type of HRT on the questionnaire. This lack of knowledge about the type of HRT used may have led to difficulties in ascertaining any specific influence of the type of HRT on radiation dose (see Discussion). The mean MGD data from 516 women were analysed using a one-way ANOVA. The mean MGD per film for each woman was analysed by assigning the women to a group depending on their menopausal status and details of their HRT use, as obtained from their questionnaires and defined in the Methods and Materials section.

Influence of menopausal status Pre- and peri-menopausal women had similar mean MGD values of 2.84 mGy and 2.91 mGy, respectively

(Figure 1). Post-menopausal women had a lower mean MGD of 2.62 mGy, although this difference was not statistically significant (p50.08). If, however, post-menopausal women who have never used HRT are compared directly with pre- and peri-menopausal women who have also never used HRT (Figure 2), then the difference in MGDs becomes statistically significant (p50.0045). These non-HRT post-menopausal women have a mean MGD of 2.52 mGy, compared with a mean MGD of 2.94 mGy for non-HRT pre- and peri-menopausal women.

Influence of HRT use In post-menopausal women, women who had never used HRT and those women who had previously used HRT, but now did not, had practically identical mean MGDs of 2.54 mGy (Figure 2). However, current users of HRT had a greater mean MGD of 2.89 mGy, which was similar to the MGD for pre- and peri-menopausal women who had never taken HRT (2.91 mGy). This difference in MGD between current users of HRT and women not currently taking HRT was found to be highly statistically significantly (p50.003; Figure 3).

Influence of HRT type As described previously, the large proportion of women who did not specifically identify their current type of HRT in the questionnaire may have adversely affected the investigation as to whether the type of HRT influences MGD. For current HRT users who specified HRT type, there was no statistically significant difference in mean MGDs between combined (2.73 mGy) and oestrogen-only (2.88 mGy) HRT (Figure 4).

Discussion The data obtained from the women’s questionnaires was generally complete, with only 14 women (2.7%) not providing enough details to classify either their menopausal status or history of HRT use (Table 1). Overall, only four women (0.8%) were excluded from the data analysis due to lack of complete radiographic data.

Figure 1. Effect of menopausal status on mean glandular dose (MGD) per film. Error bars indicate 95% confidence levels from one-way ANOVA analysis.


599


Figure 2. Effect of hormone replacement therapy (HRT) use on mean glandular dose (MGD) per film of post-menopausal women (grey bars) and pre- and peri-menopausal women combined (dark grey bar). ‘‘Never’’ indicates women who have never used HRT, ‘‘Previous’’ indicates women who have taken HRT in the past, but have now ceased HRT use, and ‘‘Current’’ indicates women currently taking HRT. Error bars indicate 95% confidence levels from one-way ANOVA analysis.

The classification of women into three menopausal groups was carried out following the criteria used by the MWS [12]. The percentages of women in each category, pre-, peri- and post-menopausal, were 5.8%, 5.8% and 88% in the study reported here (Table 1) and 6.3%, 6.5% and 87% in the MWS [12], respectively. Classification of women based on HRT use was more straightforward than menopausal status and led to 45% of women being classified as a previous or current user of HRT. This figure is similar to the MWS where questionnaires indicated that almost half of women (47%) had used HRT at some time [4]. There is obviously close agreement between our study and the MWS with regard to the woman’s menopausal details and HRT use. In our study, there was a large proportion of current HRT users who were unable to accurately recall the name or type of HRT preparation they were currently taking. Combined oestrogen and progesterone accounted for 43% of current users, with oestrogen-only preparations accounting for 29% (Table 1). However, there was a large number of current HRT users classified as using either another or unknown type of HRT (28%). These figures are comparable with those obtained by the MWS. In the MWS, the majority of current HRT users took combined oestrogen and progesterone HRT (60%), with around a third taking oestrogen-only HRT (31%) and only 8.7% of women classed as other or unknown HRT type [9]. This large difference in the proportion of current HRT users classified as ‘‘unknown HRT type’’ between

our study and the MWS is likely to be due to the difference in questionnaire protocol. In the MWS, the women completed their questionnaire at home, where they were likely to have their current HRT to hand and so were able to identify the type accurately. However, since our questionnaire was completed at the screening unit, the women were less likely to be able to remember the exact name or type of preparation they were using. This led to difficulty in interpreting the radiation dose results with regard to HRT type for current HRT users. Although our results indicated a small (5%) difference in dose received between those women taking combined oestrogen and progesterone and those taking oestrogen-only HRT preparations (Figure 4), the lack of specific details about HRT type resulted in too few women to analyse the data meaningfully.

Radiation dose The mean radiation dose received by the 516 women in this study was 2.66 mGy. This compares with an average MGD of 2.23 mGy determined by Young et al [22] for women attending routine breast screening across the UK during 2001 and 2002. The small difference between the two mean MGDs is likely to be due to the use of Lorad systems (Lorad M-IV) in four out of five of the Units which took part in our study. Lorad systems are known

Figure 3. Effect of current hormone replacement therapy (HRT) use on the mean glandular dose (MGD) per film of post-menopausal women. Error bars indicate 95% confidence levels and *** indicates highly significant difference (p,0.005) from one-way ANOVA analysis. 600


Influence of menopausal status and HRT on breast dose

Figure 4. Effect of type of hormone replacement therapy (HRT) on the mean glandular dose (MGD) per film of current HRT users. ‘‘Combined’’ indicates combined oestrogen and progesterone users. Error bars indicate 95% confidence levels from one-way ANOVA analysis.

to give a higher MGD than other systems and, according to the Young et al study, account for only 15% of the systems used nationally [22]. Ideally, the best approach to estimating the glandular dose is to apply an accurate glandularity factor, specific to each woman. In the absence of detailed knowledge of breast glandularity for each woman, an average glandularity factor has been applied. As the glandularity between groups analysed here may vary, part of the observed differences between groups will be due to this effect. It is anticipated that the actual difference between groups will be smaller.

Influence of menopausal status on radiation dose Although the mean MGD for post-menopausal women was around 9% lower than those for pre- and perimenopausal women together, this difference was not statistically significant (p50.08) when HRT status was not taken into account, perhaps due to the HRT masking any real effect the menopause has on MGD. However, when women who had never taken HRT were analysed with regard to their menopausal status, the difference between this group of women and the post-menopausal women did reach statistical significance, with postmenopausal women being found to have received a statistically significantly smaller radiation dose (2.52 mGy) than women who had not yet undergone the menopause (2.94 mGy: p50.0045). The difference in radiation dose between these groups may reflect the changes which occur to the structure of a woman’s breasts after the menopause [18].

Influence of HRT use on radiation dose It was decided that only post-menopausal women would be analysed with regard to HRT use, since the numbers of pre- and peri-menopausal women were so small (60 in total) that it would severely limit any statistical analysis. It was also likely that the vast majority of non-menopausal women would never have taken HRT. In post-menopausal women, current users of HRT received an approximately 12% larger radiation dose The British Journal of Radiology, July 2006

than women who were not taking HRT at present (Figure 2). This difference was highly statistically significant (p50.003). There was no statistical difference in radiation dose received between women who had stopped taking HRT and those who had never taken it (Figure 2). This temporary nature of the effect of HRT on radiation dose is supported by research which has shown that the effects of HRT use, both on the incidence of breast cancer [9, 14] and on the specificity of mammography [11], decline rapidly after initial cessation and return to the levels of women who have never used HRT. These results indicate that the influence of HRT on radiation dose, breast cancer and mammograms only lasts for the length of time that the woman is taking HRT, with little long term effect. However, for current HRT users, the increased radiation dose received may have important implications. The difference between radiation dose received by a current user of HRT, versus a non-user is approximately 0.35 mGy per film, which equates to 0.7 mGy per routine two-view screen (four films). Although this additional amount of radiation received is relatively small, there may still be an impact when it comes to radiation-induced breast cancer. Law and Faulkner calculated the effect of specific radiation doses on the rates of induction of breast cancer [23]. For women aged 50–64 years, this equates to approximately 11.4 breast cancers per million women for each mGy of radiation. The small amount of additional radiation incurred as a result of taking HRT is outweighed by the substantial benefit of screening women on HRT, since this group of women is known to have greater background incidence of breast cancer than women not taking HRT [1, 5, 9, 11– 15]. Other factors, such as lack of specificity and selectivity of mammograms with HRT use [24, 25], are more likely to affect the benefit/risk ratio.

Conclusions Women currently using HRT receive a statistically significantly larger radiation dose from routine breast screening than other women. However, this effect is small and only occurs during the period of HRT use. The increase in mammography radiation dose for women taking HRT may result in a small increase in the number 601


of radiation-induced cancers, but this is probably justified due to the higher incidence of breast cancer in these women. Women who have ceased using HRT show no difference in radiation dose received compared with women who have never taken HRT.

Acknowledgments We wish to thank all the women who agreed to take part in the study and completed questionnaires, along with the staff at the five Breast Screening Programmes in the North East, Yorkshire and The Humber Government Office Region who were involved in this study. Thanks to Prof. V Beral, Ms B Crossley and Ms A Brown from the Cancer Research UK Epidemiology Unit, Oxford, for their initial advice about the questionnaire and subsequent computerized data extraction from the questionnaires. The authors are also grateful to Mrs C Clayton, Mr G McGill, Mr C Moore and Dr K J Robson for their help with the Medical Physics data.

References 1. Goddard MK. Hormone replacement therapy and breast cancer; endometrial cancer and cardiovascular disease: risks and benefits. Br J Gen Prac 1992;42:120–5. 2. Hulley S, Grady D, Bush T, Furberg C, Herrington D, Riggs B, et al. Randomized trial of estrogen plus progestin for secondary prevention of coronary heart disease in postmenopausal women. JAMA 1998;280:605–13. 3. Women’s Health Initiative Study Group. Design of the Women’s Health Initiative Clinical Trial and Observational Study. Control Clin Trials 1998;19:61–109. 4. The Million Women Study Collaborative Group: The Million Women Study: design and characteristics of the study population [peer-reviewed research]. 5. Writing Group for the Women’s Health Initiative Investigators. Risks and benefits of estrogen plus progestin in healthy postmenopausal women. Principal results from the Women’s Health Initiative randomized controlled trial. JAMA 2002;288:321–33. 6. Li C, Wilawan K, Samsioe G, Lidfeldt J, Agardh C-D, Nerbrand C. Health profile of middle-aged women: the Women’s Health in the Lund Area (WHILA) study. Human Repro 2002;17:1379–85. 7. Gambrell RD Jr. Studies of endometrial and breast disease with hormone replacement therapy. In: Studd JWW, Whitehead MI, editors. The menopause. Oxford, UK: Blackwell Scientific Publications, 1988:247–61. 8. Jolleys JV, Olesen F. A comparative study of prescribing of hormone replacement therapy in USA and Europe. Maturitas 1996;23:47–53.

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9. Million Women Study Collaborators. Breast cancer and hormone-replacement therapy in the Million Women Study. Lancet 2003;362:419–27. 10. http://cancerscreening.org.uk/breastscreen/ [Accessed 2 February 2006]. 11. Banks E, Reeves G, Beral V, Bull D, Crossley B, Simmonds M, et al. Impact of use of hormone replacement therapy on false positive recall in the NHS breast screening programme: results from the Million Women Study. BMJ 2004;328:1291–2. 12. Banks E, Reeves G, Beral V, Bull D, Crossley B, Simmonds M, et al. Influence of personal characteristics of individual women on sensitivity and specificity of mammography in the Million Women Study: cohort study. BMJ 2004;329:477. 13. Marsden J. Hormone-replacement therapy and breast cancer. Lancet Oncol 2002;3:303–11. 14. Coombs NJ, Taylor R, Wilcken N, Boyages J. Hormone replacement therapy and breast cancer: estimate of risk. BMJ 2005;331:347–9. 15. Ewertz M, Mellemkjaer L, Poulsen AH, Friis S, Sorensen HT, Pedersen L, et al. Hormone use for menopausal symptoms and risk of breast cancer. A Danish cohort study. Br J Cancer 2005;92:1293–7. 16. Warren R. Hormones and mammographic breast density. Maturitas 2004;49:67–78. 17. Sterns EE, Zee B. Mammographic density changes in perimenopausal and postmenopausal women: is effect of hormone replacement therapy predictable? Breast Cancer Res Treat 2000;59:125–32. 18. Miller WR, Anderson TJ. Oestrogens, progestogens and the breast. In: Studd JWW, Whitehead MI, editors. The menopause. Oxford, UK: Blackwell Scientific Publications, 1988:234–46. 19. Greendale GA, Reboussin BA, Sie A, Singh R, Olson LK, Gatewood O, et al. Effects of estrogen and estrogenprogestin on mammographic parenchymal density. Ann Intern Med 1999;130:262–9. 20. http://www.millionwomenstudy.org/[Accessed 2 February 2006]. 21. http://www.cancerscreening.nhs.uk/breastscreen/publications/mammography-equipment.html#me-iqad [Accessed 2 February 2006]. 22. Young KC, Burch A, Oduko JM. Radiation doses received in the UK Breast Screening Programme in 2001 and 2002. Br J Radiol 2005;78:207–18. 23. Law J, Faulkner K. Cancers detected and induced, and associated risk and benefit, in a breast screening programme. Br J Radiol 2001;74:1121–7. 24. Evans A. Hormone replacement therapy and mammographic screening. Clin Radiol 2002;57:563–4. 25. Carney PA, Miglioretti DL, Yankaskas BC, Lerlikowske K, Rosenberg R, Rutter CM, et al. Individual and combined affects of age, breast density, and hormone replacement therapy use on the accuracy of screening mammography. Ann Intern Med 2003;138:168–75.



Kodak EDR2 film for patient skin dose assessment in cardiac catheterization procedures R E MORRELL,

MSc, MIPEM

and A T ROGERS,

MSc, MIPEM

Medical Physics Directorate, Nottingham City Hospital NHS Trust, Hucknall Road, Nottingham NG5 1PB, UK ABSTRACT. Patient skin doses were measured using Kodak EDR2 film for 20 coronary angiography (CA) and 32 percutaneous transluminal coronary angioplasty (PTCA) procedures. For CA, all skin doses were well below 1 Gy. However, 23% of PTCA patients received skin doses of 1 Gy or more. Dose–area product (DAP) was also recorded and was found to be an inadequate indicator of maximum skin dose. Practical compliance with ICRP recommendations requires a robust method for skin dosimetry that is more accurate than DAP and is applicable over a wider dose range than EDR2 film.

Cardiac catheterization procedures can involve prolonged fluoroscopic imaging and large numbers of acquired images. As a result, patient skin doses may approach or exceed the threshold for deterministic skin effects [1–6]. The US Food and Drug Administration [7], and more recently the International Commission on Radiological Protection (ICRP), have published guidelines for dose minimization and for recording skin doses to patients who are suspected to be at risk. ICRP Report 85 [8] recommends that the magnitude and position of the maximum skin dose should be recorded in the patient’s notes if it exceeds 1 Gy for procedures that are likely to be repeated, or 3 Gy for all procedures. Patients who are considered to be at risk should receive appropriate information and clinical follow-up. Implementation of these guidelines requires a robust method for assessment of patient skin doses. Modern cardiac X-ray units are fitted with dose–area product (DAP) meters, which indicate the total amount of radiation incident on the patient’s skin. However, there is no simple relationship between DAP and maximum skin dose. The dose distribution depends on which imaging projections are used, and for what proportion of the procedure. This can vary greatly from one patient to the next depending on operator preference, the anatomy of the patient’s disease, and the complexity of the procedure. Whilst some authors have proposed DAP values to alert the operator to the potential for deterministic effects [9–12], others have reported poor correlations between DAP and maximum skin dose [13–15]. Slow radiographic film can be used to record a map of the skin dose distribution over a large area. As long as the film is not saturated, absolute dose measurements can be made directly from the film and any regions of high dose can be easily identified. Of the films that are compatible with standard radiology processors, the wide latitude films developed for portal imaging and quality control applications in radiotherapy are responsive to the highest radiation doses, and are thus most suitable for dosimetry in high dose diagnostic and interventional procedures [13, 16–18]. The British Journal of Radiology, July 2006

Received 22 April 2005 Revised 7 October 2005 Accepted 25 October 2005 DOI: 10.1259/bjr/78359708 ’ 2006 The British Institute of Radiology

The dosemeter selected for this study was Kodak EDR2 film (Eastman Kodak Company, Rochester, NY), which currently has the widest available dose range. Guibelalde et al [19] have successfully used it for skin dosimetry during interventional cardiology procedures. They reported its saturation point at 1.4 Gy, and found saturation to occur in about 1% of cases. The performance of the film has been characterized in detail, as described in a previous paper by the authors [20]. It is available in 35 cm 6 43 cm sheets, which are large enough to capture most of the radiation fields on the patient’s back. It is pre-wrapped in light-proof paper, ready for use. It can be processed in a standard radiology processor using non-glutaraldehyde chemicals. The purpose of the study was to determine typical skin doses for patients undergoing coronary angiography (CA) and percutaneous transluminal coronary angioplasty (PTCA) in our cardiac catheterization laboratory, and to estimate the percentage of patients receiving doses of 1 Gy or more. The film measurements were compared with DAP to determine whether DAP could be used as a predictor of maximum skin dose.

Method The film was previously calibrated over the range of exposure conditions typically encountered in our cardiac catheterization laboratory [20]. Its response was characterized across the dose range 20–1000 mGy. The effects of beam energy and filtration, field size, exposure rate, film batch and processing conditions were quantified. The relationship between dose (D) and optical density (OD) was found to be: 1 ODmax {OD D~{ ln a ODmax {ODmin

ð1Þ

The constant a had a value of 0.0027 mGy21 when a 20 cm polymethylmethacrylate (PMMA) phantom was used to simulate the backscatter from a patient and with the 603

R E Morrell and A T Rogers

processing conditions used in our department. The mean optical density of unexposed film (ODmin) was 0.21, whilst the mean density of fully exposed film (ODmax) was 3.92. The film saturated at around 1 Gy. The uncertainty in dose per optical density was estimated to be 229 mGy to +62 mGy at a fixed dose of 160 mGy, across the full range of exposure and processing conditions employed. At other dose levels, this interval was assumed to scale linearly with the gradient of the calibration curve, given by: dD eaD ~ d ðODÞ aðODmax {ODmin Þ

ð2Þ

Skin doses were measured for 20 CA and 32 PTCA procedures, performed on an Integris H5000F C-arm imaging unit (Philips Medical Systems, Best, Netherlands). The imaging equipment was subject to monthly and annual quality control checks, as recommended by the Institute of Physics and Engineering in Medicine [21]. Patients were selected sequentially, and the study included only those procedures performed by our in-house consultant cardiologists and the registrars working under their supervision. Fluoroscopy was performed using the ‘‘low continuous’’ factory setting, which has a nominal input dose rate at the detector of 740 nGy s21 and employs 0.4 mm copper filtration. All acquisition runs were performed on the ‘‘12.5 FPS Coronary’’ setting, which has a nominal detector input dose rate of 870 nGy s21 and has no copper filter. Before commencing each procedure, a sheet of 35 cm 6 43 cm EDR2 film was positioned on the imaging table, underneath the mattress. The dotted rectangle in Figure 1 demonstrates the position and orientation of the film. Its long axis was perpendicular to the long axis of the table, and its top edge was approximately level with the patient’s shoulders. Each film was labelled to indicate which side was face-up, and which edge was closest to the patient’s head. Following exposure, a pinhole was made in the corner of the film packet corresponding to the patient’s left shoulder, to identify the orientation of the processed film. All films were stored in their packets overnight before processing, since the response of EDR2 film takes some time to stabilize following exposure [20, 22]. They were then processed in a Kodak X-OMAT M6B processor (Eastman Kodak Company, Rochester, NY), with Photosol developer and fixer (Photosol Limited, Basildon, UK). The maximum optical density of each film was determined manually, using a Pehamed Densoquick 2 densitometer (Pehamed, Sulzbach, Germany). The DAP for each procedure was measured using an integral PTW-DIAMENTOR-M1 DAP meter (PTWFREIBURG, Freiburg, Germany). This had previously been calibrated over the same range of exposure conditions as the film. The uncertainty in its response was estimated at ¡13%. The Pearson correlation coefficient between DAP and peak skin dose was calculated for each procedure type, for those procedures where no film saturation occurred. The significance of the correlations was determined using Student’s t-test. 604

Figure 1. Film position and orientation.

Results Figure 2 shows a dosimetry film from a PTCA procedure, viewed as if looking at the patient’s back. The patient’s left shoulder is indicated by the black spot in the top left-hand corner of the image. The region of The British Journal of Radiology, July 2006

Kodak EDR2 film for patient skin dose assessment

angiography, and 0.61 for coronary angioplasty. These correlations were both significant (p 5 0.05). However, Figures 4 and 5 show that some of the points deviate considerably from the trend lines.

Discussion

Figure 2. A dosimetry film from a coronary angioplasty, viewed as if looking at the patient’s back. The black spot in the top left-hand corner indicates the patient’s left shoulder.

maximum dose can be readily identified by the darkest patch, in the top right-hand quadrant. One film from each procedure type showed abnormally large numbers of radiation fields, with extensive areas of film saturation. Since the DAP from these examinations was not particularly high, it seemed likely that these films had been left on the imaging table for more than one procedure. Both were excluded from the data analysis. Figure 3 shows the distribution of maximum skin doses for coronary angiography and angioplasty procedures. Skin doses for the angiograms were all less than 600 mGy. Seven films from angioplasty procedures were saturated in at least one region, implying a skin dose of 1 Gy or more. Figures 4 and 5 are scatter plots of maximum skin dose against DAP for the two procedure types, for films that demonstrated no saturation. A linear trend has been fitted to each. The error bars indicate the expected uncertainties in the measurements, as described in the method. The Pearson correlation coefficients between maximum skin dose and DAP were 0.76 for coronary

For coronary angiography, all skin doses were well below 1 Gy. This is in agreement with other published studies using thermoluminescence dosimetry or film [13–15, 23, 24]. The mean peak skin dose was 195 mGy, with a range from 70 mGy to 520 mGy. Patients undergoing these diagnostic investigations are unlikely to receive doses sufficient to cause deterministic effects. For coronary angioplasty, 23% of patients received skin doses of 1 Gy or more, sufficient to saturate the film. It must be assumed that a similar proportion of our patients approach or exceed this level in routine clinical practice. Since it is fairly common for patients to undergo more than one procedure, the dose to each of these patients should be assessed and recorded. Thermoluminescence dosimetry studies by Van de Putte et al, Waite and Fitzgerald, and Verdun et al all found some skin doses approaching or exceeding 1 Gy, for PTCA patients [14, 15, 24]. The increased incidence of film saturation compared with that reported by Guibelalde et al may be at least partly explained by our lower film saturation point of 1 Gy. Even if this is increased to 1.5 Gy using a dedicated processor as Guibelalde did, the film still cannot measure doses up to the 2 Gy threshold for deterministic effects. Given the large proportion of patients receiving skin doses of at least 1 Gy, it seems likely that some of them exceed the 2 Gy threshold for deterministic effects [25]. These patients should be monitored for skin effects, and informed about potential symptoms and appropriate action to take should any skin changes occur. The study clearly identifies a need for routine assessment of patient skin doses for coronary angioplasty procedures in our cardiac catheterization laboratory. EDR2 film identifies those patients whose doses may exceed 1 Gy, and who may therefore be at risk of deterministic skin effects. However, since this film saturates at 1 Gy, it cannot be used to assess these higher doses. It is labour-intensive as a dosimetry

Figure 3. Maximum skin doses for coronary angiography (CA) and percutaneous transluminal angioplasty (PTCA) procedures.


605


Figure 4. Maximum skin dose versus dose–area product (DAP), for coronary angiography (CA).

method, because each film must be individually positioned and labelled, processed and analysed to determine the maximum dose. A further limitation of film dosimetry is that contributions from lateral and very wide oblique views are not measured. DAP was found to be a poor indicator of maximum skin dose, because clinical practice varies so much from one procedure to the next. Although there was a significant correlation between DAP and maximum skin dose for both procedure types, it is evident from Figures 4 and 5 that DAP alone cannot not reliably predict high skin doses. Alternative dosemeters used in cardiac catheterization procedures include thermoluminescent dosemeters and scintillation detectors [14, 15, 24, 26]. A major disadvantage of such detectors is their small area. The location of the maximum skin dose is not usually known in advance and, if there is no detector at this exact location, skin dose can be grossly underestimated. There is now a growing range of ‘‘Gafchromic’’ films, produced by International Speciality Products (Wayne, New Jersey). These are designed specifically for patient dosimetry applications, have higher dose ranges than EDR2 film, and do not require any processing. At present however, they are prohibitively expensive for routine dosimetric use. A potential solution would be a mathematical model, to calculate skin dose distribution from the exposure parameters recorded in the DICOM image files for each

individual acquisition run. As well as enabling assessment of doses above the film’s saturation point, this would require less staff involvement than film dosimetry. The model could be applied routinely for all patients or run for individual patients whose DAP exceeds a certain trigger level.

Conclusion Skin doses in our cardiac catheterization laboratory are unlikely to exceed 1 Gy for CA. However, around 23% of our patients receive maximum skin doses of at least 1 Gy during PTCA procedures. DAP is not an adequate indicator of patient skin dose. Practical compliance with ICRP recommendations requires a method for routine assessment of skin doses that is more accurate than DAP, and is applicable over a wider dose range than EDR2 film. This may be achievable by means of a mathematical model.

Acknowledgments We would like to thank our cardiologists for agreeing to participate in this study, and our catheterization laboratory radiographers and nurses for their invaluable assistance with data collection. We are grateful to Prof. Alan Perkins for helpful discussions.

Figure 5. Maximum skin dose versus dose–area product (DAP), for coronary angioplasty (PTCA) procedures for which no film saturation occurred. 606


Kodak EDR2 film for patient skin dose assessment

References 1. Kawakami T, Saito R, Miyazaki S. Chronic radiodermatitis following repeated percutaneous transluminal coronary angioplasty. Br J Dermatol 1999;141:150–3. 2. D’Incan M, Roger H, le Boudec MCF, Souteyrand P. Radiodermatitis following cardiac catheterization. Arch Dermatol 1997;133:242–3. 3. Dehen L, Vilmer C, Humiliere C, Corcos T, Pentousis D, Ollivaud L, et al. Chronic radiodermatitis following cardiac catheterisation: a report of two cases and a brief review of the literature. Heart 1999;81:308–12. 4. Shope TB. Radiation-induced skin injuries from fluoroscopy. Radiographics 1996;16:1195–9. 5. Vanõ E, Goicolea J, Galvan C, Gonzalez L, Meiggs L, Ten JI, et al. Skin radiation injuries in patients following repeated coronary angioplasty procedures. Br J Radiol 2001;74:1023–31. 6. Vanõ E, Arranz L, Sastre JM, Moro C, Ledo A, Garate MT, et al. Dosimetric and radiation protection considerations based on some cases of patient skin injuries in interventional cardiology. Br J Radiol 1998;71:510–6. 7. Food and Drug Administration. Recording information in the patient’s medical record that identifies the potential for serious X-ray induced skin injuries following fluoroscopically guided procedures. Rockville, MD: Center for Devices and Radiological Health, 1995. 8. International Commission on Radiological Protection. Report 85: Avoidance of radiation injuries from medical interventional procedures. Ann ICRP 2000;30:45–7. 9. Hansson B, Karambatsakidou A. Relationships between entrance skin dose, effective dose and dose area product for patients in diagnostic and interventional cardiac procedures. Radiat Prot Dosim 2000;90:141–4. 10. Skinner CL. Implementation of IR(ME)R in cardiac angiography. In: Proceedings of UK Radiological Congress 2002; 2002 June 9–11; Birmingham, UK. London, UK: British Institute of Radiology, 2002. 11. Neofotistou V, Vanõ E, Padovani R, Kotre J, Dowling A, Toivonen M, et al. Preliminary reference levels in interventional cardiology. Eur Radiol 2003;13:2259–63. 12. McFadden SL, Mooney RB, Shepherd PH. X-ray dose and associated risks from radiofrequency catheter ablation procedures. Br J Radiol 2002;75:253–65. 13. Vanõ E, Gonzalez L, Ten JI, Fernandez JM, Guibelalde E, Macaya C. Skin dose and dose-area product values for


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interventional cardiology procedures. Br J Radiol 2001;74:48–55. van de Putte S, Verhaegen F, Taeymans Y, Thierens H. Correlation of patient skin doses in cardiac interventional radiology with dose-area product. Br J Radiol 2000;73:504–13. Waite JC, Fitzgerald M. An assessment of methods for monitoring entrance surface dose in fluoroscopically guided interventional procedures. Radiat Prot Dosim 2001;94:89–92. Geise RA, Ansel HJ. Radiotherapy verification film for estimating cumulative entrance skin exposure for fluoroscopic examinations. Health Phys 1990;59:295–8. Fajardo LC, Geise RA, Ritenoure RA. A survey of films for use as dosimeters in interventional radiology. Health Phys 1995;68:595–9. Vanõ E, Guibelalde E, Fernandez JM, Gonzalez L, Ten JI. Patient dosimetry in interventional radiology using slow films. Br J Radiol 1997;70:195–200. Guibelalde E, Vanõ E, Gonzalez L, Prieto C, Fernandez JM, Ten JI. Practical aspects for the evaluation of skin doses in interventional cardiology using a new slow film. Br J Radiol 2003;76:332–6. Morrell RE, Rogers A. Calibration of Kodak EDR2 film for patient skin dose assessment in cardiac catheterization procedures. Phys Med Biol 2004;49:5559–70. Institute of Physics and Engineering in Medicine. Report 77: Recommended standards for routine testing of diagnostic X-ray imaging systems. York, UK: IPEM, 1997. Childress NL, Rosen II. Effect of processing time delay on the dose response of Kodak EDR2 film. Med Phys 2004;31:2284–8. Delichas MG, Psarrakos K, Giannoglou G, MolyvdaAthanasopoulou E, Hatziioannou K, Papanastassiou E. Skin doses to patients undergoing coronary angiography in a Greek hospital. Radiat Prot Dosim 2005;113:449–52. Verdun FR, Capasso P, Valley JF, Schnyder P. Dose evaluation in fluoroscopy. Radiat Prot Dosim 1998; 80:139–41. Wagner LK, Eifel PJ, Geise RA. Potential biological effects following high x-ray dose interventional procedures. J Vasc Interv Radiol 1994;5:71–84. Hwang E, Gaxiola E, Vlietstra RE, Brenner A, Ebersole D, Browne K. Real-time measurement of skin radiation during cardiac catheterization. Cathet Cardiovasc Diagn 1998; 43:367–70.

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Randomized phase II study of GM-CSF to reduce mucositis caused by accelerated radiotherapy of laryngeal cancer 1

J J MCALEESE,

FRCR,

2

K M BISHOP,

BSc, RGN,

2

R A’HERN,

PhD

and 2J M HENK,

FRCR

1

Belvoir Park Hospital, Hospital Road, Belfast BT8 8JR and 2Head and Neck Unit, Royal Marsden Hospital, Fulham Road, London SW3 6JJ, UK ABSTRACT. Acute mucositis is dose-limiting in many accelerated radiotherapy schedules for head and neck cancer. Cytokines may be one means of reducing the severity of mucositis. A study was designed to assess the effect of subcutaneous molgramostin (granulocyte-macrophage colony stimulating factor; GM-CSF) injections on acute radiation morbidity in patients undergoing accelerated radiotherapy for laryngeal cancer. A prospective, randomized, observer-blind, controlled trial was conducted in 29 patients who were to receive radical radiotherapy over 3 weeks for early stage laryngeal cancer. Patients were randomized to receive 150 mg (,2 mg kg21) GM-CSF subcutaneously once daily for 14 days after the second week of radiotherapy, or no GM-CSF. Patients were assessed weekly for grade of mucositis, skin reactions and related parameters. The severity of mucositis was reduced in the GM-CSF arm (p,0.05). No other end-points reached statistical significance. Two patients failed to complete their courses of GM-CSF. Three developed influenza type symptoms and in one an allergic reaction was noted. There was no difference in tumour control rates. Subcutaneous GM-CSF reduced the severity of mucositis in patients undergoing accelerated radiotherapy. Injections were well tolerated. Further studies of cytokines are warranted, to assess the feasibility of increasing the total doses of accelerated radiotherapy given, with the aim of improving tumour cure rates.

When head and neck cancers are treated with accelerated radiotherapy, acute normal tissue effects can be dose limiting, especially when concurrent chemotherapy is given. This is because severe mucositis can lead to consequential late damage [1]. Reduction of the severity of mucositis could permit dose escalation with an increased chance of tumour cure. The pathogenesis of mucositis is death of the actively dividing cells of the mucosal epithelium. The first clinical signs usually appear on the 12th or 13th day of radiotherapy. Healing takes on average 3 weeks after conventionally fractionated radiotherapy, but is prolonged in cases of severe mucositis. Various methods have been tried to reduce mucosal damage. Antiseptic and anti-inflammatory mouthwashes have no effect in reducing the severity of mucositis [2, 3]. Overgrowth of yeasts and aerobic gram negative cocci has been demonstrated, but studies of the use of local antifungals and antibiotics have given conflicting results; some suggest a small benefit, while others are negative [4–6]. Another approach is the use of growth factors to stimulate repair and proliferation of surviving mucosal cells. Granulocyte-macrophage colony stimulating factor (GM-CSF) influences the proliferation and differentiation of stem cells and regulates several functions in mature leukocytes, macrophages and dendritic cells of the submucosa and dermis [7, 8]. GM-CSF administered to Address correspondence to: Dr J J McAleese, 46 Onslow Gardens, Belfast BT6 0AQ, UK.

608


patients suffering cuts and burns decreases the healing period significantly [9]. GM-CSF has been successfully used to promote healing of chemotherapy-induced mucositis [10, 11]. It is therefore appropriate to investigate the use of GM-CSF to reduce the severity and increase the rate of healing of mucositis in patients with head and neck cancer. Several studies carried out to date have used GM-CSF with conventionally fractionated radiotherapy [12, 13]. Accelerated radiotherapy is used routinely for a variety of sites and stages of head and neck cancer, one of which is early laryngeal cancer. A large multicentre trial compared short (3–4 weeks) and long (5–6 weeks) treatment times for radiotherapy of laryngeal carcinoma. The short times gave equal local control and less late morbidity [14]. Since then, a 3-week 16-fraction regimen has been adopted as standard at the Royal Marsden Hospital. The majority of patients treated with this regimen develop symptomatic mucositis; grade 2 is observed in most, and grade 3 in about 10%. This group of patients was therefore deemed to be especially suitable for a study of the effect of GM-CSF with accelerated radiotherapy. A possible disadvantage of growth factors is risk of stimulation of cell division in tumour cells, leading to accelerated repopulation during radiotherapy and a reduced chance of tumour cure. In one study another cytokine, G-CSF, was associated with reduced tumour control [15] when given from days 15 to 19 of a 38-day course of radiotherapy, i.e. during the first half of the course. We decided to try the effect of GM-CSF daily for The British Journal of Radiology, July 2006

GM-CSF to reduce mucositis

2 weeks starting on day 14 of a 21-day course of radiotherapy. Accordingly, the drug was not given until two-thirds of the radiation dose had been accumulated, so that the risk of accelerated repopulation should be minimized. The optimal dose of GM-CSF is not known. The dose most often used to treat oral mucositis is 4 mg kg21, but other studies suggest 1 mg kg21 may be effective [13, 16]. Commercially available vials of GM-CSF contain 150 mg, sufficient for a dose of 2 mg kg21 for most patients, so this latter dose was chosen for our study.

Materials and methods Study design The study was a prospective, randomized, observerblind phase II trial. Approval was granted by the Royal Marsden Ethics Committee. Patients treated by radiotherapy for early glottic carcinoma were randomly allocated to receive or not daily injections of GM-CSF for 2 weeks, beginning on day 15 of the course of radiotherapy. A placebo injection was not used because it was not considered ethically justifiable. Accordingly, assessments of radiation effects were made and recorded by a blinded observer.

Eligibility criteria Patients were eligible for the study if they had histologically proven T1 N0 or T2 N0 glottic carcinoma and were to be treated with radiotherapy using a 16fraction 3-week regimen. They were required to be of WHO (World Health Organization) performance status grade 0 or 1. Patients were excluded if they had evidence of renal or hepatic impairment, serious infections requiring antibiotic therapy, were taking or likely to need corticosteroids, or were known to be allergic to GM-CSF.

Radiotherapy External beam radiotherapy was given with a linear accelerator using 6 MV photons. The technique was to use lateral parallel-opposed fields, except in patients with a short neck, in whom anterior oblique fields with wedge filters were used. In most cases the fields were 5 cm square centred on the vocal cord, but in T2 cases the fields were extended by up to 2 cm either above or below the vocal cord to cover supraglottic or subglottic extension, respectively. All doses were prescribed at the ICRU intersection point. Once-daily fractions of 3.125 Gy were delivered to a total dose of 50 Gy in 16 fractions in 21 days.

GM-CSF GM-CSF was administered at a dose of 150 mg by subcutaneous injection once daily for 14 days, beginning at the end of the second week of radiotherapy. It was The British Journal of Radiology, July 2006

therefore given during the final week of radiotherapy, and for 1 week after the end of radiotherapy.

Patients All patients meeting the eligibility criteria treated at the Royal Marsden Hospital between September 1997 and October 2000 were offered entry to the study. 29 agreed to enter, and were randomly assigned to the active or control arms. The characteristics of the two groups are compared in Table 1. There was an imbalance in stage distribution, with more T2 patients in the GMCSF arm. Consequently more of this group were treated with larger fields. The mean field sizes were 32.7 cm2 in the GM-CSF group and 30.6 cm2 in the control group. The two groups were well balanced for all other variables.

Assessment and follow-up Patients were seen weekly during treatment and for 3 weeks after treatment or until acute reactions settled. At each visit one of two independent observers, blinded to group allocation, scored mucositis by the Radiation Therapy Oncology Group (RTOG) system reproduced in Table 2 [17] using indirect laryngoscopy or flexible nasendoscopy to provide adequate visualization; data on skin erythema, moist and dry desquamation (as % of treatment field), pain on swallowing, severity of dysphagia, analgesic usage, evidence of candida infection and laryngeal oedema were collected at each visit. Weight was measured before starting and 2 weeks after the end of radiotherapy.

Statistical methods It was anticipated that 60% of patients receiving the 3 week course of radiotherapy would develop one of the more severe grades of mucositis. We decided that in order for the routine use of GM-CSF to be worthwhile, this figure needed to be reduced to 10%. In order to demonstrate this difference at 90% power and 5% significance, 17 patients in each group would be needed. Accordingly, the intention was to recruit 34 patients to Table 1. Pre-treatment characteristics of the study population Variable

Gender Male Female Age range (years) Median age (years) Smoking Current smoker Ex-smoker Unknown Stage I Stage II Total

GM-CSF arm

Control arm

14 1 48–79 60

12 2 32–70 65

9 5 1 6 9 15

9 4 1 11 3 14

GM-CSF, granulocyte-macrophage colony stimulating factor.

609

J J McAleese, K M Bishop, R A’Hern and J M Henk

the study. Data were analysed on an intention to treat basis. A ranking method (Mann Whitney U-test) was chosen to analyse mucositis scores, as the grading system employed uses an ordinal scale. Patients were ranked on the basis of the frequency of maximal mucositis score over the study period. The other data were also analysed by non-parametric methods. Local control, disease-free survival and overall survival rates were computed by the Kaplan-Meier method and analysed by the log-rank method.

Results Recruitment to the study proved to be disappointingly slow. There was a high refusal rate, because many patients were unwilling to receive a course of injections for a possible benefit only in terms of the side-effects of treatment. Eventually the study was terminated when 29 patients had been randomized. 15 patients were allocated to the GM-CSF treatment arm. Two patients failed to complete their prescribed course of injections. One patient discontinued after his second injection because of influenza-like symptoms. The second patient was admitted to a local hospital because of a chest infection and had his GM-CSF discontinued. All 29 patients completed their radiotherapy course as planned. Figures 1 and 2 show the proportions of patients in each group with each grade of mucositis at each week after starting radiotherapy. The overall areas of the graphs are similar, but the arms differ in the relative contributions of each grade. The number of patients in the GM-CSF arm with maximum mucositis scores of grade 0, I, II and III were 1, 4, 10 and 0, respectively. The corresponding numbers for the control arm were 1, 0, 12 and 1, respectively. There were a greater proportion of maximum grade I scores in the GM-CSF arm compared with the control arm (27% vs 0%) but decreased grade II

Figure 2. Proportion of patients with mucositis grades for granulocyte-macrophage colony stimulating factor (GM-CSF) arm.

scores (67% vs 86%), and the only patient with a grade III score was in the control arm. The difference between the mucositis grades of the two groups was significant (ranking by Mann Whitney U556.5, p,0.05) The mean time to healing of mucositis in the GM-CSF group was approximately 60% that of the control arm, but this did not reach significance on the log rank test (p50.25). One patient in the control arm required tube feeding. Otherwise, no patients had dysphagia at grade 3 (fluids only) or greater at 5 weeks. No differences were detected in dysphagia, odynophagia, analgesic usage, candida infection or laryngeal oedema. There was a difference noted in the average erythema grade in favour of the GM-CSF group (0.83 vs 1.33, p,0.05). No difference in moist desquamation or dry desquamation was noted. No statistically significant difference in weight change was noted between the groups.

Side effects of GM-CSF 12 patients who received GM-CSF had elevated white cell counts (WCC). The range of maximal WCC was 7.2–30.5 (median 19.7). All WCC had returned to normal within 3 weeks of completing injections (median 2 weeks). Three patients developed influenza-like symptoms with the GM-CSF and in one patient the injections were stopped because of this symptom. One patient developed an erythematous rash at his injection sites after completing his course of 14 injections (Figure 3). He had a past history of allergy to radiographic contrast medium.

Tumour outcomes

Figure 1. Proportion of patients with mucositis grades for control arm.

610

The median follow-up is 28 months (range 6–45 months) during which time there have been three relapses in the control arm. One occurred at the primary site at 36 months and was successfully salvaged with The British Journal of Radiology, July 2006

GM-CSF to reduce mucositis Table 2. RTOG mucositis scores RTOG Grade

Grade 0 Grade 1 Grade 2

Grade 3 Grade 4

Figure 3. Allergic reaction at granulocyte-macrophage colony stimulating factor (GM-CSF) injection sites.

laryngectomy. The other two patients had cervical node metastases, one at 10 months post-radiotherapy, the other at 13 months, and both underwent successful neck dissection. One patient in the GM-CSF arm developed a local recurrence and required a salvage laryngectomy at 12 months post-treatment. One patient in the GM-CSF group died of a myocardial infarction 5 months after his radiotherapy. The study is too small to draw conclusions about relative relapse rates, but there is no suggestion that GM-CSF adversely affected outcome. Three patients in the control arm were diagnosed with a second malignancy. Two developed non-small cell lung cancer, one of which died within 3 months; the other received a course of radical radiotherapy. The third patient was diagnosed with small cell lung cancer and has received second line chemotherapy. One patient in the GM-CSF group developed a carcinoma of the rectum, which was treated with surgery alone.

Discussion A meta-analysis of prophylactic GM-CSF in chemotherapy and radiotherapy induced mucositis [18] has shown a benefit with a relative risk of 0.51 (95% confidence intervals 0.29–0.91). Several studies have previously examined the ability of GM-CSF to modify radiation-induced mucositis (Table 3) but using differing radiotherapy regimens, doses of growth factor, and scoring systems for mucosal reactions. Five prospective randomized controlled trials of GM-CSF have been reported; in two, the agent was administered subcutaneously, and in the other three as a mouthwash. Throuvalas et al [13] reported a randomized trial in 10 patients receiving conventionally fractionated radiotherapy. GM-CSF was administered subcutaneously at a dose of 1 mg kg21 daily, beginning at the start of the third week of radiotherapy. The GM-CSF arm showed better pain control (p50.004) and less erythema (p50.09). Makonnen et al [12] conducted a randomized controlled trial comparing subcutaneous GM-CSF 150–300 mg per day plus sucralfate mouthwash versus sucralfate mouthwash control, in 40 patients receiving either The British Journal of Radiology, July 2006

Descriptor

No change over baseline Injection / may experience mild pain not requiring analgesia Patchy mucositis that may produce an inflammatory serosanguinous discharge / may experience moderate pain requiring analgesia Confluent mucositis / may include severe pain requiring narcotic Ulceration, haemorrhage or necrosis

conventionally fractionated or hyperfractionated radiotherapy with break. GM-CSF was begun at the end of the first week of radiotherapy and continued throughout the course of treatment. No difference was found in degree or frequency of mucositis in the overall group or by type of fractionation. Skin reactions, fever and bone pain were the most common side effects in the GM-CSF arm. Sprinzl et al [19] conducted a randomized control trial in 35 patients using GM-CSF mouthwash. All patients had advanced head and neck cancer and were treated with two cycles of radiotherapy (each 30 Gy in 15 fractions) and concomitant mitomycin C and 5FU chemotherapy. Data were collected during the first radiotherapy cycle. No differences in mucositis or oral pain were noted. Saarilathi et al [20] compared GM-CSF with sucralfate mouthwash in a randomized study in 40 patients with oral cancer receiving post-operative radiotherapy. Mucositis tended to be less severe and heal more quickly in the GM-CSF treated group (p50.07) and there were fewer requirements for opioid pain relief (p,0.05). Two case-control studies of subcutaneous GM-CSF have been reported. Kannan et al [16] compared the results of 10 patients treated with 1 mg kg21 body weight GM-CSF and conventionally fractionated radiotherapy with historic controls. None of the 10 patients developed grade III mucositis compared with 49% of controls. Itching in 20% and body aches in 40% were the main side effects, but both were successfully treated with simple oral medications. Wagner et al [21] compared the use of subcutaneous GM-CSF 5 mg kg21 body weight in 16 patients receiving conventionally fractionated radiotherapy with historical controls. Pain relief was improved in the GM-CSF group (p50.011). A trend towards a decreased degree of mucositis was noted. Only mild reactions to GM-CSF were reported. All of the above studies were too small to provide a conclusive result, but overall there is a trend towards a reduction in severity of radiation mucositis with GMCSF. Our study also showed a reduction in the severity of mucositis experienced by patients receiving accelerated radiotherapy. This was despite an imbalance in the two arms so that more patients in the GM-CSF arm were treated with larger fields. Side effects to GM-CSF seem to be generally mild. Kannan et al [16] reported itching in 20% and body aches in 40% of patients receiving the drug. Wagner et al [21] observed bone pain in 3% and pyrexia in 3%. Makonnen et al [12] reported local skin reaction in 13%, fever in 6% and bone pain in 5%. In our study, 20% had myalgia/flulike symptoms and one patient had an allergic reaction. 611

J J McAleese, K M Bishop, R A’Hern and J M Henk Table 3. Published studies of GM-CSF in radiation mucositis Author

Year

Number

Random

Difference

Dose

Throuvalis [13] Kannan [16] Wagner [21] Makonnen [12] Sprinzl [19] Saarilhati [20]

1995 1997 1999 2000 2001 2002

10 10 32 40 35 40

Yes No No Yes Yes Yes

Pain p50.004 Mucositis Pain p50.0011 No No Opiod use p50.042

1 mg kg21 1 mg kg21 5 mg kg21 150–300 mg Mouthwash Mouthwash

GM-CSF, granulocyte-macrophage colony stimulating factor.

Subcutaneous GM-CSF therefore appears to be well tolerated. Maurer-Schulze et al [22] suggest that low doses of GM-CSF stimulate cell proliferation in human tumour xenografts in nude mice, but inhibit proliferation at higher doses. The data in Makonnen’s paper [12] show a non-significant trend for decreased survival in the GM-CSF arm. Because of the possible potential for accelerating tumour cell re-population [23, 24], GM-CSF was not administered until the final week of radiotherapy in our study. We did not observe a difference in local control, but larger studies would be needed to confirm this finding.

Conclusions GM-CSF was well tolerated and no decrease in local control was seen. A decrease in severity of mucositis, consistent with the biological actions of GM-CSF as a promoter of wound healing, has been shown. GM-CSF appears promising as a potential modifier of radiotherapy-induced mucositis, but larger studies are required to establish its value. More specific growth factors are now under investigation: for example, Do¨rr and colleagues have demonstrated a radio-protective effect of keratinocyte growth factor on mouse oral mucosa [25]. Cytokines are a promising method of improving therapeutic ratio in radiotherapy of head and neck cancer, especially with accelerated regimes in which acute normal tissue effects are dose-limiting.

References 1. Dore W, Hendry JH. Consequential late effects in normal tissues. Radiother Oncol 2001;61:223–31. 2. Foote RL, Loprinzi CL, Frank AR, O’Fallon JR, Gulavita S, Tewfik HH, et al. Randomised trial of a chlorhexidine mouthwash for alleviation of radiation induced mucositis. J Clin Oncol 1994;12:2630–3. 3. Samaranyake LP, Robertson AG, MacFarlane TW, Hunter IP, MacFarlane G, Soutar DS, et al. The effect of chlorhexidine and benzdamine mouthwashes on mucositis induced by therapeutic irradiation. Clin Radiol 1988;39:291–4. 4. El-Sayed S, Nabid A, Shelley W, Hay J, Balogh J, Gelinas M, et al. Prophylaxis of radiation-associated mucositis in conventionally treated patients with head and neck cancer: a double-blind, phase III, randomised, controlled trial evaluating the clinical efficacy of an antimicrobial lozenge using a validated mucositis scoring system. J Clin Oncol 2002;20:3956–63. 5. Symonds RP, McIlroy P, Khorrami J, Paul J, Pyper E, Alcock SR, et al. The reduction of radiation mucositis by selective decontamination antibiotic pastilles: a placebo controlled double-blind trial. Br J Cancer 1996;73:312–7.

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6. Wijers OB, Levenda PC, Harms ER, et al. Mucositis reduction by selective elimination of oral flora in irradiated cancers of the head and neck: a placebo-controlled doubleblind randomised study. Int J Radiat Oncol Biol Phys 2001;50:341–52. 7. Berdel WE, Danhauser-Riedl S, Steinhauser C, Winton EF. Various human haemopoietic growth factors (IL-3, GM-CSF, G-CSF) stimulate clonal growth of hamatopoietic tumour cells. Blood 1990;73:80–3. 8. Dexter M. Haemopoietic growth factors. Review of biology and clinical potential. Macclesfield, Cheshire: GardinerCaldwell Ltd., 1990. 9. Masucci G. New clinical applications of granulocytemacrophage colony-stimulating factor. Med Oncol 1996;13: 149–54. 10. Chi KH, Chen CH, Chan WK, Chow KC, Chen SY, Yen SH, et al. Effect of granulocyte-macrophage colony-stimulating factor on mucositis in head and neck cancer patients after cisplatin, fluorouracil and leucovorin chemotherapy. J Clin Oncol 1995;13:2620–8. 11. Hejna M, Kostler WJ, Raderer M, Steger GG, Brodowicz T, Scheithauer W, et al. Decrease of duration and symptoms in chemotherapy-induced oral mucositis by topical GM-CSF: results of a prospective randomised trial. Eur J Cancer 2001;37:1994–2002. 12. Makkonen TA, Minn H, Jekunen A, Vilja P, Tuominen J, Joensuu H, et al. Granulocyte macrophage-colony stimulating factor (GM-CSF) and sucralfate in prevention of radiation-induced mucositis: a prospective randomised study. Int J Radiat Oncol Biol Phys 2000;46:525–34. 13. Throuvalis N, Antonadou D, Pulizzi M, Sarris G. Evaluation of the efficacy and safety of GM-CSF in the prophylaxis of mucositis in patients with head and neck cancer treated with radiotherapy. Eur J Cancer 1995;Suppl. 5: Abstract 431. 14. Wiernik G, Alcock CJ, Bates TD, Brindle JM, Fowler JF, Gajek WR, et al. Final report of the second British Institute of Radiology fractionation study: short vs long overall treatment times for radiotherapy of carcinoma of the laryngopharnx. Br J Radiol 1991;64:232–41. 15. Staar S, Rudat V, Stuetzer H, Dietz A, Volling P, Schroeder M, et al. Intensified hyperfractionation accelerated radiotherapy limits the additional benefit of simultaneous chemotherapy – results of a multicentric randomised German trial in advanced head and neck cancer. Int J Radiat Oncol Biol Phys 2001;50:1161–71. 16. Kannan V, Bapsy PP, Anantha N, Doval DC, Vaithianathan H, Banumarthy G, et al. Efficacy and safety of granulocyte macrophage-colony stimulating factor (GM-CSF) on the frequency and severity of radiation mucositis in patients with head and neck carcinoma. Int J Radiat Oncol Biol Phys 1997;37:1005–10. 17. Cox JD, Stetz J, Pajak TF. Toxicity criteria of the Radiation Therapy Oncology Group (RTOG) and the European Organization for Research and Treatment of Cancer (EORTC). Int J Radiat Oncol Biol Phys 1995;31:1341–8. 18. Clarkson JE, Worthington HV, Eden OB. Interventions for preventing oral mucositis for patients with cancer receiving


GM-CSF to reduce mucositis treatment (Cochrane Review). In: The Cochrane Library, Issue 2, 2004. 19. Sprinzl GM, Galvan O, de Vries A, Ulmer H, Gunkel AR, Lukas P, et al. Local application of granulocyte-macrophage colony stimulating factor (GM-CSF) for the treatment of oral mucositis. Eur J Cancer 2001;37:2003–9. 20. Saarilahti K, Kajanti M, Joensuu T, Kouri M, Joensuu H. Comparison of granulocyte-macrophage colony-stimulating factor and sucralfate mouthwashes in the prevention of radiation-induced mucositis: a double-blind prospective randomised phase III study. Int J Radiat Oncol Biol Phys 2002;54:479–85. 21. Wagner W, Alfrink M, Haus U, Matt J. Treatment of irradiation-induced mucositis with growth factors (rhGMCSF) in patients with head and neck cancer. Anticancer Res 1999;19:799–804.


22. Maurer-Schultze B, Bassukas ID, Hofmockel G. Effect of biological response modifiers on growth and cell proliferation of human tumour xenografts in nude mice. Cell Mol Biol 1995;41:65–78. 23. Foulke R, Marshal M, Trotta P, Van Hoff D. In vitro assessment of the effects of granulocyte-macrophage colony-stimulating factor on primary human tumours and derived lines. Cancer Res 1990;50:6264–7. 24. De Riese W, Allhoff EP, Werner M, Strief CG, Liedke S, Kircher H, et al. Effects of cytokines on clonogenic growth in vitro of primary human renal cell carcinoma. Urological Res 1992;20:369–73. 25. Do¨rr W, Spekl K, Farrell CL. Amelioration of acute oral mucositis by keratinocyte growth factor: fractionated irradiation. Int J Radiat Oncol Biol Phys 2002;54: 245–51.

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PICTORIAL REVIEW

Magnetic resonance urography: a pictorial overview 1

R GARCIÁ-VALTUILLE, 3 M C ARGUËLLO, MD

MD,

2

A I GARCIÁ-VALTUILLE,

MD,

1

F ABASCAL,

MD,

1

L CEREZAL,

MD

and

1

Instituto Radiolo´gico Cańtabro, Clıńica Mompıá, Avenida de los Condes, s/n. 39108 Santa Cruz de Bezana (Cantabria), 2Department of Pathology, Clıńica Mompıá, Santa Cruz de Bezana (Cantabria) and 3Department of Oncology, Clıńica Mompıá, Santa Cruz de Bezana (Cantabria), Spain ABSTRACT. Magnetic resonance urography (MRU) can be performed on the basis of two different imaging strategies: static-fluid MRU, based on heavily T2 weighted turbo spin echo (TSE) sequences, and gadolinium-enhanced excretory MRU. Both MR urographic techniques in combination with standard MRI permit a comprehensive examination of the entire urinary tract. This pictorial review illustrates the MRU features of the a wide spectrum of pathological conditions affecting the urinary tract.

Received 6 February 2005 Revised 10 May 2005 Accepted 23 May 2005 DOI: 10.1259/bjr/21075982 ’ 2006 The British Institute of Radiology

Introduction Magnetic resonance urography (MRU) is an emerging technique of MRI which provides a non-invasive visualization of urinary tract. Most of previous studies have used the unenhanced, heavily T2 weighted pulse sequences to obtain images in which static fluid exhibits a higher signal intensity relative to background (static MRU) [1–4]. Clinical urography requires both morphological and functional information about the kidneys and the collecting system. However, these sequences do not provide information about the renal excretory function. MRU performed with contrast material can meet all demands of clinical urography and, in some cases, could replace conventional X-ray urography [1, 4, 5]. This pictorial essay reviews the MRU features of the major urinary tract disorders in which static or excretory MRU provides information of diagnostic value.

Technique The images were acquired by a 1 T superconducting magnet (New technology [NT] Gyroscan; Philips Medical Systems, Best, The Netherlands) using a body coil. The MR sequence protocol started with localizing T1 weighted gradient-echo sequence (repetition time (TR) 18 ms/echo time (TE) 6.9 ms; flip angle 30 ˚; scan time 3 min 30 s) and T2 weighted turbo spin echo (TSE) sequence (TR 4200 ms/TE 100 ms; flip angle 90 ˚; scan time 3 min 12 s) in axial and coronal planes. In static MRU, heavily T2 weighted TSE pulse sequences are used to obtain water images of the urinary tract (three-dimensional; respiratory-triggering; TR 2000 ms/TE 700 ms; flip angle 90 ˚; TSE-factor 101; matrix size 2566256; number of excitations 2; field of view (FOV) 360–390 mm; number of slices 40–50; slice thickness 2 mm; scan time 3 min 30 s to 4 min 20 s). 614

Before the acquisition of the excretory MR urographic sequences, the patients received an intravenous dose of 0.1 mg kg21 of furosemide and 0.1 mmol kg21 of GdDTPA-BMA (Gadodiamide). A delay of 1–5 min between the administration of both drugs is necessary for achieving optimal contrast enhancement of the urinary tract. Excretory MRU was performed at our institution using a respiratory gating, three-dimensional, T1 weighted gradient-echo sequence (TR 15 ms/TE 5 ms; flip angle 70 ˚; matrix size 2566256; number of excitations 2; FOV 360–390 mm; scan time 3 min) with an anteriorly located pre-saturation slab. 60 sections, 2.2 mm thick, were obtained in coronal plane 5 min, 10 min and 20 min after diuretic and contrast material injection. In selected cases, additional transverse planes were performed to optimize visualization of anatomic structures. For the examination of children, we reduce the FOV of the sequences and adjust furosemide and gadolinium dosages (0.05 mg kg21 of furosemide and 0.05 mmol kg21 of gadolinium). The source images of static and excretory MRU were then post-processed by the use of a maximum intensity projection (MIP) algorithm. When no dilatation of the urinary tract is visible on the initial T2 weighted images we use excretory MRU. With the use of a diuretic in MRU within the dose range of 4–10 mg of furosemide, the induced distention of the urinary tract was mild and did not result in false-positive diagnosis of substantial dilatation. In patients with mild dilatation, both techniques (static and excretory MRU) are employed. In cases of marked dilatation of the urinary tract and impaired excretory function, static MRU is used. Static MRU is also used for the visualization of urinary tract disorders in women during pregnancy (Figures 1 and 2) [5–7]. The British Journal of Radiology, July 2006

Pictorial review: MR urography

(a)

(b)

Figure 1. A 52-year-old woman with an extrinsic ureteral obstruction caused by a metastasis of an ovarian carcinoma. (a) Maximum intensity projection (MIP) image from an unenhanced T2 weighted MR urograph (MRU) shows a left ureteral obstruction (arrow). Note the changes of chronic hydronephrosis and hydroureter. (b) The axial standard T2 weighted turbo spin echo (TSE) image visualizes a soft tissue mass with heterogeneous signal intensity surrounding the ureter (arrowheads).

Normal variants and congenital anomalies The main indications in children of MRU are congenital anomalies of the kidneys and collecting system [6]. Normal variants and congenital anomalies of the collecting system can be accurately identified with this technique [5]. Knowledge of the myriad appearances of congenital renal and collecting system anomalies and minor anatomic variants is essential for the correct interpretation of urograms. Congenital ureteropelvic junction (UPJ) obstruction is sharply defined UPJ narrowing with dilatation of the pelvocalyceal system, which persists even when patient is placed in a position favouring gravity drainage of the pelvis (Figure 3) [1]. Large extrarenal pelves may simulate hydronephrosis when they are stressed by diuresis. MRU can also accurately detect complete and incomplete ureteral duplication by locating the level of fusion. The British Journal of Radiology, July 2006

In cases of complete duplication, the insertion of the superior collecting system is usually ectopic [8]. Common congenital anomalies of the fusion variety have characteristic MR appearances. True congenital hypoplasia is distinctly rare or very difficult to document. Hypoplastic kidneys usually are caused by trauma, infection or ischaemic or obstructive insult during the growth phase. Renal agenesis with contralateral solitary kidney usually associates with Mu¨llerian duct abnormalities (Figure 4) [1].

Filling defects in the ureter or in the pelvocalyceal system Filling defects are demonstrated on MRU as signalvoid areas outlined by the hyperintense surrounding urine, except when they are impacted or filling the entire lumen of ureter. We sometimes perform complementary 615

R Garcıá-Valtuille, A I Garcıá-Valtuille, F Abascal et al

(a)

(b)

Figure 2. Staghorn calculus and chronic hydronephrosis in a 32-year-old pregnant patient. (a) Coronal T2 weighted turbo spin echo (TSE) image and (b) urogram from static MR urography show diffuse cortical atrophy, pyelocaliectasis and a voluminous pyelocaliceal filling defect (arrows). Note also the gestational sac (arrowheads) and a left corpus luteum cyst (white arrow).

616



obstructing stone. The ‘‘globet sign’’ and the ‘‘sipple sign’’ are also useful in differentiation with other entities. However, the morphological differentiation between a small calculus and a small early intrinsic tumour is difficult in some cases, especially if the clinical symptoms are non-specific [1, 5, 8]. Mimickers of filling defects are flow artefacts (usually with central location within the ureter) [9], vessels that can cause an extrinsic impression on the ureter (Figure 8), and ureteral spasm and peristalsis.

Obstruction of the ureter The differential diagnosis of ureteral obstruction include intraluminal (calculi, blood clots, papillary necrosis with sloughed papilla), intramural (tumour, infection diseases, post-surgery/instrumentation trauma, lesions after radiotherapy, ureterocele, megaureter) and extrinsic abnormalities (retroperitoneal fibrosis, invasion or compression by extrinsic malignancy, lymphadenopathy, inflammatory diseases) [1, 3, 8]. MRU allows the precise depiction of the site of the obstruction and the degree of ureterectasis, and may demonstrate the underlying pathology with the help of conventional T1 and T2 weighted sequences (Figures 1, 6 and 9). Figure 3. A 20-year-old man with ureteropelvic junction narrowing (arrow). Coronal maximum intensity projection (MIP) excretory MR urography. The renal pelvis shows typical dilatation and convex inferior border.

axial images because the small filling defects are better visualized in this plane. Instead of MIP images, the source images must always be reviewed because small defects may be obscured by the surrounding urine on MIP projections [5]. The acute stone colic should not be a primary indication for MRU. However, it is important to be aware of the findings of stones in MRU because most common filling defects are the calculi (Figures 2 and 5); round or oval filling defects that tend to become impacted in areas of normal anatomic narrowing – ureteropelvic and ureterovesical junctions, and the site where the ureter crosses the sacrum and the iliac vessels – and cause a variable degree of dilatation of the urinary tract [1]. Blood clots are single or multiple filling defects of various sizes and shapes that may cause temporary ureteral obstruction (Figure 6). They are usually hyperintense on T1 weighted MR images, do not enhance with gadolinium and become much smaller or disappear within several weeks [8]. Transitional cell carcinoma appears as smooth or irregular shaggy filling defects (Figure 7). The signal intensity of transitional cell carcinoma usually differs sufficiently from that of other causes of ureteral filling defects, on conventional T1 and T2 weighted images, to suggest the diagnosis. There is often localized dilatation of the ureter below the level of the expanding intraluminal tumour, in contrast to ureteral collapse distal to an The British Journal of Radiology, July 2006

Filling defects in the urinary bladder MRU provides a non-invasive mean to detect filling defects in the urinary bladder – calculus, blood clot, air bubble, neoplasm, prostatic enlargement, ureterocele or foreign body [5]. The transitional cell carcinoma of the urinary bladder is a single or multiple polypoid defect that arises from the bladder wall and is fixed in position – unlike a calculus, blood clot or air. Sometimes they may produce only focal bladder wall thickening and rigidity (Figure 10). Prostatic enlargement causes a smooth or irregular extrinsic filling defect of varying size at the base of the bladder (Figure 11). If a chronic process, there is trabeculation of the bladder wall and diverticula formation. The distal ureters often have a fishhook deformity due to elevation of the trigone.

Post-operative changes The role of MRU in patients undergoing urinary diversion (ureteroileal by-passes, ureterosigmoidostomy, skin ureterostomy, orthotopic neobladder reconstruction) or after renal transplantation is emerging. MRU allows visualization of anastomoses, as well as of associated complications such as strictures (Figure 12), ureteral compression by lymphocele or haematoma, urine leaks, fistulae (Figure 13), stones or signs of infection. Signal-void within urinary tract in postoperative patients does not always correspond to stones, but may be due to air bubbles or susceptibility artefacts caused by surgical material [1, 10, 11]. 617


(a)

(b)

(c) Figure 4. Left renal agenesia in a 14-year-old woman with didelphic uterus and a vaginal septum. Maximum intensity projection (MIP) image from (a) excretory MR urography demonstrates a normal right kidney with no evidence of left renal tissue. (b) Axial and (c) coronal T2 weighted turbo spin echo (TSE) images show a bicornuate uterus (arrows) with two cervix (arrowheads).

618



(a)

(b)

Figure 5. A 55-year-old man with left-sided ureteral stone. (a) Coronal maximum intensity projection (MIP) excretory MR urograph shows filling defect (arrow) in left distal ureter that is causing mild pyeloureterectasis. (b) Enhanced axial T1 weighted gradient-echo image shows a round dependent filling defect (arrow) in left ureter.


619


(a)

(b)

(c)

(d)

Figure 6. A 52-year-old woman with temporary ureteral obstruction caused by blood clots. (a) Coronal maximum intensity projection (MIP) excretory MR urography and (b) complementary retrograde pyelography show complete proximal ureteral obstruction (arrow) and mild dilatation of the collecting system. (c) Axial gadolinium-enhanced T1 weighted gradient-echo and (d) T2-weighted turbo spin echo (TSE) images demonstrate hypointense tissue filling completely a mildly dilated ureter (arrowhead). (Continued)

620



(e) Figure 6. (Cont.) (e) After several days, excretory urogram from conventional intravenous pyelography demonstrates patency of previously occluded ureter (arrowheads).


621


(a)

(b)

(c) Figure 7. Transitional cell carcinoma of the midureter in a 68-year-old man. (a) Coronal maximum intensity projection (MIP) and (b) source images from excretory MR urography demonstrate a large mass inside the midureter (arrows) with proximal ureteral and pelvocalyceal dilatation. (c) The axial standard T2 weighted turbo spin echo (TSE) sequence confirms the diagnosis by demonstrating a soft-tissue mass (arrowhead) with heterogeneous signal intensity.

622



(a)

(b)

Figure 8. A 73-year-old woman with mild narrowing of the midureter (arrow) caused by left common iliac artery. (a) Coronal maximum intensity projection (MIP) image from excretory MR urography and (b) composite coronal MIPs of both urogram an MR angiography.

Conclusions Static and excretory MRU are complementary methods for morphological and functional evaluation of the urinary system, which can be alternatively employed according to the degree of urinary tract dilatation and renal function. These techniques have some advantages over ultrasound, conventional urography and CT urography in the diagnosis of urological diseases. The three-dimensional nature of the data permits reformation into any plane, and thus virtually eliminates the potential of projection related errors in the diagnosis of different pathological conditions. There are also the safety advantages of eliminating ionizing radiation and the risk of medical complications due to iodinated contrast agents, and is even suitable for assessing transplanted kidneys because of the low nephrotoxicity of gadolinium. MRU, due to its non-use of ionizing radiation, is the most important tool in the diagnostic work-up of The British Journal of Radiology, July 2006

genitourinary pathologies in infants, small children and in women during pregnancy. The major drawback of MRU is its low sensitivity in detecting calcifications and subtle urothelial lesions, the latter due to the reduced spatial resolution compared with conventional excretory urography. However, MRU can be offered as an alternative to conventional urography and CT urography to avoid repetitive radiation exposure in patients with chronic urolithiasis. In conclusion, static and contrast-enhanced excretory MRU provide high-quality imaging of the urinary tract and are an accurate and safe diagnostic alternative to other urological diagnostic procedures. These techniques, combined with conventional MR images, functional MR sequences or MR angiography, in a single session yields a rapid and complete diagnostic evaluation of the entire urinary tract, and have the potential to provide the same information as can be obtained with multiple separate diagnostic studies. 623


(a)

(b)

(c)

(d)

Figure 9. A 77-year-old man with transitional cell carcinoma of the right ureter. (a) Maximum intensity projection (MIP) image from excretory MR urography (MRU) demonstrates right ureteral obstruction (arrow), hydronephrosis and hydroureter. (b) Original source image from excretory MRU shows a large hypointense filling defect inside distal ureter (arrows). (c) Axial T1 weighted image shows a hypointense soft-tissue mass (arrowhead) in the pelvis. (d) An area of subtle enhancement (arrowhead) is demonstrated on the axial section of a contrast-enhanced T1 weighted sequence.

624



(a)

(b)

Figure 10. A 61-year-old man with transitional cell carcinoma of the bladder. (a) Axial T2 weighted turbo spin echo (TSE) image shows an irregular wall thickening at the left-side of the bladder (arrows). (b) Maximum intensity projection (MIP) image from excretory MR urography confirms large irregular filling defect (arrows) on the floor and left-sided wall of the bladder. The tumour does not produce obstruction at the ureterovesical junction.

Figure 12. A 56-year-old woman with ileal loop urinary Figure 11. An excretory MR urograph in a 78-year-old man with benign prostatic hypertrophy. Large, smooth filling defect at the base of the bladder (arrowheads).


diversion. Maximum intensity projection (MIP) image from excretory MR urography shows the post-operative urinary tract anatomy. Both sides are dilated because of stenosis (arrowheads) close to the ureteroenteric implantation site.

625


Figure 13. Vesicovaginal fistula (arrowheads) formation caused by inadvertent injury to the bladder during surgery in a 48-year-old woman. Sagittal maximum intensity projection (MIP) excretory MR urography.

References 1. Nolte-Ernsting C, Adam G, Bu¨cker A. MR urography: examination techniques and clinical applications. Eur Radiol 2001;11:355–72.

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2. O’Malley ME, Soto JA, Yucel EK, Hussain S. MR urography: evaluation of a three-dimensional fast spin-echo technique in patients with hydronephrosis. AJR Am J Roentgenol 1997;168:387–92. 3. Regan F, Bohlman ME, Khazan R, Rodriguez R, SchultzeHaakh H. MR urography using HASTE imaging in the assessment of ureteric obstruction. AJR Am J Roentgenol 1996;167:1115–20. 4. Rohrschneider WK, Haufe S, Wiesel M, Tonshoff B, Wunsch R, Darge K, et al. Functional and morphologic evaluation of congenital urinary tract dilatation by using combined static-dynamic MR urography: findings in kidneys with a single collecting system. Radiology 2002;224:683–94. 5. Nolte-Ernsting C, Bu¨cker A, Adam G, Neuerburg JM, Jung P, Hunter DW, et al. Gadolinium-enhanced excretory MR urography after low-dose diuretic injection: comparison with conventional excretory urography. Radiology 1998;209:147–57. 6. Nolte-Ernsting C, Staatz G, Tacke J, Gu¨nther RW. MR urography today. Abdom Imaging 2003;28:191–209. 7. El-Diasty T, Mansour O, Farouk A. Diuretic contrastenhanced magnetic resonance urography versus intravenous urography for depiction of nondilated urinary tracts. Abdom Imaging 2003;28:135–45. 8. Blandino A, Gaeta M, Minutoli F, Salamone I, Magno C, Scribano E, et al. MR urography of the ureter. AJR Am J Roentgenol 2002;179:1307–14. 9. Girish G, Chooi WK, Morcos SK. Filling defect artefacts in magnetic resonance urography. Eur Radiol 2004;14:145–50. 10. Schubert RA, Go¨ckeritz S, Mentzel HJ, Rzanny R, Schubert J, Kaiser WA. Imaging in ureteral complications of renal transplantation: value of static fluid MR urography. Eur Radiol 2000;10:1152–7. 11. Zielonko J, Studniarek M, Markuszewski M. MR urography of obstructive uropathy: diagnostic value of the method in selected clinical groups. Eur Radiol 2003;13:802–9.



CASE OF THE MONTH

Focal pancreatic lesion: can a neoplasm be confidently excluded? J HAGUE,

FRCR

and Z AMIN,

FRCR

Department of Radiology, Middlesex Hospital, Mortimer Street, London W1T 3AA, UK

Received 18 April 2005 Revised 11 July 2005 Accepted 25 July 2005 DOI: 10.1259/bjr/19356841 ’ 2006 The British Institute of Radiology

A 56-year-old female patient was referred to our institution for further investigation of diarrhoea and weight loss over the previous 3–4 months. Extensive biochemical investigation had revealed no abnormality. The patient underwent a CT scan of the pancreas

(pre-contrast, pancreatic phase and portal venous phase), which demonstrated a low density focus in the head of the pancreas. Selected images from this study are presented below (Figure 1). What is the diagnosis? What further tests should be done?

(a)

(b)

(c)

Figure 1. (a) Unenhanced axial image through the head and neck of the pancreas demonstrating a region of hypoattenuation in the pancreatic head. (b) 5 mm reconstructed contrast enhanced axial section through the head of pancreas. (c) 5 mm reconstructed contrast enhanced axial section through the body and tail of pancreas.


627

J Hague and Z Amin

(a)

(b)

Figure 2. (a) In phase T1 weighted gradient-echo image through the head and neck of pancreas. (b) Opposed phase T1 weighted gradient-echo image through the head and neck of pancreas demonstrating signal drop in part of the head corresponding to the CT abnormality.

The patient subsequently underwent MRI. The in and opposed phase axial images are presented (Figure 2). This demonstrates uniformly high signal in the entire pancreas on the in phase T1, with signal drop in the anterior head/neck, body and tail on the opposed phase images. This is in keeping with uneven fatty infiltration in part of the head and uncinate process, which has been described as a normal variant.

Discussion Fatty change in the pancreas can be diffuse, or may be uneven, sparing certain regions (focal fatty sparing, FFS). Alternatively, fat infiltration can be confined to one region of the pancreas (focal fatty infiltration, FFI) [1, 2]. Diffuse mild fatty infiltration of the pancreas frequently occurs in elderly and obese persons, and is of little clinical importance [1–3]. Focal areas of FFS in the pancreas are analogous to focal sparing in the liver [4]. FFS can mimic a mass, in a pancreas with diffuse fatty infiltration. Areas of FFS are usually within the head or uncinate [1, 2]. FFS has been associated with a pancreas divisum. FFS can appear as a hypoechoic mass on ultrasound, and appears as an enhancing mass relative to the normal pancreas on contrast enhanced CT [4]. FFI is associated with insulin-dependent diabetes mellitus (IDDM), chronic pancreatitis, hepatic disease, 628

dietary deficiency, viral infection and steroid therapy. Distribution of FFI is variable, dominant in the body and tail. FFI of the head or uncinate or both have been described [3]. FFI can appear as a hypoechoic mass on ultrasound, and a hypoattenuating region on non-enhanced CT and contrast enhanced CT. A pancreatic neoplasm may be diagnosed on this basis. There will, however, be no dilatation of the pancreatic duct, and no contour deformation, but the appearances may be progressive on serial scans. MRI using chemical shift can detect and characterize focal fatty infiltration of the pancreas and exclude a diagnosis of neoplasm [5]. The loss of signal intensity on an opposed phase T1 weighted gradient echo image compared with a corresponding in phase image establishes the lipid content of the focal abnormality and excludes a diagnosis of a pancreatic adenocarcinoma or a neuroendocrine tumour, both of which do not contain fat. Focal lesions in the pancreas detected on CT (or ultrasound) that do not have any associated features of malignancy (for example, no venous attenuation, or duct dilatation) and do not deform the contour, may be caused by uneven pancreatic lipomatosis (FFI or FFS) and can be further evaluated with in and opposed phase MRI [5]. The patient’s symptoms resolved without treatment and she remained well on follow up 1 year later. The British Journal of Radiology, July 2006

Case of the month: Focal pancreatic lesion

References 1. Marchal G, Verbeken E, Van Steenbergen W, et al. Uneven lipomatosis; a pitfall in pancreatic sonography. Gastrointest Radiol 1989;14:233–7. 2. Donald JJ, Shorvon PJ, Lees WR. A hypo echoic area within the head of the pancreas; a normal variant. Clin Radiol 1990;41:337–8.


3. Matsumoto S, Mori H, Miyake H, Takaki H, Maeda T, Yamada Y, et al. Uneven fatty replacement of the pancreas: evaluation with CT. Radiology 1995;194:453–8. 4. Jacobs JE, Coleman BG, Arger PH, Langer JE. Pancreatic sparing of focal fatty infiltration. Radiology 1994;190:437–9. 5. Isserow JA, Siegelman ES, Mammone J. Focal fatty infiltration of the pancreas: MR characterization with chemical shift imaging. AJR Am J Roentgenol 1999;173:1263–5.

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August 2006 Volume 79 Issue 944

August 2006, Volume 79, Issue 944 Diffusion-weighted MRI: a new functional clinical technique for tumour imaging ●

In vitro evaluation of stent patency and in-stent stenoses in 10 metallic stents using MR angiography ●

Occupational exposure in the electrophysiology laboratory: quantifying and minimizing radiation burden ●

Usefulness of diffusion/perfusion-weighted MRI in patients with non-enhancing supratentorial brain gliomas: a valuable tool to predict tumour grading? ●

Imaging well-differentiated hepatocellular carcinoma with dynamic triple-phase helical computed tomography ●

Building an anonymized catalogued radiology museum in PACS: a feasibility study ●

Optimizing localization accuracy in head and neck, and brain radiotherapy ●

Evaluation of the larynx for tumour recurrence by diffusionweighted MRI after radiotherapy: initial experience in four cases ●

Non-invasive measurement of perfusion: a critical review of arterial spin labelling techniques ●

●

Book review

●

A young female with galactorrhoea and sudden onset chest pain


COMMENTARY

Diffusion-weighted MRI: a new functional clinical technique for tumour imaging 1

D-M KOH,

MRCP, FRCR

and 2A R PADHANI,

FRCR, FRCP

1

Department of Radiology, Royal Marsden Hospital, Downs Road, Sutton, Surrey SM2 9PT and Mount Vernon Hospital, The Paul Strickland Scanner Centre, Rickmansworth Road, Northwood, Middlesex HA6 2RN, UK 2

Received 30 March 2006 Revised 19 May 2006 Accepted 30 May 2006 DOI: 10.1259/bjr/29739265 ’ 2006 The British Institute of Radiology

One of the key aims of oncological imaging is to differentiate between malignant and non-malignant tissues at all stages of the patient’s cancer care. Accurate staging and precise delineation of the extent of malignancy influences therapeutic decisions, therapy outcomes and, ultimately, patient prognosis. Conventional imaging using ultrasound, CT or MRI detects cancer by identifying anatomical distortion or altered tissue appearances. Tumour tissue conspicuity may be increased after the administration of intravenous contrast medium, thus enhancing detection and delineation. However, identification of small volume active tumour, either at presentation or at early disease relapse remains challenging because small volume disease may not result in detectable structural or morphological change on conventional imaging. Furthermore, the effects of therapy and complications thereof may obscure or mimic recurrent disease. Functional imaging techniques using CT, MRI and positron emission tomography (PET) are increasingly being applied to the evaluation of tumours. These techniques exploit as their contrast mechanism unique pathophysiological changes that occur within tumours; such as altered blood flow, increased glucose metabolism, hypoxia and cellularity. Such functional techniques are increasingly used for tumour detection, for the monitoring of treatment response and to detect relapsed disease. Clinical experience has shown that functional techniques have their own unique strengths and limitations. A new, emerging functional technique that is now finding a role in cancer imaging is diffusion-weighted MRI (DWI or DW-MRI), which produces information about tissue cellularity and the integrity of cellular membranes. This technique may not be well appreciated by general radiologists. DWI can be performed on most The British Journal of Radiology, August 2006

modern MRI machines with relative ease, in a short period of time and without the need for contrast medium administration. The potential for this technique to evaluate the larynx for tumour recurrence after prior radiotherapy is demonstrated in a short communication from Vandecaveye et al in this issue [1]. At a fundamental level, DWI provides information on the random (Brownian) motion of water molecules in tissues. The Brownian displacements of millions of water molecules over time are normally distributed with a mean final value of zero for all time periods measured, but with a standard deviation that is proportional to the diffusion coefficient and time measured. This was the basis for Einstein’s diffusion equation published in 1905, which subsequently helped to earn him the 1921 Physics Nobel Prize. In tissues, DWI probes the movement of water molecules, which occurs largely in the extracellular space. However, the movement of water molecules in the extracellular space is not entirely free, but is modified by interactions with hydrophobic cellular membranes and macromolecules. Hence, diffusion in biological tissue is often referred to as ‘‘apparent diffusion’’. By comparing differences in the apparent diffusion between tissues, tissue characterization becomes possible. For example, a tumour would exhibit more restricted apparent diffusion compared with a cyst because intact cellular membranes in a tumour would hinder the free movement of water molecules. One of the simplest methods of obtaining DWI images is to apply pairs of opposing and balanced magnetic field gradients (but of differing durations and amplitudes) around a spin-echo refocusing pulse of a T2 weighted sequence. Stationary water molecules are unaffected by the paired gradients, and thus retain their signal. Nonstationary water molecules acquire phase information 633

D-M Koh and A R Padhani

from the first gradient, but are not rephased by the second gradient, leading to an overall loss of the MR signal. The signal reduction on the DWI image is proportional to the amount of diffusion water motion occurring during the pulse sequence. Hence, on DWI, there is usually less signal attenuation (i.e. higher signal intensity) of tumour compared with normal tissue due to the restricted diffusion of water molecules in tumours, which is presumed to be due to an increased cellular density. At some anatomical locations (e.g. the brain), DWI is usually performed in three or more gradient directions because of the unequal limitations to diffusion in some directions imposed by tissue organization (e.g. white matter tracts). This phenomenon is termed anisotropy and may be also observed in some visceral organs. For example, in the normal prostate gland, diffusion is greater along the line of the ducts than across the ducts. However, anisotropy is usually not seen in tumours since cancers typically grow in a disorganized fashion. The degree of diffusion-weighting applied is indicated by the b-value (measured in s mm22), which indicates the magnitude and duration of the applied gradients and time between the paired gradients. By varying the amplitudes, lengths and intervals between the diffusion gradients, the sensitivity to the degree of diffusion motion can be altered and the data processed to provide information about actual diffusion distances. Hence, DWI using a larger b-value (e.g. b5500 s mm22) is more sensitive to the slower motion of water molecules and smaller diffusion distances, whereas the converse is true with a smaller b-value (e.g. b550 s mm22). It is important to remember that the phenomenon of water molecule movement detected by DWI occurs at a length scale (typically micrometres) that is significantly larger than intracellular distances, but significantly smaller than the pixel dimensions of typical MR images (typically millimetres). When performing DWI cancer studies, images are typically acquired using different b-values (typically 0– 1000 s mm22). The images obtained at different b-values allow the calculation of the apparent diffusion coefficient (ADC; unit mm2 s21), which is usually presented as a quantitative parametric map. From the discussions above, it is not surprising that on ADC maps, tumours usually demonstrate low ADC values and appear as low signal intensity area compared with normal tissue. This appearance is the inverse of that observed on ‘‘raw’’ DWI images obtained at high b-values. The quantitative ADC values can aid in lesion characterization, and can also be applied to evaluate the treatment response of tumours. Diffusion-weighted MRI is an established tool for the evaluation of intracranial diseases. The technique has been applied successfully to detect early cerebral infarction and for the characterization of brain tumours. More recently, DWI has been used to demonstrate early response of brain tumours to radiation treatment. However, motion-related artefacts, which degrade image quality, have limited the clinical application of DWI to extracranial sites. These motion-related artefacts can now be substantially reduced by the use of parallel imaging, combined with breath-hold, single-shot, echo-planar MRI techniques. Such techniques are useful for the evaluation of the larynx, where respiratory and swallowing 634

movement, together with susceptibility effects from air within the larynx, can significantly degrade image quality. In this issue of BJR, Vandecaveye et al describe their experience of using DWI for the detection of posttreatment recurrent laryngeal tumours in a small number of patients [1]. All cases were validated by histopathology. Thin section (48 slices at 4 mm thickness) axial DWI of the larynx was performed using six b-values, which ranged from b50 s mm22 to 1000 s mm22. Although parallel imaging was not employed, each DWI study was completed in less than 6 min. The four examples presented elegantly demonstrate the potential role of DWI in distinguishing tumour from posttreatment change. Furthermore, the exquisite radiological-pathological comparison enhances our understanding of the pathological basis for their imaging appearances at DWI. In the study, two cases with recurrent tumour were detected at DWI as focal areas of restricted diffusion returning low ADC values. By comparison, two cases demonstrated focal asymmetry of the larynx due to inflammatory change. These areas returned high ADC values at DWI and were due to laryngeal necrosis and oedema at histopathology. Intriguingly, one case with inflammatory change was also evaluated using 18FDG-PET imaging, which revealed a moderately hypermetabolic focus in the larynx (false-positive). As the PET imaging was performed within 6 months of radiotherapy, this case emphasises the potential pitfall of using 18FDG-PET imaging to distinguish between tumour and inflammation in the early post-treatment period. There are challenges to the use of DWI for tumour evaluation. First, the dichotomy of identifying tumour versus post-treatment change may not always be straightforward. There can be substantial overlap in the ADC values between malignant and non-malignant tissue making it difficult to determine disease status. For example, a predominantly necrotic tumour may potentially be confused with necrosis arising from radiation treatment since both would result in higher ADC values. Thus, as with all functional imaging techniques, DWI image information should be interpreted with information from conventional imaging to improve disease assessment. Second, the averaged ADC values derived from regions of interest drawn around tumours may not sufficiently characterize tumour heterogeneity. More sophisticated methods of analysis are needed to adequately account for regional variations, which is particularly important for the evaluation of treatment response in tumours. Third, the analysis of quantitative ADC on commercial platforms lacks standardization. Some commercial software does not allow drawing of free forms to encompass tumour regions. Most commercial software also does not allow image registration and noise filtration, which can significantly impact on the quality of the quantitative data. Clearly, further collaborative work in this area would be welcomed. Despite some of the above limitations, DWI is emerging as a powerful, new diagnostic tool which will be increasingly applied to the evaluation of tumours, as has been demonstrated in the accompanying paper by Vandecaveye et al [1]. Potential applications include The British Journal of Radiology, August 2006

Commentary: Diffusion-weighted MRI

distinguishing tumour from non-tumour tissue, assessing of treatment response and for the prediction of treatment outcome. As the examination is quick and can be conveniently incorporated into existing protocols, assessment of its role in everyday clinical practice could be expediently achieved.

The British Journal of Radiology, August 2006

Reference 1. Vandecaveye V, De Keyzer F, Poorten VV, Deraedt K, Alaerts H, Landuyt W, et al. Evaluation of the larynx for tumour recurrence by diffusion-weighted MRI after radiotherapy: initial experience in four cases. Br J Radiol 2006;79:681–7.

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In vitro evaluation of stent patency and in-stent stenoses in 10 metallic stents using MR angiography O W HAMER, MD, I BORISCH, N ZORGER, MD

MD,

C PAETZEL,

MD,

W R NITZ,

PhD,

J SEITZ,

MD,

S FEUERBACH,

MD

and

Department of Radiology, University Hospital of Regensburg, Franz-Josef-Strauss-Allee 11, 93042 Regensburg, Germany ABSTRACT. In vitro study to investigate the suitability of contrast enhanced magnetic resonance angiography (CEMRA) for determination of stent patency and grading of instent stenoses in 10 metallic stents. The Acculink carotid, DynaLink, Easy Wallstent, JostentSelfX XF, Luminexx, Omnilink, sinus-SuperFlex, SMART, Symphony and ZA stent were separately placed in a vascular phantom. Dedicated stenoses inside the stents generated a concentric lumen narrowing of 50%. CEMRA was performed for each stent. Signal loss inside the stents and artificial lumen narrowing were assessed objectively using the evaluation software of the MR imager. Moreover, three blinded observers determined visibility of stent patency and in-stent stenoses subjectively on a 3-point scale and graded in-stent stenoses. Loss of signal intensity within the stent lumen ranged between 90% (Wallstent) and 5% (ZA), artificial lumen narrowing between 56% (Symphony) and 22% (ZA). For the Symphony and Wallstent, visibility of patency and in-stent stenoses was impaired and the observers’ grading exaggerated the degree of stenoses (by 23% and 33%, respectively). For the remainder of stents, patency and stenoses were visible and stenoses were graded accurately (less than 10% discrepancy from reference standard). In this in vitro study, eight of 10 stents presented with MRI characteristics which enabled determination of stent patency and accurate grading of clinically relevant in-stent stenoses.

Contrast-enhanced three-dimensional magnetic resonance angiography (CEMRA) is a non-invasive alternative to digital subtraction angiography (DSA) and has meanwhile gained wide acceptance regarding the assessment of arterial occlusive disease in virtually all arterial territories [1–5]. Patient monitoring after stent placement is still a challenge for CEMRA due to image distortion by artefacts emanating from the stent. Computed tomography angiography (CTA) has been advocated as a suitable technique for follow-up after stent implantation. However, CTA is associated with the need for ionizing radiation and potentially harmful iodinated contrast material [6, 7]. Moreover, it has been demonstrated that stent visualization on CTA images can be hampered by markedly artificial lumen narrowing and pseudoenhancement within the stent lumen [8]. Colour duplex ultrasound is another non-invasive technique which is capable of assessing stent patency and eventual in-stent stenoses. Performance of this technique, however, can be very difficult in patients with obesity and gaseous distension of the bowel when iliac stents are the target. However, monitoring of endoprostheses in iliac arteries is of particular interest because iliac stenting is a frequently performed procedure. Several experimental studies have been performed to evaluate the nature and amount of artefacts of a variety of stent designs and materials on MR images [9–16]. Some of these studies evaluated whether stent patency can be determined, reporting promising results for a subgroup of stents [17–24]. However, for a clinically 636

Received 23 August 2005 Revised 9 January 2006 Accepted 30 January 2006 DOI: 10.1259/bjr/57301879 ’ 2006 The British Institute of Radiology

valuable follow-up technique, it is crucial to detect not only stent patency, but also significant in-stent stenoses. The purpose of this in vitro study simulating iliac anatomy was to investigate the suitability of CEMRA for determining stent patency as well as detecting clinically relevant in-stent stenoses in 10 metallic stents.

Materials and methods Pre-study to determine the MR background signal MR background signal in vivo Within a pre-study, the MR signal of perivascular pelvic soft tissue was retrospectively determined for 27 patients who had received a CEMRA of the iliac vasculature for clinical purposes. The CEMRA sequence was identical to the sequence used in the main in vitro study (see below). The coronal source image best showing the aortic bifurcation was determined for each patient. Using the MR evaluation software (NUMARIS 3.5 a1.1b; Siemens AG, Medical Solutions, Erlangen, Germany), one of the authors placed a circular userdefined region of interest (ROI) with a diameter of 3– 4 cm in the soft tissues beneath the aortic bifurcation of each patient. The mean signal intensity as well as standard deviation (SD) was registered for each patient. The mean signal intensity¡SD for all 27 patients was calculated. The British Journal of Radiology, August 2006

Detection of stent patency and in-stent stenoses with MRA

MR background signal in vitro Blankets which had been soaked in water at room temperature were positioned in the centre of the magnet and imaged with the identical CEMRA sequence as applied in the main study (see below). The same author who had performed the in vivo measurements for the 27 patients placed ROIs with a diameter of 3–4 cm on a coronal source image of the blankets and determined the mean signal intensity¡SD.

Main in vitro study to determine the MRI characteristics of 10 metallic stents Stents Ten metallic stents were evaluated (Acculink carotid, DynaLink, JostentSelfX XF, Luminexx vascular, Omnilink, sinus-SuperFlex, SMART, Symphony, Easy Wallstent, and ZA-stent). The dimensions, material composition and manufacturers are listed in Table 1. Nine of these 10 stents were self-expanding and one was balloon-expanding (Omnilink). All stents had a dedicated diameter of 10 mm and thus were suitable for the treatment of iliac stenoses (the Acculink carotid is intended for the treatment of carotid stenoses; the stent’s dimensions, however, render an implantation in the iliac vasculature possible also). The stents’ lengths ranged from 38 mm to 60 mm (mean 45.0 mm¡7.6 mm). Vascular phantom The stents were placed in a vascular phantom (endovascular model # 1803; Sawbones, Malmo¨, Sweden). The phantom consisted of tubes made from flexible urethane polymer and simulated the abdominal Table 1. Technical stent parameters and manufacturers Stent

Diameter Length Material (mm) (mm)

Manufacturer

Acculink carotid DynaLink

10

40

nitinol

10

56

nitinol

JostentSelfX 10 XF

44

nitinol

Luminexx Vascular

10

40

nitinol

Omnilink

10

38

10 sinusSuperFlex

60

stainless steel (316L) nitinol

Guidant, Diegem, Belgium Guidant, Diegem, Belgium Abbott Laboratories, Chicago, IL Bard/Angiomed, Karlsruhe, Germany Guidant, Diegem, Belgium

SMART

10

60

Symphony

10

60

Easy 10 Wallstent ZA 10

42 60

Optimed, Ettlingen, Germany nitinol Cordis, Miami Lakes, FL nitinol Boston Scientific, Natick, MA cobalt alloy Boston Scientific, Natick, MA nitinol Cook, Bjaeverskov, Denmark


aorta, renal arteries, common iliac arteries and external iliac arteries of both sides. The common and external iliac arteries were made out of the same tube and thus were of identical dimensions without noticeable transition. All stents were separately implanted in the right common/ external iliac artery of the phantom, which had an inner diameter of 10 mm and a wall thickness of 2 mm. Because the tubes were transparent, complete stent opening and attachment to the tube wall was assured visually. The phantom was filled with a 2.00 mmol l21 gadopentetate dimeglumine in saline solution at room temperature: 1 l of saline solution was mixed with 4 ml of gadopentetate dimeglumine (0.5 mol l21; Magnevist; Schering GmbH, Berlin, Germany). This dilution was chosen according to settings used by other authors who applied similar MR sequences [11, 13, 25]. The MR signal value of this dilution had been shown to be similar to the signal of contrast filled vessels on CEMRA images in vivo. The phantom was embedded in water soaked blankets to simulate perivascular pelvic soft tissue and to enable proper shimming.

Artificial in-stent stenosis A defined lumen narrowing was created by placing a dedicated stenosis made from rigid urethane (model # 1802-2; Sawbones, Malmo¨, Sweden) in each stent. The artificial stenoses were 20 mm in length and cylindrically shaped, showing a central lumen and thus simulating concentric stenoses. The lumen narrowing for all employed stenoses was set to be 50%. MRA imaging All images were acquired at a commercial 1.5 T MRI unit (Magnetom Sonata; Siemens AG, Medical Solutions, Erlangen, Germany). The vascular phantom was positioned in the centre of the magnet. The amount of susceptibility artefacts caused by the stent depends on the orientation of the stent compared with the main magnetic field B0. The distortion of the magnetic field by a cylindrical object is minimal when the cylinder’s long axis is aligned with the direction of the main magnetic field [9, 12, 14]. The artefacts grow when the object is rotated toward a direction in which its long axis is perpendicular to B0. This dynamic was demonstrated in several experimental studies for numerous stents [13, 15]. With the purpose of comparing the stent specific artefacts rather than the orientational dependence, it was decided to choose one orientation of the stent containing tube to B0, which is representative for the common/ external iliac arteries. The image distortion by all of the stents would be more severe for an angle setting more perpendicular to B0 and less severe for an angulation more parallel to B0. However, the relative severity of image distortion and thus the ranking regarding suitability for CEMRA imaging seen for the 10 evaluated stents would be identical, independent of the specific angle. Thus, in order to identify those stents which are best suitable for CEMRA and which should be chosen if follow-up with CEMRA is intended, one angle setting is sufficient as long as the setting is identical for all stents. Therefore, the tube which contained the stent was orientated at a lateral deviation of 30 ˚ as well as an upward deviation of 45 ˚ to the z-axis and main magnetic 637

O W Hamer, I Borisch, C Paetzel et al

field (B0), respectively, to simulate the course of the common/external iliac artery. A phased-array body coil (circularly polarized) served as a receiver coil. Oblique coronal source images (parallel to the orientation of the stent containing tube) were obtained. The applied CEMRA sequence was a three-dimensional gradientecho sequence with RF spoiling (fast low-angle shot, repetition time (TR) 3.37 ms, echo time (TE) 1.24 ms, flip angle 25 ˚, matrix 246 6 512, read-out bandwidth 390 Hz pixel21, 1.1 mm partition thickness (after Fourier interpolation), voxel size 1.3 mm 6 0.7 mm 6 1.1 mm, frequency encoding parallel to the main magnetic field) identical to the sequence usually used for CEMRA of the pelvis in the authors’ department. For each stent, one angiographic data set was acquired. The data sets were post-processed by application of a standard maximum intensity projection (MIP) algorithm (22.5 ˚ rotational intervals around the craniocaudal axis covering 180 ˚). Finally, the images were sent to a MagicView workstation (Siemens AG).

Subjective data analysis Three radiologists, each with 4 years of experience in MR angiogram interpretation, evaluated the images independently on softcopy displays at a MagicView workstation. The observers were blinded regarding the type of stent and degree of in-stent stenosis. Each observer evaluated each stent once; there were no repeated measurements. They were instructed to rank their overall impression of the quality of each CEMRA study (source and MIP images were evaluated together for this analysis) on a 4-point scale (15good, 25minor limitations, 35major limitations, 45not diagnostic). Furthermore, both stent patency and visibility of in-stent stenoses had to be ranked on a 3-point scale (15visible, 25limited visibility, 35not visible). This was done separately for source and MIP images. If ‘‘limited visibility’’ or ‘‘visible’’, the degree of stenosis had to be determined at its narrowest location on the source images. The reference diameter was defined as the diameter of the tube segment adjacent to the proximal stent end. The severity of stenosis had to be expressed as a percentage of the reference diameter. The mean¡SD of the three observers’ gradings and the difference of this mean to 50% (reference standard) were computed for every stent. Objective data analysis For each stent, the author who had performed the measurements of the pre-study assessed (1) the background signal intensity on CEMRA images, (2) the apparent diameter of the stent lumen outside the artificial stenosis and (3) the loss of signal intensity within the stent outside the artificial stenosis. The author was blinded to the type of stent and acquired one data set per stent. Measurements were made on source images by using the evaluation software of the MR imager (NUMARIS 3.5 a1.1b; Siemens AG, Medical Solutions, Erlangen, Germany). Ad (1): The background signal intensity was determined as follows: ROIs with a diameter of 3–4 cm were placed beneath the ‘‘aortic bifurcation’’ of the vascular phantom on coronal source images. The mean signal 638

intensity¡SD was determined for each of the 10 studies. The mean signal value¡SD of these 10 signal measurements was calculated. Ad (2): The apparent stent lumen and artificial lumen narrowing, respectively, were determined as follows: signal intensity plots were drawn perpendicular to the long axis of the stent. The plots were placed within the visible stent lumen outside the in-stent stenosis. The location of the artificial in-stent stenosis was easily visible for eight stents (Acculink, DynaLink, JostentSelfX XF, Luminexx, Omnilink, sinus-SuperFlex, SMART and ZA). For two stents (Symphony and Wallstent), the margins of the in-stent stenoses could not be determined unequivocally on MR images. However, the examiner was aware of the exact location because he knew the experimental setup. The outer contour of the visible stent lumen was determined using a method similar to that used by other groups [13, 26]: Those two points (left and right margin of the stent lumen) on the signal intensity plot were identified where the signal exceeded twice the signal of the background. The apparent stent lumen diameter was defined as the distance between these two points (Figure 1). The ratio of apparent stent lumen diameter and the respective value of the reference tube, which was the tube segment adjacent to the proximal stent end, was calculated and represented the artificial lumen narrowing (the higher the value, the less artificial lumen narrowing, and vice versa). Ad (3): The signal intensity within the stent was determined as follows: a circular ROI was placed (a) within the visible stent lumen without covering parts of the stent struts or the in-stent stenosis and (b) within the reference tube. The mean signal values within both ROIs were measured. The loss of signal intensity within the stent was determined by calculating the ratio of the signal within the stent and the signal in the reference tube for each stent (the higher the value, the less signal loss, and vice versa).

Figure 1. Schematic of signal intensity plot (dotted line) perpendicular to long axis of stent. The apparent stent lumen was defined to be the distance between the two points on the plot where the signal exceeded twice the background signal.



To compare the suitability of the stents for CEMRA, three groups were established on the basis of the degree of artificial lumen narrowing and loss of in-stent signal intensity. The thresholds were chosen according to the strategy published by other authors [13]. Group 1 stents were stents with a minor reduction of signal intensity within the stent and a minor artificial lumen narrowing (signal intensity and apparent stent lumen 66–100% relative to the reference tube). These stents were considered to be well suited for CEMRA. Group 2 stents were considered to be partially suited for CEMRA: one of the two parameters (signal intensity or apparent stent lumen) was less than 66%. Group 3 stents were considered to be not well suited for CEMRA: both parameters were less than 66%.

Results MR background signal The mean background signal intensity, as retrospectively assessed for the 27 in vivo studies, was 10.4¡3.9. The background signal intensity determined within the pre-study for the water soaked blankets was 13.3¡3.5. Because the difference between these two signal values was negligible, the water soaked blankets were considered to be suitable to simulate perivascular pelvic soft tissue, and thus were applied in the main study. The mean background signal intensity for the 10 experimental stent studies was 14.8¡3.9.

Subjective data analysis Statistics Interobserver variability for categorical data (ranking of visibility stent patency and in-stent stenoses) was determined using kappa statistics. In case kappa statistics was not applicable because the cross tabulated scores were not symmetrical, Spearman’s correlation coefficient was calculated. For continuous data (severity of in-stent stenoses), kappa statistics was not applicable because calculation of kappa requires that all observers use the same rating categories. In the present study, however, the observers’ grading of stenoses was not classified, but resulted in a large number of possible responses. Thus, interobserver variability regarding determination of severity of instent stenoses was evaluated by calculating Pearson’s correlation coefficient. A p-value # 0.05 was considered to be statistically significant.

The source and MIP images of all 10 stents are given side by side in Figure 2. The quality of CEMRA images regarding the overall diagnostic value was ranked to be good (score 1) for all stents and observers. The observers’ performance for visibility of stent patency and in-stent stenoses is summarized in Table 2. The scores for the Acculink, DynaLink, JostentSelfX XF, Luminexx, Omnilink, sinus-SuperFlex, SMART, and ZA were equally excellent on both CEMRA source images and MIP images. All observers agreed that visibility of stent patency and in-stent stenoses was worse for the Symphony and Wallstent. Moreover, visibility of patency and in-stent stenoses in these two stents was ranked lower on MIP images compared with the respective source images for 21/24 (88%) of corresponding score pairs. Kappa was 0.51–1.00 (p50.027 to p ,0.0005) and Figure 2. Source images (upper row) and maximum intensity projection (MIP) images (lower row) obtained from oblique coronal 3D contrast enhanced MR angiography (CEMRA) images of 10 metallic stents. The stents from left to right: group 1: DynaLink, ZA; group 2: Acculink carotid, JostentSelfX XF, Luminexx, Omnilink, SMART, sinusSuperFlex; group 3: Symphony, Wallstent. The signal loss within the stents and the degree of artificial lumen narrowing differed considerably between the stents. Imaging characteristics of group 1 and 2 stents (for classification see Table 1) enabled determination of stent patency and accurate delineation of in-stent stenoses by the observers. The MIP algorithm can cause loss of contrast which might artificially worsen the severity of a stenosis or mimic an occlusion. This phenomenon is best appreciated for the group 3 stents.


639

O W Hamer, I Borisch, C Paetzel et al Table 2. Subjective and objective evaluation of contrast enhanced MR angiography (CEMRA) images

Acculink carotid DynaLink JostentSelfX XF Luminexx Vascular Omnilink sinus-SuperFlex SMART Symphony Easy Wallstent ZA

Visibility of stent patencya

Visibility of in-stent stenosisa

Accuracy of assessment of in-stent stenosisb

Intraluminal signal intensityc

Apparent stent lumen diameterd

Source images

MIP images

Source images

MIP images

[%]

[%]

[%]

1-1-1 1-1-1 1-1-1 1-1-1 1-1-1 1-1-1 1-1-1 1-1-1 2-3-2 1-1-1

1-1-1 1-1-1 1-1-1 1-1-1 1-1-1 1-1-1 1-1-1 2-3-2 3-3-3 1-1-1

1-1-1 1-1-1 1-1-1 1-1-1 1-2-1 1-1-1 1-1-1 1-2-2 2-3-3 1-1-1

1-1-1 1-1-1 1-1-1 1-1-1 1-2-1 1-1-1 1-1-1 2-3-3 3-3-3 1-1-1

(+) 1.1 (–) 1.7 (+) 3.5 (+) 4.3 (–) 2.8 (–) 6.3 (+) 8.2 (+) 22.6 (+) 33.3 (+) 3.0

64.8 75.6 84.9 62.8 47.0 64.1 40.3 34.4 10.1 95.8

70.1 76.2 62.3 76.7 73.4 75.8 74.6 43.9 65.0 77.7

Group classificatione

2 1 2 2 2 2 2 3 3 1

a

Three observers ranked the visibility according to the score: 1: visible, 2: limited visibility, 3: not visible. Each number refers to one observer. b Expressed as the difference between the reference standard (stenosis of 50%) and the mean of severity of stenoses, as assessed by the three observers on CEMRA; (+): degree of stenosis was overestimated by the observers, (2): degree of stenosis was underestimated by the observers. c Expressed as percentage of signal intensity in reference tube. d Expressed as percentage of diameter of reference tube. e Classification on the basis of signal intensity and apparent stent lumen diameter; group 1: both parameters >66%, group 2: one of the two parameters ,66%, group 3: both parameters ,66%. MIP, maximum intensity projection.

Spearman’s correlation coefficient 0.65–1,00 (p50.044 to p ,0.0005), indicating high interobserver agreement. The observers’ performance regarding accuracy of grading of in-stent stenosis is given in detail in Table 2. For eight (Acculink, DynaLink, JostentSelfX XF, Luminexx, Omnilink, sinus-SuperFlex, SMART, and ZA) out of 10 evaluated stents, the observer’s grading differed by less than 10% from the reference standard. For five (Acculink, JostentSelfX XF, Luminexx, SMART, and ZA) of these eight stents, the degree of stenosis was overestimated; for three stents (DynaLink, Omnilink, sinus-SuperFlex) it was underestimated. For the Symphony and Wallstent, the discrepancy between the observers’ grading and the reference standard was 22.6% and 33.3%, respectively. For both stents, in-stent stenoses were overestimated. Pearson’s correlation coefficients ranged between 0.62 and 0.85 (p50.002 to p50.059) indicating high interobserver agreement.

Objective data analysis The magnitude of the in-stent signal intensities and degree of artificial lumen narrowing and, based on these criteria, the group classification are summarized in Table 2. The measurements differed considerably between the 10 stents. Eight of the 10 stents (Acculink, DynaLink, JostentSelfX XF, Luminexx, Omnilink, sinus-SuperFlex, SMART, and ZA) exerted less severe artefacts and were consecutively ranked to be partially or well suited for CEMRA. Two stents (Symphony and Wallstent) caused marked image distortion and thus were found to be not well suited for CEMRA. Interestingly, the measurements for the subgroup of eight stents made from nitinol (Acculink, DynaLink, 640

JostentSelfX XF, Luminexx, sinus-SuperFlex, SMART, Symphony, and ZA) were inhomogeneous with the instent signal intensity ranging between 34.4% (Symphony) and 95.8% (ZA) and the apparent stent lumen diameter between 43.9% (Symphony) and 77.7% (ZA) relative to the reference tube. The magnitude of artefacts exerted by the stent made from stainless steel (Omnilink: signal intensity within the stent 47.0%, apparent stent lumen diameter 73.6%) was found to be less extensive than that seen in some nitinol stents.

Discussion MRA is an alternative to digital subtraction angiography providing several advantages. It is a non-invasive technique without the need for ionizing radiation and potentially harmful iodinated contrast material [6, 7]. In recent years, there have been rapid developments regarding the optimization of MRA sequences. With the introduction of CEMRA, the limitations of time-offlight and phase contrast techniques in terms of long imaging times and artefacts caused by slow and turbulent flow have widely been overcome [2, 4]. CEMRA is meanwhile well accepted for assessing arterial occlusive disease. However, the monitoring of vascular endoprostheses is still a challenge because of image distortion due to artefacts caused by the stents. According to data in the literature, stent patency can be assessed by CEMRA for a subgroup of stents [20, 21, 23, 24, 27, 28]. A clinically valuable follow-up technique, however, should also enable a reliable detection of critical in-stent stenoses which cause a significant compromise to blood flow. In general, this occurs for stenoses generating a diameter reduction of 50% or more The British Journal of Radiology, August 2006


[29, 30]. Consecutively, these stenoses are considered to be relevant for surgical or interventional treatment. In this in vitro study, 10 stents suitable for the implantation in iliac arteries were evaluated with regard to their imaging characteristics on CEMRA images. Emphasis was placed on subjective criteria (visibility of stent patency and artificial in-stent stenoses, as well as grading of lumen narrowing) and on objective parameters (signal loss within the stent and artificial lumen narrowing). The stents were made from nitinol, stainless steel and a cobalt-based alloy. The quality of images regarding the overall diagnostic value was ranked to be good for all studies. Thus, there was no bias due to a potentially differing image quality. There are three main types of artefacts associated with metallic vascular implants causing image distortion: (1) radiofrequency (RF) artefacts originating from RFinduced eddy currents in the stent; (2) susceptibility artefacts caused by differences between the magnetic susceptibility of the stent material and the surrounding tissue, leading to local field inhomogeneities; and (3) flow-related artefacts caused by turbulences and consecutive signal loss due to dephasing [9, 12, 14, 16, 27]. The absence of flow and pulsatility in the presented study is a limitation and may reduce the comparability to an in vivo study. However, CEMRA depends on the T1 shortening effects of gadopentetate dimeglumine and is relatively insensitive to flow-related artefacts because of extremely short echo times and thin slices [31]. Thus, the presented setup can be assumed to obtain valid results regarding interstent visibility. In order to quantify RF induced and susceptibility artefacts, the degree of signal loss inside the stent and artificial lumen narrowing was assessed objectively. According to these measurements, the stents were classified into three groups representing their suitability for MRI. This classification correlated well with the observers’ performance regarding the subjective grading of in-stent stenoses. Group 1 and 2 stents (Acculink, DynaLink, JostentSelfX XF, Luminexx, Omnilink, sinusSuperFlex, SMART, and ZA) were considered to be at least partially suited for CEMRA imaging. For these stents, the observers’ accuracy of grading of in-stent stenoses did meet the needs of clinical routine (difference to the standard of reference ,10%). For group 3 stents (Symphony and Wallstent), the number of artefacts was more pronounced. Consecutively, the grading of in-stent stenoses was markedly worse, resulting in an overestimation of 23–33%. Hence, implantation of these two stents seems not be favourable if follow-up with CEMRA is intended. Published data regarding the evaluation of type and amount of artefacts of various stent materials revealed nitinol to be among the favourable alloys and stainless steel to be less suited regarding MR compatibility [12, 13, 15, 16, 32–34]. In the present study, substantial differences between the imaging characteristics within the subgroup of nitinol stents (Acculink, DynaLink, JostentSelfX XF, Luminexx, sinus-SuperFlex, SMART, Symphony and ZA) were noted. This fact emphasises that the magnitude of artefacts not only depends on material composition, but also on other features like wire thickness, total weight and geometry of mesh or slotted tubes. The observation that the Omnilink stent, made The British Journal of Radiology, August 2006

from stainless steel, presented in this in vitro study with less severe artefacts than some nitinol stents confirmed that an advantageous design may be able to compensate for disadvantages due to an unfavourable stent material. For the Symphony and Wallstent, the visibility of stent patency and in-stent stenoses was ranked worse on MIP images than on source images. This was most likely caused by an artificial loss of contrast on MIP images. Loss of contrast is a result of the characteristics of the MIP algorithm which projects noise or distant high signal into the image [26]. MIP images can mimic a stenosis or even a stent or vessel occlusion when the original signal is low, as seen within the Symphony and Wallstent. Therefore, it should be emphasised that for a thorough evaluation of CEMRA studies, interpretation of source and MIP images is mandatory. In contradiction to the remainder of stents evaluated in this study, the Wallstent presented with a complex morphology of the in-stent signal in terms of bandlike artefacts at both stent endings, a circumscribed region of high signal intensity adjacent to the bandlike artefacts and a severe signal reduction in the central part of the stent. This morphology is the same observed in experimental studies performed by other authors who did not apply in-stent stenoses but evaluated the imaging characteristics of the stent itself [13, 14, 32]. The length of the central region of severe signal reduction observed in the present study exceeded the length of the artificial in-stent stenosis. The signal intensity plots for determination of the apparent in-stent lumen and the region of interest for assessment of in-stent signal intensity were placed within the area of severe signal reduction, but outside the in-stent stenosis. To our knowledge, there are two studies in literature which evaluated in-stent stenoses on CEMRA images with, however, different setups: on the basis of objective measurements, Letourneau-Guillon et al determined the degree of artificial lumen narrowing and the delineation of a 50% in-stent stenosis in a Luminexx stent (diameter 8 mm), revealing comparable results as observed in our study for this stent [25]. An observer based subjective evaluation was not performed in this study. Maintz et al employed a vascular phantom with an inner tube diameter of 8 mm and evaluated, among others, the Symphony, SMART, Wallstent and ZA stent [35]. The stent-containing tubes were oriented along the main magnetic field B0. The signal loss inside the stent was assessed objectively, the degree of artificial lumen narrowing and the visibility of in-stent stenoses were determined subjectively in a consensus decision. Grading of in-stent stenoses was not performed. As far as comparisons are appropriate, the results are consistent with the data evaluated in the present study. However, the authors observed differing imaging characteristics for the Wallstent in terms of no signal loss inside the stent and a good visibility of the in-stent stenosis. This discrepancy is most likely due to the different orientation of the stent relative to the main magnetic field because it is well documented that the artefact size increases with a larger angle to B0 [9, 12, 13, 15, 25, 36]. This hypothesis is confirmed by another study performed by the same group of authors evaluating the Wallstent at varying angles relative to B0 [14]. According to this latter analysis, the Wallstent caused severe artefacts when the 641

O W Hamer, I Borisch, C Paetzel et al

longitudinal axis of the stent was diagonal to B0. Furthermore, two other groups observed similar results as seen in our study in terms of a subtotal to total signal void within the Wallstent even though the stent’s orientation was parallel to the main magnetic field [13, 32]. There are limitations to our study. For the sake of comparability, all evaluated stents had a dedicated diameter of 10 mm. Thus, the observed results apply for this specific dimension only. It is possible that the relative artefact size changes with different stent diameters, which might be expected for smaller diameters in particular. However, considering the already outlined data published by Letourneau-Guillon et al and Maintz et al, the imaging characteristics of the Luminexx, Symphony, SMART, and ZA stent were similar for the 10 mm and 8 mm diameter models, indicating low impact of the stent diameter on artefact size [25, 35]. Although the vascular phantom and the experimental setting were designed to simulate the conditions of the iliac vasculature, clinical studies are warranted to investigate if the observations are confirmed in an in vivo setting. First clinical experiences in 27 patients regarding the detection and grading of in-stent stenoses in the JostentSelfX (Abbott Laboratories, Chicago, IL) by CEMRA confirmed the in vitro results of the present study because this stent, which is very similar to the JostentSelfX XF, revealed to be suitable for MR [37]. However, in single cases susceptibility artefacts at the stent’s ends simulated significant in-stent stenoses. This phenomenon, which limited the reliability of CEMRA significantly, was not observed under in vitro conditions. Hence, in vitro studies seem to be suitable to roughly classify stents with regard to their MR suitability. However, the final decision has to be made in vivo. In conclusion, the amount of stent related artefacts differed considerably for the 10 evaluated stents. Two out of 10 tested stents were not suitable for CEMRA follow-up because stent-related artefacts were severe and in-stent stenoses could not reliably be determined. In eight out of 10 stents, artefacts were less pronounced and the observers’ performance regarding the determination of stent patency and grading of in-stent stenoses on CEMRA images did meet the requirements of clinical routine. Radiologists should be familiar with this subgroup of stents if follow-up with CEMRA is intended. However, confirmation of the evaluated data under in vivo conditions is mandatory.

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4. Prince MR, Yucel EK, Kaufman JA, Harrison DC, Geller SC. Dynamic gadolinium-enhanced three-dimensional abdominal MR arteriography. J Magn Reson Imaging 1993;3:877–81. 5. Sueyoshi E, Sakamoto I, Matsuoka Y, Ogawa Y, Hayashi H, Hashmi R, et al. Aortoiliac and lower extremity arteries: comparison of three-dimensional dynamic contrastenhanced subtraction MR angiography and conventional angiography. Radiology 1999;210:683–8. 6. Cochran ST, Bomyea K, Sayre JW. Trends in adverse events after IV administration of contrast media. AJR Am J Roentgenol 2001;176:1385–8. 7. Laroche D, Namour F, Lefrancois C, Aimone-Gastin I, Romano A, Sainte-Laudy J, et al. Anaphylactoid and anaphylactic reactions to iodinated contrast material. Allergy 1999;54 Suppl. 58:13–6. 8. Strotzer M, Lenhart M, Butz B, Volk M, Manke C, Feuerbach S. Appearance of vascular stents in computed tomographic angiography: in vitro examination of 14 different stent types. Invest Radiol 2001;36:652–8. 9. Bartels LW, Smits HF, Bakker CJ, Viergever MA. MR imaging of vascular stents: effects of susceptibility, flow, and radiofrequency eddy currents. J Vasc Interv Radiol 2001;12:365–71. 10. Bartels LW, Bakker CJ, Viergever MA. Improved lumen visualization in metallic vascular implants by reducing RF artifacts. Magn Reson Med 2002;47:171–80. 11. Graf H, Klemm T, Lauer UA, Duda S, Claussen CD, Schick F. [Systematics of imaging artifacts in MRT caused by metallic vascular implants (stents)]. Rofo Fortschr Geb Rontgenstr Neuen Bildgeb Verfahr 2003;175:1711–9. 12. Klemm T, Duda S, Machann J, Seekamp-Rahn K, Schnieder L, Claussen CD, et al. MR imaging in the presence of vascular stents: A systematic assessment of artifacts for various stent orientations, sequence types, and field strengths. J Magn Reson Imaging 2000;12:606–15. 13. Lenhart M, Volk M, Manke C, Nitz WR, Strotzer M, Feuerbach S, et al. Stent appearance at contrast-enhanced MR angiography: in vitro examination with 14 stents. Radiology 2000;217:173–8. 14. Maintz D, Kugel H, Schellhammer F, Landwehr P. In vitro evaluation of intravascular stent artifacts in three-dimensional MR angiography. Invest Radiol 2001;36:218–24. 15. Meyer JM, Buecker A, Schuermann K, Ruebben A, Guenther RW. MR evaluation of stent patency: in vitro test of 22 metallic stents and the possibility of determining their patency by MR angiography. Invest Radiol 2000;35:739–46. 16. Wang Y, Truong TN, Yen C, Bilecen D, Watts R, Trost DW, et al. Quantitative evaluation of susceptibility and shielding effects of nitinol, platinum, cobalt-alloy, and stainless steel stents. Magn Reson Med 2003;49:972–6. 17. Hagspiel KD, Leung DA, Nandalur KR, Angle JF, Dulai HS, Spinosa DJ, et al. Contrast-enhanced MR angiography at 1.5 T after implantation of platinum stents: in vitro and in vivo comparison with conventional stent designs. AJR Am J Roentgenol 2005;184:288–94. 18. Amano Y, Gemma K, Kawamata H, Kumazaki T. Fatsuppressed gadolinium-enhanced three-dimensional magnetic resonance angiography adequately depicts the status of iliac arteries following atherectomy and stent placement. Cardiovasc Intervent Radiol 1998;21:345–7. 19. Cavagna E, Berletti R, Schiavon F. In vivo evaluation of intravascular stents at three-dimensional MR angiography. Eur Radiol 2001;11:2531–5. 20. Juergens KU, Tombach B, Reimer P, Vestring T, Heindel W. Three-dimensional contrast-enhanced MR angiography of endovascular covered stents in patients with peripheral arterial occlusive disease. AJR Am J Roentgenol 2001;176: 1299–303.


Detection of stent patency and in-stent stenoses with MRA 21. Link J, Steffens JC, Brossmann J, Graessner J, Hackethal S, Heller M. Iliofemoral arterial occlusive disease: contrastenhanced MR angiography for preinterventional evaluation and follow-up after stent placement. Radiology 1999;212:371–7. 22. Matsumoto AH, Teitelbaum GP, Barth KH, Carvlin MJ, Savin MA, Strecker EP. Tantalum vascular stents: in vivo evaluation with MR imaging. Radiology 1989;170:753–5. 23. Schurmann K, Vorwerk D, Bucker A, Neuerburg J, Grosskortenhaus S, Haage P, et al. Magnetic resonance angiography of nonferromagnetic iliac artery stents and stent-grafts: a comparative study in sheep. Cardiovasc Intervent Radiol 1999;22:394–402. 24. Tello R, Thomson KR, Witte D, Becker GJ, Tress BM. Dynamic gadolinium DTPA-enhanced magnetic resonance of intravascular stents. Invest Radiol 1998;33:411–4. 25. Letourneau-Guillon L, Soulez G, Beaudoin G, Oliva VL, Giroux MF, Qin Z, et al. CT and MR imaging of nitinol stents with radiopaque distal markers. J Vasc Interv Radiol 2004;15:615–24. 26. Anderson CM, Saloner D, Tsuruda JS, Shapeero LG, Lee RE. Artifacts in maximum-intensity-projection display of MR angiograms. AJR Am J Roentgenol 1990;154:623–9. 27. Amano Y, Ishihara M, Hayashi H, Gemma K, Kawamata H, Amano M, et al. Metallic artifacts of coronary and iliac arteries stents in MR angiography and contrast-enhanced CT. Clin Imaging 1999;23:85–9. 28. Matsumoto AH, Teitelbaum GP, Carvlin MJ, Barth KH, Savin MA, Strecker EP. Gadolinium enhanced MR imaging of vascular stents. J Comput Assist Tomogr 1990;14:357–61.


29. May AG, van de Berg L, Deweese JA, Rob CG. Critical arterial stenosis. Surgery 1963;54:250–9. 30. Rose SC. Noninvasive vascular laboratory for evaluation of peripheral arterial occlusive disease: Part I--hemodynamic principles and tools of the trade. J Vasc Interv Radiol 2000;11:1107–14. 31. Prince MR. Gadolinium-enhanced MR aortography. Radiology 1994;191:155–64. 32. Hilfiker PR, Quick HH, Debatin JF. Plain and covered stentgrafts: in vitro evaluation of characteristics at threedimensional MR angiography. Radiology 1999;211:693–7. 33. Barras CD, Myers KA. Nitinol - its use in vascular surgery and other applications. Eur J Vasc Endovasc Surg 2000;19:564–9. 34. Duerig T, Tolomeo D, Wholey M. An overview of superelastic stent design. Min Invas Ther Allied Technol 2000;9:235–46. 35. Maintz D, Tombach B, Juergens KU, Weigel S, Heindel W, Fischbach R. Revealing in-stent stenoses of the iliac arteries: comparison of multidetector CT with MR angiography and digital radiographic angiography in a Phantom model. AJR Am J Roentgenol 2002;179:1319–22. 36. Trost DW, Zhang HL, Prince MR, Winchester PA, Wang Y, Watts R, et al. Three-dimensional MR angiography in imaging platinum alloy stents. J Magn Reson Imaging 2004;20:975–80. 37. Hamer OW, Finkenzeller T, Borisch I, Paetzel C, Zorger N, et al. In vivo evaluation of patency and in-stent stenoses after implantation of nitinol stents in iliac arteries using MR angiography. AJR Am J Roentgenol 2005;185:1282–8.

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Occupational exposure in the electrophysiology laboratory: quantifying and minimizing radiation burden 1,2

N THEOCHAROPOULOS, MSc, 1J DAMILAKIS, PhD, 1K PERISINAKIS, PhD, 3E MANIOS, MD, 3P VARDAS, 4 MD, PhD, FESC, FACC and N GOURTSOYIANNIS, MD, PhD 1

Departments of Medical Physics, 3Cardiology and 4Radiology, Faculty of Medicine, University of Crete, P.O. Box 2208, Iraklion 71003, Crete and 2Department of Natural Sciences, Technological Education Institute of Crete, P.O. Box 140, Iraklion 71004, Crete, Greece ABSTRACT. Fluoroscopically guided procedures in the electrophysiology room, such as radiofrequency catheter ablation and implantation of cardiac resynchronization devices, may result in high radiation exposure of electrophysiologists and assisting staff. Our aim was to provide accurate and applicable data on occupational doses to the electrophysiology laboratory personnel. We exposed fluoroscopically an anthropomorphic phantom at three projections common in electrophysiology studies. For each exposure, scattered radiation was measured at 182 sites of the cardiology room at four body levels. Effective dose values, eye lens, skin and gonadal doses to the laboratory staff were calculated. Our study has shown that a procedure requiring 40 min of fluoroscopy yields a maximum effective dose of 129 mSv and a maximum value of gonadal dose of 56.8 mSv to staff using a 0.35 mm lead-equivalent apron. A conservative estimate of the electrophysiologist’s annual maximum permissible workload is 155 procedures. Staff effective dose values vary by a factor of 40 due to positioning during fluoroscopy and by a factor of 11 due to radiation protection equipment. Undercouch protective shields may reduce gonadal doses up to 98% and effective dose up to 25%. Consequently, radiation levels in the electrophysiology room are not negligible. Mitigation of occupational exposure is feasible through good fluoroscopy and working practices.

In recent years, due to technological advances, a plethora of new fluoroscopically guided procedures has been introduced in the electrophysiology suite [1–9]. Radiofrequency catheter ablation and implantation of rhythm devices and cardioverter defibrillators have been proven to supersede the effectiveness of other therapeutic approaches and, as a result, their application is increasing [7, 8]. Hence, electrophysiologists and support personnel may be exposed to considerable levels of radiation, depending on the laboratory workload and complexity of the procedures [9–11]. The need for quantifying the risk of radiation detriment to laboratory staff is imperative. Because of the different radiosensitivities of the various body organs and tissues, for non-homogeneous irradiations, radiation detriment is assessed by a radiation quantity termed effective dose, introduced by the International Commission on Radiological Protection (ICRP). The effective dose E is the weighted average of the mean absorbed dose DT to 22 organs/tissues, where the organ/tissue-specific weighting factor wT is the fractional organ/tissue contribution to the total body detriment [12]: P E~ wT DT T

Address correspondence to: J Damilakis, Assistant Professor, Department of Medical Physics, Faculty of Medicine, University of Crete, P.O. Box 2208, Iraklion 71003, Crete, Greece.

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Received 7 June 2005 Revised 9 December 2005 Accepted 3 January 2006 DOI: 10.1259/bjr/76128583 ’ 2006 The British Institute of Radiology

The risk of genetic effects, i.e. impairment on progeny, depends on the preconceptual radiation dose delivered to staff gonads. Other organ/tissue doses of interest are the skin and the eye lens dose, which together with effective dose, are controlled by regulatory annual limits [12]. To the best of our knowledge, there is no reported experience on occupational effective and gonadal doses as well as maximum permissible workloads regarding the electrophysiology suite. The present study aims to provide accurate and applicable data on occupational doses to the electrophysiology laboratory personnel.

Methods and materials Simulated exposures and scatter measurements An anthropomorphic tissue-equivalent phantom (RANDO; Alderson Research Labs, Stanford, CA) simulating an adult was positioned supine on the surgical table and fluoroscopically exposed at three distinct projections common in interventional cardiology: (a) posteroanterior (PA), (b) left anterior oblique 45 ˚ (LAO) and (c) right anterior oblique 25 ˚ (RAO). A mobile undercouch C-arm fluoroscopic unit was used (Philips BV300-R2; Philips Medical Systems, Best, The Netherlands) at a focus to image intensifier distance of The British Journal of Radiology, August 2006

Occupational exposure in the electrophysiology laboratory

100 cm. The half value layer of the X-ray tube was 4.7 mm aluminium at 70 kVp and the diameter of the input field size was 23 cm. Scatter air kerma rates were measured using a handheld ionization chamber (LB 1236, EG&G BERTTHOLD ) connected to a portable digital meter (UMo LB 123, EG&G BERTTHOLD). The operating theatre floor was divided into two grids relative to the long sides of the table. The grids were 3 m61.5 m in dimensions and consisted of 25 cm625 cm square cells. 91 measuring points were thus defined on each grid. Measurements were performed at the height of the gonads (80 cm above the floor), the waist (100 cm), the neck (150 cm) and the face (165 cm). The air kerma readings were divided by the dose–area product (DAP) rate of each exposure to obtain data independent of exposure parameters. All measurements were repeated with two removable flexible undercouch 0.5 mm lead-equivalent protective shields (60 cm670 cm) attached to the table’s long sides (Mavig, Muenchen, Germany).

Calculation of occupational doses Using coefficients provided by ICRP [13], the face level air kerma measurements were converted into eye lens dose and to personal dose equivalent Hp (0.07)F [14]. The latter corresponds to the dose received by the facial skin. The neck level air kerma measurements were converted into personal dose equivalent Hp (0.07)N while the waist level measurements were converted to personal dose equivalent Hp (10)W. The genital levels measurements were converted to male and female staff gonadal dose with use of coefficients provided by ICRP [13]. Previously published attenuation data were used for the calculation of gonadal and deep doses under a 0.35 mm and 0.50 mm lead-equivalent apron, and of the eye lens doses under 0.35 mm lead-equivalent goggles [15]. The polychromatic nature of X-rays was accounted for by weighting the available monoenergetic dose conversion coefficients on the basis of the characteristics of the utilized spectra. The values of Hp W under apron protection and the Hs N values at neck level were used for the calculation of staff effective dose according to the following equations [16]: E~0:06|(Hp (0:07) N Hp (10) W )zHp (10) W

ð1Þ

when a thyroid shield is not used, and E~0:02|(Hp (0:07) N Hp (10) W )zHp (10) W

ð2Þ

when a thyroid shield is used. Equation (1) was also used for the calculation of effective dose to unshielded personnel by means of attenuation-free Hp (10)W values. The neck level air kerma measurements were also converted into to personal dose equivalent Hp (10)N, to enable the correlation of the reading of an over-apron chest dosemeter with the actual value of effective dose to monitored staff. To investigate the effect of beam quality on staff doses, the Hp (0.07)N, Hp (10)W and Hp (10)N values were calculated for four tube voltage settings: 70 kVp, 85 kVp, 100 kVp and 120 kVp. The British Journal of Radiology, August 2006

Calculation of radiation burden from a complex procedure Derived occupational doses are projection specific and DAP normalized since they originate from the initial DAP normalized scatter radiation measurements. This normalization eliminates the dependence on exposure technique factors (kilovoltage and milliamperage) or instrumentation and enables occupational dose calculations from any complex procedure on the basis of the individual DAP value of each projection involved [17]. The fluoroscopy time required per patient strongly depends on the type of electrophysiological study. Reported fluoroscopy course durations range from 3.54 min per procedure for conventional cardiac rhythm device implantation [18] to 77 min per patient for biventricular pacing [10]. For the estimation of the staff radiation burden per patient, a complex procedure requiring 40 min of fluoroscopy was considered. The total DAP of the examination was assumed to be 4300 cGy cm2 and the relative contribution of the PA, RAO and LAO projection was 58%, 15% and 27%, respectively [3]. Adopted fluoroscopy time and total DAP value are similar to those derived from patient studies performed at our institution regarding biventricular pacing (35.2 min and 4765 cGy cm2) [19] and radiofrequency catheter ablation (41 min) [3]. For the hypothesised complex procedure, staff doses and workloads were calculated at three positions at the left hand side of the patient and at the symmetrical positions at the right hand side of the patient. Two sites of interest adjacent to (12 cm from) the surgical table were considered, at the level of the patient’s groin (femoral area) and heart (subclavian position), while the third position was selected 1 m away from the operating table, also at the level of the patient’s heart. The effect of projection ratios constituting a complex procedure on staff effective dose values was investigated by considering an alternative procedure in which the roles between the LAO and RAO projections were interchanged, i.e. the RAO contribution was increased by a factor of 1.8 and the LAO contribution was decreased by a factor of 0.56.

Calculation of workloads Maximum annual permissible workloads were calculated on the basis of the ICRP recommendations regarding effective dose (20 mSv per year), eye lens dose (150 mSv per year) and skin dose (500 mSv per year) [12].

Results Distribution of effective dose values per projection The highest occupational exposure during a PA or RAO projection is delivered next to the patient’s chest at the left hand side (subclavian position), while during a LAO projection the symmetrical site at the right hand side of the patient receives the highest radiation burden (Figure 1). The maximum values of effective dose to the electrophysiology laboratory staff, 645

N Theocharopoulos, J Damilakis, K Perisinakis et al

protected by 0.5 mm lead-equivalent apron and thyroid collar, are 14 nSv cGy21, 14.8 nSv cGy21 and 21 10.5 nSv cGy for the PA, RAO and LAO projection, respectively.

Radiation burden from complex procedure During an electrophysiological study, the highest radiation levels occur adjacent to the patient’s heart. A complex 40 min procedure yields between 54.2 mSv and 129 mSv depending on radiation protection measures (Table 1). Effective dose values are considerably lower at the right hand side of the patient where, at 1 m from the table, they range from 3.38 mSv to 8.47 mSv, depending on protective equipment used. The annual permissible workload derived for the effective dose constraint is at minimum for a staff member using only a 0.35 mm leadequivalent apron while occupying the site of highest exposure, and equals 155 procedures. Since cancer risk increases with effective dose and workloads decrease with dose, their variation with staff positioning and protection devices can be readily appreciated. Interchanging roles between the LAO and RAO projections, as in the alternative complex procedure considered, would increase effective doses to staff working adjacent to the tableside at the left hand side of the patient by 10.1%, and reduce effective dose values at the right hand side of the patient by 10.6%. By analogy to effective dose, eye lens and face skin dose present their peak values, 389 mSv and 446 mSv per procedure, respectively, at the left hand side of the patient (Table 2). Maximum permissible workloads derived from the skin dose limit always exceed that derived from the eye lens dose constraint, unless eye protection is used. Gonadal doses are also higher at the left hand side of the patient. Moreover, male gonads receive approximately twice the dose delivered to female gonads. The

maximum estimated gonadal dose was 56.8 mSv per procedure for the male staff using a 0.35 mm leadequivalent apron (Table 3).

Correlation between effective dose and dosemeter reading An over-apron dosemeter worn at chest height records an Hp value that is from 3.6 to 27 times higher than the effective dose actually received by the monitored staff member in the electrophysiology laboratory, depending on radiation protection equipment used and the tube voltage (Table 4).

Effective dose and tube voltage The penetrating ability of radiation through protective apron and tissue increases with beam energy, and hence effective dose values depend also on tube voltage. The correction factors of Table 5 show the effect of voltage on staff effective dose and can be used as multipliers to adjust the data of Table 1 or Figure 1, which were derived for 100 kVp.

The effect of removable protective shields For the complex 3-projection procedure considered, the effect of the removable under-apron protective drapes on gonadal doses is markedly different from that on effective dose values, within the confines of the electrophysiology laboratory (Figure 2). Effective dose is reduced up to 20% for staff members working at the left hand side of the patient while for those working at the contralateral side, reduction up to 40% occurs in the vicinity of the patient’s groin and up to 50% at 0.5 m from the couchside. The reduction in gonadal doses is

Figure 1. Distribution of effective dose in the electrophysiology laboratory to 0.5 mm lead-equivalent apron and collar protected staff from (a) posteroanterior (PA), (b) right anterior oblique 25 ˚ (RAO) and (c) left anterior oblique 45 ˚ (LAO) heart fluoroscopy. Provided values are dose–area product (DAP)-normalized and expressed in nSv cGy cm22.

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Occupational exposure in the electrophysiology laboratory Table 1. Effective dose and maximum permissible annual workloads for electrophysiology laboratory staff, calculated for four radiation protection conditions: (a) 0.35 mm lead-equivalent apron, (b) 0.35 mm lead-equivalent apron and 0.5 mm leadequivalent collar, (c) 0.5 mm lead-equivalent apron and (d) 0.5 mm lead-equivalent apron and 0.5 mm lead-equivalent collar Side of the patient working from

Left hand side

Position (coordinates)*

Subclavian position (0,0)

Femoral area (20.5,0)

1 m from table (0,1)

Right hand side




Radiation protection measures

a b c d a b c d a b c d a b c d a b c d a b c d

Effective dose mSv per procedure{

129 106 79.1 54.2 19.7 13.7 14.1 7.82 21.2 14.8 15.1 8.42 67.8 54.2 42.9 28.2 18.4 13.7 12.4 7.47 8.47 5.99 5.98 3.38

Maximum permissible workloads{ Procedures

Gy m2

155 189 253 369 1015 1460 1418 2558 943 1353 1325 2375 295 369 466 709 1087 1460 1616 2677 2361 3339 3344 5917

0.7 0.8 1.1 1.6 4.4 4.8 6.2 11.0 4.1 5.8 5.7 10.3 1.3 1.6 2.0 3.1 4.7 6.3 7.0 11.6 10.2 14.5 14.5 25.6

*As in Figures 1 and 2. {4300 cGy cm2 at 100 kVp, 40 min of fluoroscopy. {Based on the 20 mSv annual limit on effective dose.

Table 2. Eye lens dose, face skin dose and corresponding maximum permissible annual workloads for the electrophysiology laboratory staff Side of the patient staff are working from

Position (coordinates)

Left hand side


Eye lens Eye protection

Dose mSv per procedure*

Unprotected

389

Goggles Femoral area (20.5,0)

Unprotected Goggles


Unprotected Goggles

Right hand side


Unprotected Goggles


Unprotected Goggles


Face skin

Unprotected Goggles

19.4 153 7.7 125 6.3 194 9.7 107

Maximum permissible workloads{ procedures

386

Dose mSv per procedure

Maximum permissible workloads{ procedures

446

1121

176

2844

144

3481

223

2246

123

4072

7721 979 19587 1199 23972 773 15466 1402

5.3

28046

31.2

4809

1.6

96175

35.8

13964

*4300 cGy cm2 at 100 kVp, 40 min of fluoroscopy. {Based on the 150 mSv annual limit on eye lens dose. {Based on the 500 mSv annual limit on skin dose.


647

N Theocharopoulos, J Damilakis, K Perisinakis et al Table 3. Female and male gonadal doses for the electrophysiology laboratory staff protected by (a) 0.35 mm lead-equivalent apron and (b) 0.5 mm lead-equivalent apron Side of the patient working from

Position (coordinates)

Left hand side

Subclavian position (0,0) Femoral area (20.5,0) 1 m from table (0,1)

Right hand side

Subclavian position (0,0) Femoral area (20.5,0) 1 m from table (0,1)


Female staff

Male staff

Gonadal dose mSv per procedure*

Gonadal dose mSv per procedure

a b a b a b a b a b a b

56.8 25.0 6.4 2.8 3.6 1.6 34.6 15.2 8.6 3.8 2.5 1.1

119 52.2 13.3 5.8 7.4 3.3 72.3 31.8 17.9 7.9 5.1 2.3

*4300 cGy cm2 at 100 kVp, 40 min of fluoroscopy.

also asymmetrical. However, protection is more effective since staff gonadal doses are reduced by 90–98% at both sides of the patient’s groin. It is evident that the undercouch shields protect mostly the lower part of the body.

Discussion Applicability of provided dosimetric data A recent study has shown that scattered radiation patterns from fluoroscopically exposed patients are similar to those obtained from simulated exposures of a RANDO phantom and that variations in the body habitus have a minor impact on scatter air kerma rates [20]. Hence, presented data allow the accurate estimation of staff occupational doses and radiogenic risks in the electrophysiology laboratory. Although not explicit, doses of Tables 1–3 are in effect DAP-normalized, since a strictly defined complex procedure of 4300 cGy cm2 was considered. Hence, staff doses in any institution from procedures similar in geometry to the procedure adopted in the current study may be estimated on the basis of the tabulated data, corrected for total DAP value. Even major alterations in the ratios of the constituent components of fluoroscopic imaging will have minor

impact on doses to tableside workers. However, if more accurate calculations are required, the isodose maps of Figure 1 and the individual DAP values of the constituent projections may be used. Effective dose estimation using the above methods is demonstrated in the following example. A cardiologist has performed an electrophysiology study positioned next to the patient’s right hip (femoral area, x520.5, y50). The total DAP value recorded during fluoroscopy at 85 kVp was 5000 cGy cm2 and the estimated relative contribution of PA, LAO and RAO projections is 60%, 25% and 15%, respectively. From the data of Table 1 and the high voltage correction factors of Table 5, the effective dose (E) received by the cardiologist may be obtained as follows (1st method): E~

7:47 mSv |5000 cGy cm2 |0:58~5:04 mSv 4300 cGy cm2

More precisely (2nd method), the isodose curves of Figure 1 can be combined with Table 5 correction factors to yield E: E~(

1:5 nSv |5000 cGy cm2 |0:60|0:58) cGy cm2 1:5 nSv |5000 cGy cm2 |0:25|0:58) z( cGy cm2 3 nSv |5000 cGy cm2 |0:15|0:58)~5:00 mSv z( cGy cm2

Table 4. Ratios of dosemeter reading* to effective dose, for an over-apron chest dosemeter Radiation protection measures

70

85

100

No protection 0.39 0.47 0.53 0.35 mm lead-equivalent apron 9.4 7.6 5.6 0.35 mm lead-equivalent apron 15 10 6.8 and 0.5 mm lead-equivalent collar 0.5 mm lead-equivalent apron 13 11 9.1 0.5 mm lead-equivalent apron 27 19 13 and 0.5 mm lead-equivalent collar *Hp (10)N.

648

Table 5. Correction factors (multipliers) for effective dose

Tube voltage (kVp)

calculation relative to data of Table 1 and Figure 1 120

0.56 3.6 4.7

6.2 7.7


0.35 mm lead-equivalent apron 0.35 mm lead-equivalent apron and 0.5 mm lead-equivalent collar 0.5 mm lead-quivalent apron 0.5 mm lead-equivalent apron and 0.5 mm lead-equivalent collar

Tube voltage (kVp) 70

85

120

0.52 0.36

0.72 0.62

1.57 1.84

0.44 0.34

0.65 0.58

1.66 1.55


Occupational exposure in the electrophysiology laboratory

Figure 2. Percent reduction of (a) effective dose and (b) gonadal dose delivered to electrophysiology laboratory staff from the complex three-projection procedure considered, accomplished with use of removable undercouch protective drapes.

A third, retrospective, method proposed for effective dose estimation is based on personal dosemeter reading and the conversion coefficients of Table 4. Suppose that during an electrophysiological study performed at an average voltage of 70 kVp, the over-apron dosemeter positioned over a 0.35 mm lead-equivalent apron records an Hp N value of 4.7 mSv. The actual effective dose delivered will be: E~

dosemeter reading Hp N 4:7 mSv ~ ~0:5 mSv 9:4 9:4

protective devices, is of little value in developing a radiation protection strategy. Mapping the electrophysiology laboratory in terms of effective dose and gonadal dose for several beam geometries and different protective equipment is essential for the precise calculation of staff radiogenic risk. Moreover, the use of any available removable radiation barriers should be accurately duplicated in a separate set of measurements, since there is no other reliable method of predicting their efficiency in dose reduction [21, 22].

The actual radiation risks

Maximum permissible workloads

Occupational exposure in the electrophysiology laboratory is most unlikely to incur deterministic effects to staff. Not only are dose limits well below the threshold of the induction of such effects, but also the fluoroscopy times required for the accumulation of threshold doses are extremely high. The 2 Sv-threshold for the induction of erythema [12] will be reached after at least 3000 h of fluoroscopy, while cataract formation occurs after at least 1700 h of fluoroscopy or 1 Sv. The threshold dose for temporary sterility under prolonged exposure is 0.4 Gy per year to male gonads [12], which corresponds to 2240 h of fluoroscopy. Hence, the actual concern for the laboratory staff is the induction of stochastic effects, such as cancer and impairment to progeny.

For each working site of interest, the overall maximum permissible workload is the minimum of the three partial workloads imposed by the constraints on eye lens, face skin and effective dose. Since workloads depend on positioning and protection measures, a simplified and safe approach is to adopt the workload derived for the most burdened position with the minimum protection of a 35 mm lead-equivalent apron. Hence, no dose limit will be violated if the maximum DAP value of fluoroscopy performed annually is less than 0.7 Gy m2. This corresponds to 6200 min of fluoroscopy per year or 155 procedures.

Factors affecting occupational doses The quantity directly related to the risk of cancer induction and used for the expression of occupational dose limits is effective dose. Genetic risk, on the other hand, calls for the estimation of gonadal doses to staff. Our study has shown that, within a 1.5 m radius from the tableside, effective dose varies by a factor of 40 due to staff positioning and by a factor of 11 due to differences in radiation protection measures. Gonadal doses present even higher discrepancies. Hence, providing limited data regarding scatter dose levels, or even effective doses, without exact information on beam geometry, location or The British Journal of Radiology, August 2006

Comparison with previous studies Calkins et al [24] measured radiation exposure to medical personnel during radiofrequency catheter ablation using thermoluminescent dosimetry. Using the mean exposure, they recorded per case at waist and thyroid level (532 mSv and 156 mSv, respectively) and by assuming a 0.5 lead-equivalent apron and collar, the current study methodology yields a mean effective dose value of 15.2 mSv for 44 min of fluoroscopy. Lindsay et al [25] estimated an effective dose equivalent, which is the predecessor of effective dose, of 18 mSv per case or 55 min of fluoroscopy when a thyroid shield is used, and of 28 mSv per case without thyroid protection, for the physician located at the femoral area. The effective dose equivalent was approximately twice as high near the 649

N Theocharopoulos, J Damilakis, K Perisinakis et al

influenced by the volume of tissue lateral to the scattering site, real patients, who may differ considerably in size compared with the RANDO phantom, may produce a somewhat different scatter pattern.

subclavian position. Regarding eye lens dose, Calcins et al [24] reported 281 mSv per case, Vano et al [26] 294 mSv and Kuon et al [21] more than 165 mSv per year for an annual workload of 1000 interventions. Given the multiplicity and complexity of factors affecting occupational doses, reported data are in broad agreement with the current study estimates. It should be stressed, however, that the vast majority of previously reported occupational doses are based on fluoroscopy duration observations and not on DAP. Hence, they are strongly dependent on fluoroscopy technique factors and equipment used, and cannot serve the purpose of accurate dose estimation in other electrophysiology laboratories.

Conclusions

Mitigation of radiation risks

References

Two easy-to-apply practices are recommended for minimizing personnel exposure in the electrophysiology laboratory. Approaching the patient from the right hand side, specifically at the groin level, is preferable since the radiation backscatter effect is less pronounced. Also, reducing tube voltage and milliamperage is beneficial for staff doses in two ways: the total DAP per patient is diminished and staff effective dose values per unit DAP fall as the beam energy decreases. Personal protection is also extremely important. A 0.35 mm lead-equivalent apron reduces effective dose to staff by a factor of 10, compared with the unshielded values. Additional use of a 0.5 mm thyroid collar provides a further 1.5-fold decrease. Use of both 0.5 mm lead-equivalent apron and collar provides a protection factor of 26. Furthermore, operators should avoid or limit the use of LAO projection when positioned on the right side of the patient and of RAO projection when positioned on the left [22]. Other good fluoroscopy practices, such as strict beam collimation use of pulsed fluoroscopy and last frame hold feature, act to reduce radiation administered per patient and keep staff radiation doses as low as reasonably achievable [21, 27, 28]. However, the degree of radiation reduction depends on specific characteristics of the equipment used. Therefore, optimization of imaging protocols should be based on a comprehensive evaluation of the dosimetric characteristics and the performance of automatic exposure control of the employed fluoroscopic unit, preferably in terms of DAP rate [29].

1. Calkins H, Canby R, Weiss R, Taylor G, Wells P, Chinitz L, et al. 100W Atakr II Investigator Group. Results of catheter ablation of typical atrial flutter. Am J Cardiol 2004;94:437–42. 2. Lickfett L, Mahesh M, Vasamreddy C, Bradley D, Jayam V, Eldadah Z, et al. Radiation exposure during catheter ablation of atrial fibrillation. Circulation 2004;110:3003–10. 3. Perisinakis K, Damilakis J, Theocharopoulos N, Manios E, Vardas P, Gourtsoyiannis N. Accurate assessment of patient effective radiation dose and associated detriment risk from radiofrequency catheter ablation procedures. Circulation 2001;104:58–62. 4. Leclercq C, Hare JM. Ventricular resynchronization: current state of the art. Circulation 2004;109:296–9. 5. Roguin A, Bomma CS, Nasir K, Tandri H, Tichnell C, James C, et al. Implantable cardioverter-defibrillators in patients with arrhythmogenic right ventricular dysplasia/cardiomyopathy. J Am Coll Cardiol 2004;3:1843–52. 6. Leclercq C, Kass DA. Retiming the failing heart: principles and current clinical status of cardiac resynchronization. J Am Coll Cardiol 2002;39:194–201. 7. Boehmer JP. Device therapy for heart failure. Am J Cardiol 2003;91:53D–9D. 8. Lee KL, Hafley G, Fisher JD, Gold MR, Prystowsky EN, Talajic M, et al. Multicenter Unsustained Tachycardia Trial Investigators. Effect of implantable defibrillators on arrhythmic events and mortality in the multicenter unsustained tachycardia trial. Circulation 2002;106:233–8. 9. Izutani H, Quan KJ, Biblo LA, Gill IS. Biventricular pacing for congestive heart failure: early experience in surgical epicardial versus coronary sinus lead placement. Heart Surg Forum 2002;6:E1–6. 10. Balter S. Stray radiation in the cardiac catheterisation laboratory. Radiat Prot Dosim 2001;94:183–8. 11. Vano E. Radiation exposure to cardiologists: how it could be reduced. Heart 2003;89:1123–4. 12. International Commission on Radiological Protection. Recommendations of the International Commission on Radiological Protection, ICRP Publication 60. Ann ICRP 1991;21(1/3). 13. International Commission on Radiological Protection. Conversion coefficients for use in radiological protection against external radiation, ICRP Publication 74. Ann ICRP 1996;26(3/4). 14. International Commission on Radiation Units and Measurements. Measurement of dose equivalents from external photon and electron radiations, ICRU Report 47. Bethesda, MD: ICRU, 1992. 15. Rawlings DJ, Faulkner K, Harrison RM. Broad-beam transmission data in lead for scattered radiation produced at diagnostic energies. Br J Radiol 1991;64:69–71.

Limitations of the study The major limitations of the present study are that not all of the beam geometries possibly encountered in electrophysiological studies were simulated, and that no ceiling-suspended radiation barriers were available in our institution. However, the selected projections are the major constituents of any cardiac study. Moreover, the effect of commercially available removable radiation barriers should be individually examined at each institution, according to their specific use, with extensive radiation measurements regarding beam angulation and staff/barrier positioning. Furthermore, since the distribution and attenuation of scatter is likely to be 650

Radiation hazards in the electrophysiology laboratory should not be overstated nor ignored. Continuous risk assessment and minimization is required. Good fluoroscopy and radiation protection practices to mitigate occupational exposure to radiation should be proposed and implemented.


Occupational exposure in the electrophysiology laboratory 16. Niklason LT, Marx MV, Chan HP. The estimation of occupational effective dose in diagnostic radiology with two dosimeters. Health Phys 1994;67:611–5. 17. Theocharopoulos N, Perisinakis K, Damilakis J, Papadokostakis G, Hadjipavlou A, Gourtsoyiannis N. Occupational exposure from common fluoroscopic projections used in orthopaedic surgery. J Bone Joint Surg Am 2003;85:1698–703. 18. Wiegand UKH, Bode F, Bonnemeier H, Eberhard F, Schlei M, Peters W, et al. Long-term complication rates in ventricular single lead VDD and dual chamber pacing. Pacing Clin Electrophysiol 2003;26:1961–9. 19. Perisinakis K, Theocharopoulos N, Damilakis J, Manios E, Vardas P, Gourtsoyiannis N. Fluoroscopically guided implantation of modern cardiac resynchronization devices: radiation burden to the patient and associated risks. J Am Coll Cardiol 2005;46:2335–9. 20. Damilakis J Perisinakis K, Theocharopoulos N, Manios E, Vardas P, Gourtsoyiannis N. Anticipation of radiation dose to the conceptus from occupational exposure of pregnant staff during fluoroscopically guided electrophysiological procedures. J Cardiovasc Electrophysiol 2005;16: 773–80. 21. Kuon E, Schmitt M, Dahm JB. Significant reduction of radiation exposure to operator and staff during cardiac interventions by analysis of radiation leakage and improved lead shielding. Am J Cardiol 2002;9:44–9.


22. Kuon E, Dahm JB, Empen K, Robinson DM, Reuter G, Wucherer M. Identification of less-irradiating tube angulations in invasive cardiology. J Am Coll Cardiol 2004;44:1420–8. 23. Hall EJ. Radiobiology for the radiologist. Philadelphia, PA: JB Lippincott Co, 1994:363–78. 24. Calkins H, Niklason L, Sousa J, el-Atassi R, Langberg J, Morady F. Radiation exposure during radiofrequency catheter ablation of accessory atrioventricular connections. Circulation 1991;84:2376–82. 25. Lindsay BD, Eichling JO, Ambos HD, Cain ME. Radiation exposure to patients and medical personnel during radiofrequency catheter ablation for supraventricular tachycardia. Am J Cardiol 1992;70:218–23. 26. Vano E, Gonzalez L, Guibelalde E, Fernandez JM, Ten JI. Radiation exposure to medical staff in interventional and cardiac radiology. Br J Radiol 1998;71:954–60. 27. Balter S. Radiation safety in the cardiac catheterization laboratory: operational radiation safety. Catheter Cardiovasc Interv 1999;47:347–53. 28. Kuon E, Birkel J, Schmitt M, Dahm JB. Radiation exposure benefit of a lead cap in invasive cardiology. Heart 2003;89:1205–10. 29. Kuon E, Dorn C, Schmitt M, Dahm JB. Radiation dose reduction in invasive cardiology by restriction to adequate instead of optimized picture quality. Health Phys 2003;84:626–31.

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Usefulness of diffusion/perfusion-weighted MRI in patients with non-enhancing supratentorial brain gliomas: a valuable tool to predict tumour grading? G G FAN,

PhD, MD,

Q L DENG,

MD,

Z H WU,

MD

and Q Y GUO,

PhD, MD

Department of Radiology, Second Hospital of China Medical University, No.36 Sanhao St., Heping Dist., Shenyang, Liaoning,110004, People’s Republic of China ABSTRACT. 22 patients with non-enhancing supratentorial gliomas on contrastenhanced MRI underwent both diffusion- and perfusion-weighted MRI (DWI/PWI) before surgical resection. 14 low-grade gliomas (WHO Grade I and II) and 8 anaplastic gliomas were verified histologically. Both apparent diffusion coefficient (ADC) values and relative cerebral blood volume (rCBV) ratios were calculated on the solid portion of the tumour, on peritumoural area, as well as on the contralateral normal white matter, respectively. The results showed that lower ADC values were present in the solid portions of anaplastic gliomas, but not in low grade (p,0.01). All ADC values in peritumoural regions of tumours were decreased compared with the contralateral normal white matter. However, there was no significant difference between anaplastic gliomas and low-grade gliomas. Meanwhile, higher rCBV ratios were present in both solid portions and peritumoural regions of anaplastic gliomas, but not in low grade gliomas (p,0.01). In conclusion, non-enhancing brain gliomas with lower ADC values in the solid portions and higher rCBV ratios in both solid portions and peritumoural regions of tumours are significantly correlated with anaplasia. Therefore, DWI and PWI should be integrated in the diagnostic work-up of non-enhancing gliomas in order to predict grading.

Gliomas are the most common primary neoplasms of the brain, varying histologically from low grade to high grade [1]. MRI plays a crucial role in the evaluation of patients with gliomas [2]. The use of gadolinium-based contrast agents yields further improvement in the demonstration and detection of cerebral gliomas. Patterns and extent of contrast enhancement have been suggestive of a malignant potential [3]. However, this approach is limited because 14–45% of non-enhancing supratentorial gliomas are malignant (especially in older patients) and some enhancing gliomas (i.e. pilocytic astrocytoma) are benign [4, 5]. Moreover, large cerebral gliomas are often histopathologically heterogeneous and may have components of varying grades of malignancy within them. Hence, accurate pre-operative grading of gliomas and planning of adequate treatment strategies are often difficult with conventional MRI [5, 6]. Diffusion-weighted MRI (DWI), which is sensitive to the molecular diffusion of water, has been well established as a reliable non-invasive method for the early detection of cerebral ischaemic stroke, and DWI has been reported to be helpful in differentiating necrotic cavities associated with malignant gliomas from the benign ones [7]. The use of DWI to better characterize enhancing tumours and vasogenic oedema has been explored, but the results obtained have been conflicting [7, 8]. Recent developments in perfusion-weighted MRI (PWI) techniques have permitted the creation of relative cerebral blood volume (rCBV) maps, leading to the qualitative and quantitative assessment of tumour vascularity. 652

Received 9 January 2006 Revised 27 January 2006 Accepted 2 February 2006 DOI: 10.1259/bjr/25349497 ’ 2006 The British Institute of Radiology

These maps have helped in the assessment of tumour grade and in targeting the site of biopsy [8, 9]. Although DWI and PWI have been widely used in preoperative grading of gliomas, to the best of our knowledge, few studies evaluating the usefulness of diffusion and perfusion MRI solely in non-enhancing gliomas have been reported, although the utility of perfusion MRI in gliomas without enhancement on conventional contrastenhanced MRI was performed by Maia et al [10]. In addition, the efficiency of peritumour region in the grade assessment of cerebral gliomas is still being investigated. Our hypothesis was that DWI/PWI could provide additional useful information in the assessment and tumour grading of supratentorial glial neoplasms, which lacked contrast enhancement on pre-operative neuroimaging.

Materials and methods Patient selection and clinical data collection 22 patients (10 women and 12 men; median age 48.3 years, age range 38–68 years) with non-enhancing supratentorial gliomas on contrast-enhanced MRI underwent both diffusion/perfusion-weighted MRI (DWI/ PWI) before surgical resection. Informed consent was obtained from all patients prior to the investigation, and all procedures were performed under the approval of our institutional review board for clinical studies. 14 low The British Journal of Radiology, August 2006

Diffusion/perfusion weighted MRI and supratentorial gliomas

grade gliomas (WHO Grade I and II) and 8 high grade (WHO Grade III) were verified histologically. Low grade gliomas consisted of low grade astrocytomas (Grade I, n57), low grade oligodendrogliomas (Grade II, n54) and low grade mixed oligoastrocytomas (Grade II, n53); while high grade gliomas consisted of anaplastic astrocytomas (Grade III, n55) and anaplastic oligodendrogliomas (Grade III, n53).

MRI examination MRI examinations were performed on 1.5 T superconduction whole-body MR system (Intera Gyroscan; Philips Medical Systems, Best, The Netherlands) and an eight-channel Sensitivity Encoding (SENSE) head coil. After scout view MRI, the examination protocol consists of pre-contrast conventional MRI followed by DWI, PWI, and finally post-contrast T1 weighted images. Conventional MR images were obtained with T1 (spin echo (SE), 442/15) and turbo T2 (turbo spin echo (TSE), 3235/100) weighted spin echo sequences (both with a 2566192 matrix, 6 mm slice thickness and 2 averages). DWIs were acquired using single-shot echo-planar imaging (EPI) sequence at multiple levels. 18 slices of 6 mm thickness were obtained (repetition time (TR) 5000 ms, echo time (TE) 104 ms, field of view 40620, matrix size 2566128, b values of 0 and 1000 mm2 s21) in 3 orthogonal directions. For susceptibility-based PWI, the transitory signal loss during the bolus passage was detected with a T2* weighted fast field echo (FFE) EPI sequence (TR 232 ms, TE 25 ms, 6 slices with 6 mm slice thickness; matrix, 896128; and 1 average). 40 dynamic scans with a time resolution of 1.6 s per image were performed after intravenous bolus injection of 20 ml Gd-DTPA (Magnevist; Schering AG) at a flow rate of 4 ml s21 and a 20 ml saline flush.

cover the region of expected maximum rCBV on maps of rCBV. Additionally, we measured the ratio of maximum rCBV in either the tumour or peritumoural area of that in the contralateral normal white matter in order to standardize variations in each examination. All data obtained were summarized as the mean ¡ standard deviation (SD). The Student t-test was used to determine if there were statistically significant differences in both averaged ADC value and maximum rCBV ratio between anaplastic gliomas (Grade III) and low grade (Grade I and Grade II) gliomas. A p-value of less than 0.05 was considered to indicate statistical significance.

Results Conventional MRI revealed heterogeneous signal intensity of tumours with clear evidence of central necrosis in all 22 patients. Of all these patients, the solid tumour tissue exhibited typical hyperintense on T2 weighted images (T2WI) and hypointense on T1 weighted images (T1WI) (Figure 1). Six of the seven patients with Grade II tumours and all eight patients with Grade III tumours (Figure 2) were found to have peritumoural oedema and mass effect. Infiltration of the tumour producing nodular thickening of the grey matter was seen in three patients with Grade II tumours and five patients with Grade III tumours (Figures 3 and 4). All tumours, however, did not produce significant

MR data analysis and statistics The apparent diffusion coefficient (ADC) maps and values were calculated on a separate workstation (Easy Vision Intera workstation, release 8.1.3; Philips Medical Systems). We recorded the ADC values from both the solid portion of the tumour (seen as the highest signal intensity lesion at b value 0 and 1000) and peritumoural area. The ADC values in our study represented averaged ADCs of three to five regions of interest (ROIs). A ROI, varying from 40 mm3 to 60 mm3, was positioned carefully to avoid contamination from adjacent different tissues. The ROI was drawn as large as possible using a circular ROI on the workstation. As a control, the ADC was obtained from contralateral normal white matter. Raw PWIs were transferred to a PC workstation (Easy Vision Intera workstation, release 8.1.3; Philips Medical Systems) for post-processing. With the aid of the implemented software, the rCBV could be calculated on the basis of the indicator dilution method and were displayed as spectral colour images. ROI analyses were performed on the solid portion of the tumour, on peritumoural area, as well as on the contralateral normal white matter, respectively. ROI was placed carefully to The British Journal of Radiology, August 2006

Figure 1. MRI and diffusion weighted imaging (DWI) from a patient with anaplastic astrocytoma. (a) T2 weighted image. (b) T1 weighted MR image with contrast material. (c) DWI (b51000). (d) Apparent diffusion coefficient (ADC). The tumour exhibited typical hyperintense on T2 weighted images; peritumoural oedema and mass effect was present; no contrast enhancement. On DWIs (b51000), the signal intensity in the solid portion of the tumour was hyperintense with respect to the white matter; lower ADC values were present in the solid portions of high grade gliomas.

653

G G Fan, Q L Deng, Z H Wu and Q Y Guo

Figure 2. Perfusion weighted imaging (PWI) and histopathological photograph from a patient with anaplastic astrocytoma (same patient as in Figure 1). (a) Relative cerebral blood volume (rCBV) colour map. (b) Signal-intensity time-curve. (c) Histopathological photograph. rCBV maps were inhomogeneous with various increases of signal intensity in both solid portion and peritumoural region of tumour. Histopathological photograph of tumour confirmed the diagnosis of anaplastic astrocytoma.

contrast enhancement on visual assessment, suggesting lack of blood–brain barrier breakdown. On DWIs (b50, 1000), the signal intensity in the solid portion of the tumour was hyperintense with respect to the white matter (Figures 1, 3 and 5). The averaged ADC values for the solid tumour component ranged from

Figure 3. MRI and diffusion weighted imaging (DWI) from a patient of anaplastic astrocytoma. (a) T2 weighted image. (b) T1 weighted MR image with contrast material. (c) DWI (b51000). (d) Apparent diffusion coefficient (ADC). The tumour exhibited typical hyperintensity on T2 weighted images; infiltration of the tumour producing nodular thickening of the grey matter was observed. The tumour did not produce significant contrast enhancement on visual assessment, On DWIs (b51000), the signal intensity in the solid portion of the tumour was hyperintense with respect to the white matter; lower ADC values were present in the solid portions of high grade gliomas.

654

0.4361023 mm2 s21to 1.3961023 mm2 s21 (mean 0.89¡0.3661023 mm2 s21) and for white matter from 1.0261023 mm2 s21to 1.7761023 mm2 s21 (mean 1.40¡0.3561023 mm2 s21) (Table 1). Lower ADC values were present in the solid portions of high grade gliomas (Grade III) (0.52¡0.1161023 mm2 s21), but not in low grade (Grade I and Grade II) (1.15¡0.1661023 mm2 s21); the difference was statistically significant (p,0.01). All ADC values in peritumoural regions of tumours were decreased compared with the contralateral normal white matter; however, there was no significant difference between high grade (0.69¡0.1361023 mm2 s21) and low grade gliomas (0.79¡0.0761023 mm2 s21) (p.0.05). The signal intensity in the peritumoural region was homogeneous on rCBV map in all 14 cases of low-grade (Grade I and Grade II) tumours (Figure 6). However, a slight increase of signal intensity of the tumour was observed in three cases of Grade II tumours and in one case of Grade I tumour. In contrast, all rCBV maps for Grade III tumours were inhomogeneous with various increases of signal intensity in solid portions of tumour (Figures 2 and 4). Table 2 summarizes the measurements of maximum relative rCBV ratios of all non-enhancing gliomas. Measured maximum rCBV ratio in the solid portion of tumour varied from 0.76 to 5.01, with a mean of 2.02¡0.73 (¡SD); whereas in peritumoural region, maximum rCBV ratio varied from 0.92 to 3.55, with a mean of 1.79¡1.03 (¡SD). Higher maximum rCBV ratios were present in both solid portions and peritumoural regions of anaplastic gliomas (3.27¡1.54; 2.92¡0.44), but not in low grade (Grade I and Grade II) (1.52¡0.49; 1.36¡0.83) (p,0.01).

Discussion MRI plays a critical role in the pre-operative assessment of brain gliomas. Mass effect, cyst formation and necrosis on MRI studies do correlate significantly with The British Journal of Radiology, August 2006


Figure 4. Perfusion weighted imaging (PWI) and histopathological photograph from a patient with anaplastic astrocytoma (same patient as in Figure 3). (a) Relative cerebral blood volume (rCBV) colour map. (b) Signal-intensity time-curve. (c) Histopathological photograph. rCBV maps were inhomogeneous with various increases of signal intensity in solid portion of tumour. Histopathological photograph of tumour confirmed the diagnosis of anaplastic astrocytoma.

malignant behaviour [2, 3, 4, 6]. Traditionally, the extent of contrast enhancement has been used as a mark of malignancy: most high-grade gliomas on post-contrast T1 weighted images generally show moderate to strong enhancement; on the other hand, the low-grade gliomas have minimal or no enhancement [6, 10]. However, lack of contrast enhancement on MRI studies does not equate with low tumour grade. In fact, a broad spectrum of histological types may present as non-enhancing lesions [5]. Even after contrast administration, up to 25% of highgrade gliomas may show faint or no detectable enhancement, such as our cases (8 of 22 cases of anaplastic gliomas have been confirmed) [6, 11]. In addition, the

risk of anaplasia in non-enhancing lesions on MRI increases significantly with the patient’s age [12]. Therefore, in the case of tumours that do not enhance, pre-operative evaluation of tumour grade is often difficult on conventional MRI. DWI has been used by some to evaluate intra-axial tumours [7, 8, 13]. DWI and calculation of ADC values have been used to distinguish the normal white matter areas from necrosis, cyst formation, oedema, and solid tumour by measuring differences in ADC values caused by water proton mobility alterations [13, 14]. These differences are thought to result from both changes in the balance between intracellular and extracellular water and changes in the structure of the two compartments [14]. Our result showed that calculated ADC values from tumoural core added more information to MRI in the differentiation and grading of non-enhancing gliomas; these results suggested different ADC values due to different tumour grades [15]. Although ADC values of biological tissue are determined by many factors, results of previous studies have already confirmed that the lower ADC values in tumour core were mainly affected by tumour cellularity. With higher diffusivity found in the extracellular volume, the increase of intracellular space due to highly cellular tissue is coupled with a decrease of the ADC [8, 15, 16]. Therefore, higher cellularity in anaplastic neoplasm would contribute to the lower ADC values. Contrary to most previous studies, averaged ADC values in peritumoural area were also analysed in our Table 1. Apparent diffusion coefficient values comparison of non-enhancing gliomas with contralateral white matter (WM) (61023 mm2 s21) Grade

Cases

Solid tumoural region

Peritumoural Contralateral region WM

I II III

7 7 8

1.17¡0.14 1.01¡0.17 0.52¡0.11*

0.83¡0.16 0.74¡0.09 0.69¡0.13

Figure 5. MRI and diffusion weighted imaging (DWI) from a patient with low-grade oligodendrogliomas. (a) T1 weighted image. (b) T 2 weighted image. (c) DWI (b51000). (d) Apparent diffusion coefficient (ADC). The tumour exhibited inhomogeneous with various increase of signal intensity on both T2 weighted image and DWI.


1.44¡0.32 1.35¡0.28 1.40¡0.47

Data are the mean¡SD; * is statistical significance (p,0.01).

655

G G Fan, Q L Deng, Z H Wu and Q Y Guo

Figure 6. Perfusion weighted imaging (PWI) and histopathological photograph from a patient with low-grade oligodendrogliomas (same patient as in Figure 5). (a) Relative cerebral blood volume (rCBV) colour map. (b) Signal-intensity time-curve. (c) Histopathological photograph. The signal intensity in both peritumoural region and solid portion of tumour was homogeneous on rCBV map. Histopathological photograph of tumour confirmed the diagnosis of oligodendrogliomas (Grade II).

study. The results showed that ADC values in peritumoural regions were decreased compared with the contralateral normal white matter. It is assumed that malignant gliomas are not strictly focal lesions, but rather are characterized by intracerebral dissemination of malignant glial cells along the myelinized axons and blood vessels, or through the subarachnoid space [8]. We therefore speculated that there might be a difference in the ADC values of the peritumoural region between low grade and high grade gliomas. However, no significant difference was found. We postulate that the ADC value in the peritumoural region is unreliable as a means of differentiating between high grade and low grade gliomas, because a partial volume effect contaminated by surrounding oedema may affect the accurate outcome [8, 17]. Recently, usefulness of diffusion tensor imaging (DTI) in the study of peritumoural region of gliomas has been investigated and the result seems promising [18]. DTI, which is capable of visualizing the anisotropy of proton motion, may make it possible to eliminate the partial volume effect of peritumoural oedema, and the difference between high-grade and low-grade gliomas in this respect seems worthy of evaluation. For planning the optimal treatment strategy, accurate determination of tumour grade is critical, and in most histological grading systems, vascular proliferation of gliomas is a diagnostic criterion for malignancy [19]. New blood vessel growth is a critical phase of solid tumour growth. The growth of a solid tumour mass at 1– 2 mm3 depends upon simple diffusion of oxygen, Table 2. Maximum relative cerebral blood volume (rCBV) ratio in both tumoural and peritumoural regions of nonenhancing gliomas Grade

Cases

Solid tumoural region

Peritumoural region

I II III

7 7 8

1.09¡0.26 1.91¡0.78 3.27¡1.54*

1.27¡0.33 1.54¡0.60 2.92¡0.44*

Data are the mean¡SD; * is statistical significance (p,0.01).

656

nutrients and other essential materials. However, tumour mass growth over 1–2 mm3 can not occur and metastasise without angiogenesis [20]. Although conventional MRI with gadolinium-based contrast enhancement has been useful for grading gliomas, contrast enhancement itself reflects disruption of the blood–brain barrier, not tumour angiogenesis [3, 5, 6]. PWI techniques now have been used for the assessment of tumour vascularity in vivo. rCBV maps and measurements have been shown to correlate reliably with tumour grade and histological findings of tumoural microvascular density, the current standard for assessing the degree of angiogenesis [8–11]. The correlation between the histopathological grade of cerebral gliomas and rCBV has already been evaluated by various groups. Although these studies showed a wide range of rCBV ratios and overlapping between tumours of different grades, there were statistically significant differences between high grade and low grade gliomas [8–11], even in the study of non-enhancing gliomas, as our results confirm. In our study, nonenhancing anaplastic gliomas often demonstrated significant heterogeneity and areas of high rCBV. The presence of contrast enhancement on conventional MRI only represents a pathological alteration in the blood–brain barrier (with or without concomitant angiogenesis), whereas the degree of perfusion MR abnormality can truly reflect the degree of angiogenesis (with or without destruction of the blood–brain barrier) [6, 21]. Therefore, the advantage of perfusion MRI over contrastenhanced MRI is in depicting tumour angiogenesis and hence in pre-operative grading. Moreover, because large cerebral gliomas are often histopathologically heterogeneous, areas with higher rCBV values, which may be regarded as greater tumour vascularity, can be selectively targeted by stereotactic biopsies to reduce tumour under-grading [6, 22]. This is especially true for nonenhancing gliomas with relative intact blood–brain barrier [10]. On the contrary, for high grade enhancing gliomas with concomitant breakdown of the blood–brain barrier, the first pass of contrast material may leak into extravascular space, and thus the produced The British Journal of Radiology, August 2006


susceptibility effects may be decreased between intravascular and extravascular space near the disrupted blood–brain barrier, which is considered to cause the underestimation of the true tumour vascularity [8]. In contrast to measuring average rCBV within tumoural core alone, the measurement of maximum rCBV in peritumoural regions was also performed in our study. The results showed that elevated rCBV ratios were also present in peritumoural brain regions in high grade gliomas, suggests increased peritumoural perfusion due to tumour infiltration. In anaplastic tumours, peritumoural areas demonstrate not only altered capillary morphologic findings but also scattered tumour cells infiltrating along newly formed or pre-existing but dilated vascular channels [8, 22, 23]. In low grade gliomas, on the other hand, the peritumoural region contains less infiltrating tumour cells. This interpretation is consistent with elevated blood volume preceding the appearance of enhancement, which reflects blood–brain barrier breakdown [22]. The fact that we found comparable elevated rCBV in the peritumoural area of grade III gliomas is both novel and significant: Information regarding heterogeneity of peritumoural region in terms of vascularity as depicted by PWI can be effectively used to best estimate of the true brain tumour size pre-operatively [24]. In PWI study, we chose spin-echo echo-planar images because of their presumed higher sensitivity in detecting tumour vascularity at capillary level than at large vessel level [10]. Two different sequences, including spin echo and gradient-echo echo-planar sequences are generally used in first-pass perfusion MR study [21]. Because the gradient-echo echo-planar technique is sensitive to susceptibility effects from the total volume of blood contained in both capillaries and large vessels, we prefer spin echo echo-planar sequences in PWI study of nonenhancing gliomas in order to eliminate the interference of susceptibility artefact [25]. In conclusion, usefulness of DWI and PWI in nonenhancing cerebral gliomas is not only feasible, but also offers clinically relevant physiological data not obtainable by conventional MRI. Non-enhancing brain gliomas with lower ADC values in the solid portions and higher rCBV ratios in both solid portions and peritumoural regions of tumours are significantly correlated with anaplasia. Therefore, DWI and PWI should be integrated in the diagnostic work-up of non-enhancing gliomas in order to predict grading. However, with the advent of advanced MR techniques, a more sophisticated study using a larger sample is needed in the near future.

Acknowledgments We wish to thank Dr Bing Yu and Dr Songmin Quan for their expert neuroradiological opinion and assistance, and Prof. Liying Chen for helpful advice and discussion. The authors also thank Department of Neurosurgery of China Medical University for their fellowship.

References 1. Daumas-Duport C, Scheithauer B, O’Fallon J, Kelly P. Grading of astrocytomas: a simple and reproducible method. Cancer 1988;62:2152–65.


2. Rees J. Advances in magnetic resonance imaging of brain tumours. Curr Opin Neurol 2003;16:643–50. 3. Kurki T, Lundbom N, Kalimo H, Valtonen S. MR classification of brain gliomas: value of magnetization transfer and conventional imaging. Magn Reson Imaging 1995;13:501–11. 4. Mihara F, Numaguchi Y, Rothman M, Kristt D, Fiandaca M, Swallow L. Non-enhancing supratentorial malignant astrocytomas: MR features and possible mechanisms. Radiat Med 1995;13:11–7. 5. Scott JN, Brasher PM, Sevick RJ, Rewcastle NB, Forsyth PA. How often are nonenhancing supratentorial gliomas malignant? A population study. Neurology 2002;59:947–9. 6. Batra A, Tripathi RP, Singh AK. Perfusion magnetic resonance imaging and magnetic resonance spectroscopy of cerebral gliomas showing imperceptible contrast enhancement on conventional magnetic resonance imaging. Australas Radiol 2004;48:324–32. 7. Tien RD, Felsberg GJ, Friedman H, Brown M, MacFall J. MR imaging of high-grade cerebral gliomas: value of diffusionweighted echoplanar pulse sequences. AJR Am J Roentgenol 1994;162:671–7. 8. Fan G, Zang P, Jing F, Wu Z, Guo Q. Usefulness of diffusion/perfusion-weighted MRI in rat gliomas: correlation with histopathology. Acad Radiol 2005;12:640–51. 9. Law M, Yang S, Babb JS, et al. Comparison of cerebral blood volume and vascular permeability from dynamic susceptibility contrast-enhanced perfusion MR imaging with glioma grade. AJNR Am J Neuroradiol 2004;25:746–55. 10. Maia AC Jr, Malheiros SM, da Rocha AJ, et al. MR cerebral blood volume maps correlated with vascular endothelial growth factor expression and tumor grade in nonenhancing gliomas. AJNR Am J Neuroradiol 2005;26:777–83. 11. Law M, Yang S, Wang H, et al. Glioma grading: sensitivity, specificity, and predictive values of perfusion MR imaging and proton MR spectroscopic imaging compared with conventional MR imaging. AJNR Am J Neuroradiol 2003;24:1989–98. 12. Bruner JM. Neuropathology of malignant gliomas. Semin Oncol 1994;21:126–38. 13. Bulakbasi N, Guvenc I, Onguru O, Erdogan E, Tayfun C, Ucoz T. The added value of the apparent diffusion coefficient calculation to magnetic resonance imaging in the differentiation and grading of malignant brain tumors. J Comput Assist Tomogr 2004;28:735–46. 14. Bulakbasi N, Kocaoglu M, Ors F, Tayfun C, Ucoz T. Combination of single-voxel proton MR spectroscopy and apparent diffusion coefficient calculation in the evaluation of common brain tumors. AJNR Am J Neuroradiol 2003;24:225–33. 15. Hein PA, Eskey CJ, Dunn JF, Hug EB. Diffusion-weighted imaging in the follow-up of treated high-grade gliomas: tumor recurrence versus radiation injury. AJNR Am J Neuroradiol 2004;25:201–9. 16. Sugahara T, Korogi Y, Kochi M, et al. Usefulness of diffusion-weighted MRI with echo-planar technique in the evaluation of cellularity in gliomas. J Magn Reson Imaging 1999;9:53–60. 17. Pauleit D, Langen KJ, Floeth F, et al. Can the apparent diffusion coefficient be used as a noninvasive parameter to distinguish tumor tissue from peritumoral tissue in cerebral gliomas? J Magn Reson Imaging 2004;20:758–64. 18. Lu S, Ahn D, Johnson G, Law M, Zagzag D, Grossman RI. Diffusion-tensor MR imaging of intracranial neoplasia and associated peritumoral edema: introduction of the tumor infiltration index. Radiology 2004;232:221–8. 19. Zhen HN, Zhang X, Hu PZ, et al. Survivin expression and its relation with proliferation, apoptosis, and angiogenesis in brain gliomas. Cancer 2005;104:2775–83.

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G G Fan, Q L Deng, Z H Wu and Q Y Guo 20. Bello L, Giussani C, Carrabba G, Pluderi M, Costa F, Bikfalvi A. Angiogenesis and invasion in gliomas. Cancer Treat Res 2004;117:263–84. 21. Hakyemez B, Erdogan C, Ercan I, Ergin N, Uysal S, Atahan S. High-grade and low-grade gliomas: differentiation by using perfusion MR imaging. Clin Radiol 2005;60: 493–502. 22. Lupo JM, Cha S, Chang SM, Nelson SJ. Dynamic susceptibility-weighted perfusion imaging of high-grade gliomas: characterization of spatial heterogeneity. AJNR Am J Neuroradiol 2005;26:1446–54.

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23. Fan G, Sun B, Wu Z, Guo Q, Guo Y. In vivo single-voxel proton MR spectroscopy in the differentiation of high-grade gliomas and solitary metastases. Clin Radiol 2004;59:77–85. 24. Bulakbasi N, Kocaoglu M, Farzaliyev A, Tayfun C, Ucoz T, Somuncu I. Assessment of diagnostic accuracy of perfusion MR imaging in primary and metastatic solitary malignant brain tumors. AJNR Am J Neuroradiol 2005;26:2187–99. 25. Sugahara T, Korogi Y, Kochi M, Ushio Y, Takahashi M. Perfusion-sensitive MR imaging of gliomas: comparison between gradient-echo and spin-echo echo-planar imaging techniques. AJNR Am J Neuroradiol 2001;22:1306–15.



Imaging well-differentiated hepatocellular carcinoma with dynamic triple-phase helical computed tomography 1,3

C-S LI, 1,3,4R-C CHEN, 1H-Y TU, 2L-S SHIH, 2T-A ZHANG, 1J-M LII, 1W-T CHEN, 1S-J DUH and 1 L-C CHIANG Departments of 1Radiology, 2Pathology, Renai Branch, Taipei City Hospital, Taipei, 3Department of Radiology, School of Medicine, National Yang-Ming University, Taipei and 4Department of Radiology, School of Medicine, Taipei Medical University, Taipei, Taiwan ABSTRACT. To investigate the imaging appearance of well-differentiated hepatocellular carcinoma (HCC) on dynamic CT, a total of 38 histopathologically proven well-differentiated HCC were included in a retrospective study. We reviewed the contrast-enhanced dynamic CT of all 38 tumours for attenuation of each tumour in unenhanced scan, arterial-dominant and delayed portal venous phases. Our results showed that dynamic CT identified 26 (68.4%) out of the 38 lesions. The remaining 12 lesions were isodense compared with surrounding liver parenchyma in each dynamic CT phase. There was no statistically significant difference between the mean size of tumours detected by dynamic CT and that of tumours not detected by dynamic CT (p50.1). Of a total of 38 tumours, most were isodense (n519) or hypodense (n516) in unenhanced scan, mostly hyperdense (n518) or isodense (n515) in arterial-dominant phase and mostly isodense (n522) or hypodense (n515) in delayed portal venous phase. Enhancement of tumour was observed in 19 (50.0%) of 38 lesions. In conclusion, the ability of dynamic CT to detect welldifferentiated HCC is poor, and negative CT findings cannot exclude the presence of well-differentiated HCC, especially if there is well-grounded clinical suspicion for HCC.

Hepatocellular carcinoma (HCC) is one of the most common malignancies in the world. Globally, there is an increasing incidence of HCC in both eastern and western countries [1, 2]. HCC usually develops in the setting of chronic liver disease and cirrhosis [3]. Screening with ultrasound and a-fetoprotein levels to detect HCC in patients with chronic liver disease has become common practice [4, 5]. Consequently, an increasing number of nodular hepatic lesions has been detected on screening ultrasound. It is important to distinguish well-differentiated HCC, an early form of HCC in hepatocarcinogenesis, from dysplastic hepatic nodule [3, 6]. Treatment should be applied at this stage with local ablation therapy, surgical resection or transplantation [7]. The biological behaviour of well-differentiated HCC is uncertain, but it is thought to have relatively low malignant potential and rarely invade vessels or metastasise to other sites [8, 9]. When hepatic nodules are identified on ultrasound, CT or MRI is usually used for further characterization because ultrasound appearance of a hepatic nodule may be non-specific [10]. However, imaging characteristics of well-differentiated HCCs have been described in only a few papers in the English literature [11–15]. In this retrospective study, we would like to investigate the imaging appearance of well-differentiated HCC on dynamic dual-phase helical CT. The British Journal of Radiology, August 2006

Received 7 November 2005 Revised 26 January 2006 Accepted 13 February 2006 DOI: 10.1259/bjr/12699987 ’ 2006 The British Institute of Radiology

Materials and methods Patients From January 2000 to July 2004, we retrospectively reviewed the histopathological reports of ultrasoundguided liver tumour biopsy in our hospital. A total of 68 patients with 71 well-differentiated HCCs were identified. All patients received abdominal ultrasound for screening of HCC or follow-up after treatment for HCC. Among this group, 37 patients with 38 well-differentiated HCCs were referred for dynamic CT scan because of hepatic nodules detected on abdominal ultrasound, and they were included in our study. They comprised 25 men and 12 women, with ages ranging from 31 years to 82 years (mean age 59.8 years). All patients suffered from chronic liver disease including hepatitis B in 22 patients, hepatitis C in 12 patients, alcoholic liver disease in 1 patient and cryptogenic liver cirrhosis in 2 patients. The dynamic CT images of each patient were reviewed. Written informed consent was obtained from each patient. All 37 patients underwent ultrasound-guided biopsy of hepatic nodules before or soon after (within 1 month) CT scan. All biopsies were performed with a 3.5 MHz guiding probe with a Logiq 400 unit (GE pro series). The biopsy needle was either a 1.2 mm Surecut needle (Top Surgical Manufacturing, Tokyo, Japan) or a 0.8 mm 659

C-S Li, R-C Chen, H-Y Tu et al

Majima needle (Top Surgical Manufacturing, Tokyo, Japan). The needle biopsy specimens were fixed in formalin immediately and processed according to standard procedures in the department of pathology. Both haematoxylin and eosin (H&E) and reticulin stains were routinely performed on each case. The specimens were interpreted by one of the two authors (LSS and TAZ); both of them were experienced pathologists in the diagnosis of differentiation of HCC. The histopathological criteria of well-differentiated HCC included: (1) increased cell density (more than 1.8 times nonneoplastic hepatic tissue of the same patient); (2) increased nuclear/cytoplasmic ratio of individual tumour cells; and (3) disorganized reticulin framework with irregular thin-trabecular pattern [16]. In most cases, the cancer cells of well-differentiated hepatocellular carcinoma were smaller in size with increased cytoplasmic eosinophilia. In addition, fatty changes of tumour tissue were occasionally observed [17]. Correlation between imaging characteristics and histopathological evidence of fatty change and vascular density was not a focus of our study, because the histopathological diagnosis in our study was based on liver biopsy.

Image acquisition Hepatic CT scan was performed with a helical CT scanner (HiSpeed Advantage; General Electric Medical Systems, Milwaukee, WI). All patients received oral contrast material before the CT examination. The unenhanced scans of entire liver were performed with slice thickness of 7 mm and space of 10 mm. The enhanced CT scans were then performed by intravenous injection of 100–120 ml iopromide (Ultravist 370; Schering, Berlin, Germany), depending on patients’ body weight. A mechanical power injector was used. The rate of injection was 2.5–3.5 ml s21 depending on accessibility to patients’ peripheral veins. The length of delay between intravenous contrast material administration and scanning was 30–35 s for the arterial-dominant phase, according to cardiac function estimated by patient’s age and general medical condition. The delayed portal venous phase was obtained at 110 s after contrast material administration. The delayed portal venous phase, in-between the usual portal venous phase and delayed phase, was the standard dynamic CT protocol for evaluation of nodular liver lesions in our institution. The dual-phase enhanced CT scan was performed with a slice thickness of 7 mm and a pitch of 1.2:1. Images acquired during each phase of contrast were obtained using a breath-hold technique and required two split acquisitions to cover the entire liver. For most cases, a total of 15 slices were obtained in each acquisition and there was a 6 s interval between the two acquisitions.

Image interpretation The ultrasound features of each patient in this study were carefully correlated with the CT scan findings in respect of tumour location. The tumour sizes were determined by the maximal diameter of each tumour 660

on the basis of ultrasound. The attenuations of these lesions in unenhanced scan, arterial-dominant phase and delayed portal venous phase of dynamic CT scan were independently analysed by two experienced radiologists specialized in liver imaging. The attenuations of these lesions were compared with surrounding liver and were categorized as hyperdense (or high-density), isodense (or iso-density) and hypodense (or low-density) by visual inspection. Identification of extremely low attenuation representing the fat component within these lesions was also noted. Positive enhancement of tumour was defined as when a hypodense tumour in unenhanced scan becomes isodense or hyperdense in arterial-dominant phase, or when an isodense tumour in unenhanced scan turns hyperdense in arterial-dominant phase. The window levels of images were kept constant during interpretation. Disagreements of images interpretation were resolved by consensus.

Statistical analysis All 38 tumours in the study were subdivided into two groups, one that could be visualized by CT scan and another group that could not. The size of tumour in each group was expressed as mean ¡ standard error (SE). Comparison of sizes between different groups was performed using the Student’s t-test. A threshold p-value of 0.05 was chosen for statistical significance.

Results The sizes of these 38 well-differentiated HCCs detected by ultrasound ranged from 0.5 cm to 2.7 cm (mean: 1.7¡0.6 cm). 27 (71.1%) out of 38 tumours were small HCCs with the sizes less than 2 cm. Dynamic CT scan depicted 26 (68.4%) out of the 38 nodules. The remaining 12 nodules were isodense in unenhanced, arterial-dominant and delayed portal venous phases of dynamic CT scan (we coded this as ‘‘iso-iso-iso’’ pattern) and, as a result, could not be identified (Figure 1). The mean size of the 26 nodules detected by dynamic CT scan was 1.8¡0.6 cm; and that of the remaining 12 nodules not identified by dynamic CT scan was 1.5 ¡0.4 cm. There was no statistically significant difference in size between the two groups of nodules (p50.1). Besides that, there was no statistically significant difference in size between all 38 nodules and the group of 26 nodules that could be visualized by CT scan (p50.4), as well as the group of 12 nodules that could not be visualized by CT scan (p50.3). In unenhanced CT scans, these 38 well-differentiated HCCs were mostly isodense (n519, 50.0%) or hypodense (n516, 42.1%); in arterial-dominant phase, they were mostly hyperdense (n518, 47.4%) or isodense (n515, 39.5%); in delayed portal venous phase, they were mostly isodense (n522, 57.9%) or hypodense (n515, 39.5%) (Table 1). Positive enhancement in arterial-dominant phase was observed in 19 (50.0%) of 38 tumours (Figure 2). One of the lesions was hypodense and enhanced to be isodense on the arterial-dominant phase images. All of the 19 tumours demonstrated washout of enhancement in delayed portal venous phase. The British Journal of Radiology, August 2006

Dynamic CT of well-differentiated HCC

(a)

(b)

(c)

(d)

Figure 1. Images of a 1.1 cm-diameter well-differentiated hepatocellular carcinoma in segment IV of the liver in a 73-year-old man. (a) Abdominal ultrasound disclosed a hepatic tumour with high echogenicity (arrow). (b,c) The tumour was isodense to surrounding liver parenchyma in unenhanced (not shown), arterial-dominant phase (b) and portal venous phase (c) of dynamic CT and, as a result, could not be identified. (d) Photomicrograph of the needle biopsy specimen showed a well-differentiated hepatocellular carcinoma. Haematoxylin and eosin, original magnification 6100.

In our study, the most common appearance of welldifferentiated HCC in each phase of dynamic CT was the ‘‘iso-iso-iso’’ pattern (n512, 31.6%). These tumours could not be identified on dynamic CT. The other welldifferentiated HCC in our study presented varying imaging appearances, and common patterns included: a ‘‘low-high-low’’ pattern (n57, 18.4%) (Figure 3), an ‘‘iso-high-iso’’ pattern (n55, 13.2%), a ‘‘low-high-iso’’ pattern (n54, 10.5%), and a ‘‘low-low-low’’ pattern (n54, 10.5%).

Discussion Early detection of HCC is becoming feasible owing to the wide use of ultrasound for screening [18]. High rates of ultrasound detection for small HCCs, ranging from 89% to 98.6%, have been reported [19, 20]. However, the differentiation of a hepatic nodule between early malignant tumour and pre-malignant nodule is often The British Journal of Radiology, August 2006

difficult. The distinction between well-differentiated HCC and dysplastic nodule is clinically important because of the difference in their management [10]. Therefore, advanced imaging modalities such as CT and MRI play an important role in characterizing these hepatic nodules in a non-invasive fashion. The appearance of well-differentiated HCC on CT scan has been studied before, but these prior studies either had too small a case number of well-differentiated HCCs or were lacking in dynamic CT protocols [5, 13, 14, 21]. Takayasu et al reported a poor detection rate of 56% for well-differentiated HCC, which was based on mixed table incremental CT and helical CT protocols [5]. Our study presented a 68.4% detection rate on the basis of dynamic helical CT protocol. The detection rate in our study was still low. Our study revealed that the mean size of well-differentiated HCC undetected by dynamic CT was not statistically different from tumours detected by dynamic CT. The reason behind the low detection rate was that well-differentiated HCC was commonly 661


(a)

(b)

(c)

(d)

Figure 2. Dynamic CT images of a 1.5 cm-diameter well-differentiated HCC in segment II of the liver in a 43-year-old man. (a) The tumour was hypodense in unenhanced scan (arrow), (b) hyperdense in arterial-dominant phase (arrow) and (c) isodense in portal venous phase. (d) Photomicrograph of the needle biopsy specimen with reticulin stain showed a disorganized reticulin framework with irregular thin-trabecular pattern. Original magnification 6100.

isodense in unenhanced, arterial-dominant and delayed portal venous phases of dynamic CT scan. Only half of our cases showed tumour enhancement in arterial-dominant phase of dynamic CT scan, which means the other half were hypovascular hepatocellular carcinoma. Well-differentiated HCC had less neoplastic angiogenesis and incomplete vascularization of the sinusoid-like blood spaces of the tumour, and therefore frequently presented as a hypovascular tumour [22–25]. In contrast, most moderately- and poorly-differentiated HCCs present as hypervascular lesions, and are easier to detect on dynamic CT [21]. Choi et al considered that combination of arterial and portal venous phases on dynamic CT was enough for detecting HCC, and delayed phase images could be 662

omitted to decrease scanning time and radiation hazard [26]. However, other authors suggested that delayed phase CT images were important in detecting early or hypovascular hepatocellular carcinoma [5, 27, 28]. The CT protocols used in these studies were composed of arterial-dominant phase of 30–35 s, portal venous phase of 60 s and delayed phase of 180 s after commencing intravenous contrast medium administration [27, 28]. In our institution, we used arterial-dominant phase to detect hypervascular lesions, followed by delayed portal venous phase at 110 s after contrast medium administration. The purpose of our dual-phase CT protocol was to avoid excessive radiation exposure of triple-phase CT; and our study showed that the CT protocol could offer a comparable (68.4%) detection rate for well-differentiated The British Journal of Radiology, August 2006


(a)

(b)

(c) Figure 3. Dynamic CT images of a 2.1 cm-diameter well-differentiated HCC in segment III of the liver in a 54-year-old man. (a) The tumour was hypodense in unenhanced scan (arrow). (b) Positive enhancement of the tumour was demonstrated in arterial-dominant phase (arrow). (c) In portal venous phase, the tumour was again hypodense (arrow), representing the ‘‘lowhigh-low’’ pattern on dynamic CT scan.

HCC. According to the results of our study, roughly half of well-differentiated HCCs are hypovascular; we hypothesized that dynamic CT scan with additional delayed-phase images might increase the detection rate of well-differentiated HCC. However, the hazard of increased radiation exposure should also be a concern. Further study with the triple-phase dynamic CT protocol to detect well-differentiated HCCs may be needed. It is clinically important to distinguish pre-malignant hepatic dysplastic nodules from well-differentiated HCCs. However, our study showed some overlapping dynamic CT appearances between the well-differentiated The British Journal of Radiology, August 2006

HCCs and hepatic dysplastic nodules. Choi et al had reported that hepatic dysplastic nodules were mostly isodense or hypodense in arterial-phase and portalphase CT images; most of them were relatively avascular [7]. Compared with radiographic features demonstrated in our study, high density or positive enhancement of a small hepatic nodule in arterial-dominant phase of dynamic CT could be an important feature pointing to the diagnosis of a HCC rather than a pre-malignant dysplastic nodule. However, in the appropriate clinical settings (chronic B or C viral hepatitis, elevation of serum a-fetoprotein level, positive ultrasound findings), 663


negative CT finding may not be sufficient to exclude a well-differentiated HCC. Further imaging study or biopsy should be considered in such situations. Lee et al reported a 96% arterial-phase enhancement rate of HCC, using triple-phase multidetector CT (MDCT) as well as rapid injection of contrast material at 5 ml s21 [29]. However, only a small number of welldifferentiated HCCs were presented in their study. MDCT, with its rapid scanning time, should have better detection ability than single-detector scanner; but the role of MDCT in diagnosis of well-differentiated HCC still needs further investigation. Dynamic gadolinium-enhanced MRI has been considered the method of choice for HCC diagnosis [30–32]. In cirrhotic patients, dynamic gadolinium-enhanced MRI is also a useful modality for detection and characterization of regenerative or dysplastic nodules [33]. Through its superior tissue contrast, MRI might detect more welldifferentiated HCCs based on signal intensity change on pre-contrast T1 weighted or T2 weighted images than CT [11, 13, 34]. Amano et al studied the CT and MRI patterns of HCC and concluded that they were useful in predicting the degree of histological differentiation of cancer cells in HCC [35]. However, no specific CT patterns or MR signals can offer accurate diagnosis of well-differentiated HCC, and absence of arterial phase enhancement cannot exclude the possibility of early or ‘‘borderline’’ malignancy [32]. In recent years, superparamagnetic iron oxide (SPIO)-enhanced MRI has been reported as useful for the detection of hepatic tumours [36, 37]. SPIO-enhanced MRI provided a fundamentally different approach for HCC diagnosis: SPIO would be taken up by Kupffer cells in normal liver parenchyma, but not in hepatic tumours which generally lacked reticulo-endothelial cells. However, the diagnosis of well-differentiated HCC could still be problematic because well-differentiated HCC may contain Kupffer cells as in normal liver parenchyma. Consequently, it may show similar SPIO uptake comparable with surrounding normal liver parenchyma and would be poorly delineated on SPIO-enhanced MRI [30, 38, 39]. Contrastenhanced MRI could be a better imaging modality for diagnosis of well-differentiated HCC than CT, but further investigation would be needed to achieve better sensitivity and specificity. We recognized two limitations in our study. First, this is a retrospective study in which there might be case selection bias, with respect to referral and patients undergone CT scanning. Second, there is a limitation in the correlation between CT appearances and histopathological findings in our study. All of our cases only had needle biopsy specimen for histopathological diagnosis, and underwent percutaneous ethanol injection therapy (PEIT) or transcatheter arterial chemoembolisation (TACE) for subsequent treatment. HCC may have different grades of differentiation within the tumour and the biopsy sample may not be representative of the whole tumour. In a study from Japan [24], about 20% of small HCCs (less than 2 cm in diameter) of the distinctly nodular type were composed of varying mixtures of well-differentiated and moderately-differentiated cancerous tissues. Needle biopsy of a tumour cannot fully reflect the differentiation of HCC. Better imaging-pathology 664

correlation could be obtained if we were able to examine the entire tumours.

Conclusion In conclusion, the detecting ability of dynamic dualphase helical CT for well-differentiated HCC is low. This type of tumour is frequently isodense to surrounding liver parenchyma in unenhanced, arterial-dominant or portal venous phase of CT scan, and positive enhancement is seen only in 50% of them. It is difficult to diagnose well-differentiated HCCs based solely on the density changes on CT scans. Hepatic nodules with high density or positive enhancement in arterial-dominant phase of dynamic CT scans suggest a diagnosis of a HCC rather than a hepatic dysplastic nodule in cirrhotic liver. However, in the appropriate clinical settings, a negative CT scan cannot be relied upon in exclusion of welldifferentiated HCCs. Further imaging study, including contrast enhanced-MDCT or MRI, and liver biopsy should be considered in this situation.

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26. Choi BI, Lee HJ, Han JK, Choi DS, Seo JB, Han MC. Detection of hypervascular nodular hepatocellular carcinomas: value of triphasic helical CT compared with iodizedoil CT. AJR Am J Roentgenol 1997;168:219–24. 27. Lim JH, Choi D, Kim SH, Lee SJ, Lee WJ, Lim HK, et al. Detection of hepatocellular carcinoma: value of adding delayed phase imaging to dual-phase helical CT. AJR Am J Roentgenol 2002;179:67–73. 28. Iannaccone R, Laghi A, Catalano C, Rossi P, Mangiapane F, Murakami T, et al. Hepatocellular carcinoma: role of unenhanced and delayed phase multi-detector row helical CT in patients with cirrhosis. Radiology 2005;234:460–7. 29. Lee KH, O’Malley ME, Haider MA, Hanbidge A. Triplephase MDCT of hepatocellular carcinoma. AJR Am J Roentgenol 2004;182:643–9. 30. Simon G, Link TM, Wortler K, Doebereiner F, SchulteFrohlinde E, Daldrup-Link H, et al. Detection of hepatocellular carcinoma: comparison of Gd-DTPA- and ferumoxides-enhanced MR imaging. Eur Radiol 2005;15:895–903. 31. Peterson M, Baron R, Murakami T. Hepatic malignancies: usefulness of acquisition of multiple arterial and portal venous phase images at dynamic gadolinium-enhanced MR imaging. Radiology 1996;201:337–45. 32. Rode A, Bancel B, Douek P, Chevallier M, Vilgrain V, Picaud G, et al. Small nodule detection in cirrhotic livers: evaluation with US, spiral CT, and MRI and correlation with pathologic examination of explanted liver. J Comput Assist Tomogr 2001;25:327–36. 33. Martin DR, Semelka RC. Magnetic resonance imaging of the liver: review of techniques and approach to common diseases. Semin Ultrasound CT MR 2005;26:116–31. 34. Kajiwara M. MR imaging of small hepatocellular carcinoma (, or 5 20 mm)--correlation with vascularity and histological features. Kurume Med J 1997;44:327–8. 35. Amano S, Ebara M, Yajima T, Fukuda H, Yoshikawa M, Sugiura N, et al. Assessment of cancer cell differentiation in small hepatocellular carcinoma by computed tomography and magnetic resonance imaging. J Gastroenterol Hepatol 2003;18:273–9. 36. Paley MR, Mergo PJ, Torres GM, Ros PR. Characterization of focal hepatic lesions with ferumoxides-enhanced T2weighted MR imaging. AJR Am J Roentgenol 2000;175:159–63. 37. Tang Y, Yamashita Y, Arakawa A, Namimoto T, Mitsuzaki K, Abe Y, et al. Detection of hepatocellular carcinoma arising in cirrhotic livers: comparison of gadolinium- and ferumoxides-enhanced MR imaging. AJR Am J Roentgenol 1999;172:1547–54. 38. Lim JH, Choi D, Cho SK, Kim SH, Lee WJ, Lim HK, et al. Conspicuity of hepatocellular nodular lesions in cirrhotic livers at ferumoxides-enhanced MR imaging: importance of Kupffer cell number. Radiology 2001;220:669–76. 39. Kato H, Kanematsu M, Kondo H, Goshima S, Matsuo M, Hoshi H, et al. Ferumoxide-enhanced MR imaging of hepatocellular carcinoma: correlation with histologic tumor grade and tumor vascularity. J Magn Reson Imaging 2004;19:76–81.

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Building an anonymized catalogued radiology museum in PACS: a feasibility study A P TOMS,

FRCS, FRCR,

B KASMAI,

MSc, MIPEM,

S WILLIAMS,

MRCP, FRCR

and P WILSON,

MRCP, FRCR

Department of Radiology, Norfolk & Norwich University Hospital NHS Trust, Colney Lane, Norwich, Norfolk NR4 7UY, UK ABSTRACT. The aim of this study was to test the feasibility of a software application that would allow the anonymization and cataloguing of whole DICOM datasets in order to build searchable radiology museums within PACS. The application was developed on a dedicated networked PC, using C# and HL7 coding. Whole DICOM datasets were pushed from PACS to a networked PC on which the application, Museum Builder, was developed. Museum Builder works by replacing the patient specific data (the forename, surname and hospital number) within each header of each DICOM file with terms from anatomical and surgical sieve menus. The date of birth is anonymized to 1 January of the same year. Whole DICOM datasets comprising hundreds of images can be anonymized and catalogued in a single episode. Museum Builder primes PACS with an HL7 script to receive a ‘‘new’’ patient. DICOM datasets are then pushed back to PACS where they are added to the database as ‘‘new’’ cases. The museum cases can then be searched for, on PACS, by any combination of terms that correspond to appropriate anatomical units, surgical sieve headings or radiological specialty. New radiology reports containing clinical histories, radiological descriptions, differential diagnoses and discussion can be added through the report window. Our institution has developed and used this tool to generate a PACS based radiology museum containing not only full DICOM datasets, but also relevant histological and clinical photographs. In conclusion, this technique offers a mechanism for generating anonymized catalogued radiology museums in PACS. Museum Builder represents a working prototype that demonstrates some of the archiving functions that are expected by teaching institutions from PACS.

The radiology museum is an integral component of every radiological training scheme. Over the past 5 years, the practice of radiology has moved from film to PACS, but the ability to build radiology museums has not kept pace [1]. For many of us, the hardcopy ACR collection provided hours of study and exam practice material. Those cases we saw in our formative years of radiological training often become ‘‘index’’ cases against which those that followed were measured. These museums, which are becoming increasingly obsolete [2], often comprised cupboard-like rooms filled with shelf upon shelf of ageing radiographs in various states of disorganization. Radiological museums have now diversified into multiple digital formats. DICOM files can easily be converted and saved in a number of manageable formats [3]. Large institutional collections can be acquired on CD-ROM [4, 5]. Personal teaching collections can be created in any number of readily available image databases [6–10]. Some of these databases are specifically designed for archiving radiological teaching cases and sometimes for storage on servers [4, 11, 12] for sharing access across networks or the World Wide Web [13]. Online database applications allow free text searches across thousands of cases, This projected was supported by a Kodak Radiology Fund Scholarship awarded by the Royal College of Radiologists in 2003.

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Received 22 June 2005 Revised 6 December 2005 Accepted 9 December 2005 DOI: 10.1259/bjr/11861604 ’ 2006 The British Institute of Radiology

sometimes in multiple institutions [13]. These archives are used for research, teaching [2, 9, 10, 14] and for assessment of radiological expertise [15]. With this digital diversity has come a subtle change in the look and feel of the radiology museum. Hardcopy museum cases must be read in the same way that hardcopy radiology is practiced; with a light box. Digital museum cases are read from personal computers and not in the PACS environment in which many of us now work. One of the obstacles to replicating the PACS environment on a PC is the prohibitive size of the DICOM files. A solution to these problems is not to replicate PACS in order to build a work-like radiology museum, but to build a radiology museum within PACS. To the best of our knowledge, the major PACS manufacturers provide only limited tools for archiving radiology teaching cases, whereas most radiologists consider this sort of functionality important or even essential when considering the purchase of PACS [16]. Some provide a system of academic folders that require a system administrator to set up. These provide inadequate archiving and retrieval mechanisms for generating usable databases within PACS [1]. Neither does there appear to be any third party solutions that meet these criteria. PACS manufacturers prohibit access to their databases other than by their employees and, therefore, a novel approach is required for third parties to generate teaching cases on PACS. The aim of this study was to see if it was feasible to The British Journal of Radiology, August 2006

Building anonymized radiology museums in PACS

develop a software application that could edit the DICOM headers to anonymize and catalogue teaching cases in order to build a radiology museum within PACS.

RAM running Microsoft Windows XP Professional, version 2002, with Service Pack 1). An academic licensed version of eFilm Workstation 1.9.4 [17] was installed as the helper application.

Materials and methods Results Principle Radiological studies stored on PACS can be identified and retrieved using a number of search fields common to all PACS – namely, the patient’s surname, forename, middle names and unique hospital number. This information is stored within a header in every DICOM file (Figure 1). This usually means that every image, in every series, in every radiological study contains this information embedded within it. Multiframe images, generated with ultrasound, contain the same information in a single header. Both of these types of DICOM file can be handled in the same way. After DICOM data has been generated by a radiological investigation, it is pushed to the PACS server to be archived. As PACS receives the DICOM data, it reads the DICOM header and stores the patient specific data in a database. When PACS is queried to search for a particular patient, it is this database that is searched and not the DICOM archive itself. This entry in the database, however, points to the DICOM dataset within the archive, which can then be retrieved and opened. Our application, called Museum Builder, exploits this process by replacing the patient specific data within the header of DICOM files that have been exported from PACS. When the DICOM files are returned to PACS, the new information within the header is added to the PACS database as a ‘‘new’’ patient. The DICOM files for the museum case are effectively duplicated, but with a new DICOM header. In effect, PACS sees Museum Builder as any other radiological modality contributing to the local PACS archive.

Hardware and software Our institution is a film-free hospital with a GE Centricity PACS (General Electric, Milwaukee, WI). Images can be viewed from a mixture of dedicated reporting workstations and PC-based web-browsers, which cover the whole hospital on a network with a 2 Gb s21 backbone and a 100 Mb s21 link to workstations. Museum Builder was developed using .NET technology and C# as the preferred language (Microsoft .Net Framework to run and Visual Studio to compile the C# source code). Museum Builder was installed on a networked PC (Pentium 4 CPU 2.80 Ghz with 1.0 Gb of

Figure 1. A screenshot of a readout from a hexadecimal editor of the header from a DICOM file demonstrating the ASCII values of the binary data. The patient specific binary data can be identified and edited with Museum Builder.


Process Museum Builder can work with most PC based DICOM browsers, which for the purpose of this article will be referred to as the helper application. After opening Museum Builder, the helper application’s database can be browsed or searched using search fields for the patient’s name or hospital number. Cases are then identified for museum archiving and highlighted. Once selected, the patient’s surname, forename, hospital number and date of birth are displayed in a row of text fields. Below this a second row contains the text fields for the anonymized museum case. The patient’s ‘‘new’’ surname and forename are selected from pre-defined menus. Optional middle names can also be entered. The new hospital number comprises a three letter code, again selected from a menu, followed by a unique four digit number generated by Museum Builder (Figure 2). The drop-down menu selections can be over-ridden by entering text into the fields manually. The patient’s date of birth is automatically anonymized to 1 January of the year of their birth, thus preserving age-related information in the DICOM header. Once the minimum number of fields has been completed (surname, forename and hospital number), the case is ready for anonymization (Figure 2). Clicking the ‘‘anonymize’’ button will then replace the patient specific data in each header of each DICOM file within each directory of the chosen radiological study or studies and the helper application database is modified accordingly. All other identifiable labels, private or otherwise, are changed or erased. This one click anonymizes every radiological study in the eFilm database for that patient; there is no limit to the number of types of examination, e.g. conventional radiography, CT, ultrasound or MRI, that can be processed at once. Even for large datasets with multiple examinations this only takes a few seconds on the workstation described. The process is simple and is currently performed in our department by clerical staff. Every DICOM file generated by a medical imaging device contains a unique identifier (the study SOP instance UID). At this stage, Museum Builder assigns each new case with a new unique identifier, generated internally from a sub-delegate range offered by Medical Connections [18], so that PACS does not recognize it when it returns. PACS just sees another new case arriving from a medical imaging device. When coding and anonymization is complete, Museum Builder composes an HL7 Radiology order message, which is sent to the PACS HIS/RIS broker. On receipt of this acknowledgment the new museum case is pushed from the helper application back to PACS (Figure 3).

Catalogue headings The American College of Radiologists (ACR) has established a well-recognized classification system for 667

A P Toms, B Kasmai, S Williams and P Wilson

Figure 2. A screenshot from Museum Builder’s anonymization window. The patient specific fields are replaced by catalogue headings that are selected from drop down menus.

radiology [19]. This system allows high-level discrimination of radiological diagnoses, and therefore accurate retrieval of data, particularly for research. However, the system is not entirely intuitive, and therefore a different classification system has been used in Museum Builder. The catalogue headings are almost universal within radiology. The patient’s surname is changed to a radiological anatomical unit, which broadly defines the sections of the human body that radiological investigations cover. These consist of head, neck, chest, abdomen, pelvis, extremities and breast. The patient’s forename is selected from a surgical sieve consisting of normal, developmental, trauma, infection, neoplasia, inflammation, vascular, metabolic. Thereafter there is an option to add one, two or more middle names from a selection of organ specific titles such as liver, lung, brain, adrenal and so on. The hospital number is replaced by a unique museum number that comprises a three-letter code followed by a four-digit number. The code reflects subspecialty interests within radiology and include MSK for musculoskeletal, GIT for gastrointestinal and H&N for head and neck (Figure 2).

Non-radiological images Many DICOM browsers will import non-DICOM image files such as JPEG and TIFF files. During the 668

import process, DICOM header fields are entered manually and new DICOM files are generated. This allows non-radiological images to be added to the

Figure 3. A schematic diagram demonstrating data flow between Museum Builder, the helper application, PACS and RIS/HIS.



Figure 4. A montage of screenshots demonstrating material from a single museum case including computed radiography, MRI, arthroscopy and histology.

museum case, including histology, arthroscopy, endoscopy and clinical photographs (Figure 4). The report field of these ‘‘new studies’’ can then contain pathology reports, operative notes and clinical findings.

Reports Reports for the museum cases can be added to the PACS RIS. These can be added by pasting text into a report window in Museum Builder (Figure 5). These reports can include copies of the original radiological report or can be entirely new and contain clinical histories, updated radiological reports, results of other special investigations, differential diagnoses and discussion.

Discussion Museum Builder is a fully working prototype that has some of the functionality required by teaching institutions to develop radiology museums within PACS. It uses a novel approach to generating anonymized searchable teaching cases without accessing the PACS database directly. Teaching cases can then be read in the PACS environment in which the radiology trainee and his or her trainers work. Instead of being presented with The British Journal of Radiology, August 2006

single or selected digital images from a teaching archive or CD-ROM, the trainee has access to the whole DICOM dataset. When reading cross-sectional investigations, the trainee would have to interrogate the entire dataset, including localizers and sequences repeated because of technical problems, to gather the signs necessary to yield a diagnosis. It is this process that cannot be replicated by non-DICOM museums and teaching collections. There are limitations to radiological museums created by Museum Builder. There is no free-text search function within the PACS browser window, which would allow the user to search for a specific diagnosis. However, the objective was to create a radiological museum that functioned in a similar manner to the ACR hard copy museum and therefore did not require the ability to immediately recall specific cases. Trainees can search through the database through catalogue headings based on anatomical site, disease process and radiology subspecialty. Museum Builder can be used with any PACS and, in theory, integrated with any PACS broker, but this has only been tested with PACS in our institution. Whilst it should work with any PACS broker, the concept of Museum Builder does not allow for a ‘‘plug and play’’ solution. Museum Builder needs to be configured for each PACS broker in the same way that any CT or ultrasound machine must be configured to work with a 669

A P Toms, B Kasmai, S Williams and P Wilson

Figure 5. A screenshot demonstrating the ‘‘Report Window’’ where the text of the museum case has been pasted in prior to sending to RIS in an HL7 script.

particular PACS. However, configuring Museum Builder has been simplified by using variables in the HL7 code that can be defined from within the Access database according to the local PACS broker profiles. The concept of Museum Builder is relatively simple and the coding is mainstream. It currently works as a bolt-on application to PC-based DICOM browsers, but there are a number of options for further development. Museum Builder could be coded to work as a DICOM client and, therefore, could stand alone in its integration with PACS and RIS. However, it does not make sense to repeat the work done by many affordable or free, readily available proprietary DICOM browsers. It would be easier to add Museum Builder’s functionality to these 670

applications. The most elegant solution would be for the PACS manufacturers to add this functionality to their current systems. That way the radiologist could build his or her DICOM radiology museum without leaving PACS. This sort of functionality within PACS has certain implications for governance of the educational material because other allied healthcare workers outside radiology, and IT personnel, also have access to the database. In our institution we developed a governance protocol that was approved by the Caldicott Guardian to ensure that the limitations of patients’ consent to procedures and investigations were adhered to. Rather than being a risk, the radiology museum is considered a valuable The British Journal of Radiology, August 2006


hospital-wide resource. Generating validated case material is always time consuming, and therefore the number of museum cases has, so far, made a negligible impact on the PACS archive capacity. In theory, however, duplicating large volumes of archive material could have serious implications for storage and therefore needs to be carefully controlled. In our institution, all museum material must be approved by a Radiology Museum Committee, which acts as a gatekeeper safeguarding the quality of the PACS museum and controlling its impact on the clinical archive.

Conclusion Museum Builder demonstrates that it is feasible to build anonymized catalogued radiology museums within PACS, by editing the patient specific headers within the DICOM files, and therefore without directly accessing the PACS database. Teaching cases generated with this tool allow the trainee to read the full DICOM datasets within the normal PACS working environment. By using PACS as the radiology museum repository, the problems of storing and transmitting large image files and directories can be overcome. Museum Builder provides a model for the some of functionality that many academic institutions would like to see added to PACS by the PACS manufacturers.

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Optimizing localization accuracy in head and neck, and brain radiotherapy 1

M McJURY, PhD, 2K DYKER, MBChB, MRCP, FRCR, 3R NAKIELNY, MA(Cantab), BM BCh, FRCR, 1J CONWAY, PhD and 2M H ROBINSON, MD, FRCP, FRCR Departments of 1Radiotherapy Physics and 2YCR Clinical Oncology, Weston Park Hospital, Whitham Road, Sheffield and 3Department of Radiology, Royal Hallamshire Hospital, Glossop Road, Sheffield, UK ABSTRACT. The purpose of this study was to investigate the impact on localization of utilizing contrast-enhanced CT scans and the formal input of a radiologist in the planning process. 25 head and neck/brain patients had pre- and post-contrast CT scans in the treatment position. Radiotherapy treatment was planned on the unenhanced CT images as per standard practice. Retrospectively, their scans (unenhanced and enhanced) were re-contoured by two oncologists and a radiologist. These new contours were compared with the original unenhanced treatment contours and differences in contour volume, geographical isocentre position and tolerance coverage of the associated planning target volumes (PTVs) were evaluated using the original plans. The use of contrast enhanced CT data during localization by the oncologist shows little change in gross tumour volumes (GTVs) or PTVs, geographical position or tolerance coverage for the targets in the brain studied here. Larger changes in mean volume are seen for the head and neck cases alone. Changes are greater and statistically significant (p,0.05, Wilcoxon signed rank test) for localization by the radiologist. Furthermore, when comparing the original PTV marked by the oncologist with a new PTV recontoured by the oncologist, but based on a GTV marked-up by the radiologist, again statistically significant (p,0.01) changes in percentage volume are noted. Intraoperator precision is good, percentage volume differences being of the order 3–6%. PTVs also show improved standard deviations compared with GTVs. Geographic shifts are generally within our departmental tolerance levels for daily patient setup. Comparing precision of unenhanced data with enhanced, mean percentage volume changes are smaller, but not statistically significant. The use of enhanced scan data for localization has little effect on size, geographical position or tolerance coverage of PTVs marked up by the oncologists in this study. However, more important is the input from a radiologist. Statistically significant differences due to mark-up on enhanced scans by the radiologist are shown. Furthermore, significant differences are also seen between PTVs based on oncologist-generated GTVs, and those based on radiologist-generated GTVs. The goal of modern three-dimensional (3D) conformal radiotherapy is to accurately conform dose to the tumour target whilst minimizing dose to nearby normal tissue. For patients with brain and head and neck tumours, the targets are in close proximity to many critical structures, making tissue avoidance a high priority. Failure of locoregional tumour control is also a particular issue for these patients as brain tumours do not tend to metastasise and nodal involvement in head and neck tumours is one of the major prognostic factors. These factors determine the quality and length of life in many patients undergoing radiotherapy, and are known to be due to biological and technical factors [1]. The interCurrent address for Dr M McJury: Department of Medical Physics, The Northern Ireland Cancer Centre, Belfast City Hospital, Lisburn Road, Belfast BT9 7AB, UK. The authors gratefully acknowledge support from Weston Park Research Fund (MM) and Yorkshire Cancer Research (KD, MHR).

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Received 15 December 2005 Accepted 7 February 2006 DOI: 10.1259/bjr/14663755 ’ 2006 The British Institute of Radiology

clinician variability is known to be high across all sites [2–4]. Many methods to improve the consistency of tumour and organ delineation have been tried, including the development of volume delineation protocols [5, 6], the application of additional imaging modalities [7–9], the use of contrast-enhanced data [10] and publication of anatomical maps [11]. The first step in the treatment process is localization of the tumour, usually using radiographic films and/or CT data. Tepper et al [12] showed that performing a planning CT scan in addition to existing diagnostic information enabled improvements in target localization in 49% of patients. In the authors’ oncology centre, intravenous (IV) contrast is used routinely in diagnostic CT scanning of these patients, but not used routinely when acquiring CT scans for treatment planning purposes. The use of enhanced CT scans can offer improved tumour visibility in many cases and may enable improved localization for planning [13, 14]. The British Journal of Radiology, August 2006

Optimizing localization in radiotherapy

Whilst seemingly obvious, improvements may be available in marking-up visible gross tumour volumes (GTVs), although the overall impact on the target or planning target volume (PTV) and general plan quality is unknown. The use of contrast-enhanced data has been shown (for other sites, e.g. Zhou et al and Valcenti et al [10, 15]) to provide improved tumour delineation. For brain tumours and head and neck cancer, there have been no previous reports that assess the impact of contrast-enhanced CT on the delineation of the GTV and the effect of such change on the PTV. When contouring, the oncologist will mark the gross tumour volume (GTV), which is the visible extent of the tumour, with the help of diagnostic MRI images. A margin is added to this to allow for non-visible tumour infiltration, creating a clinical target volume (CTV). An additional margin is added to the CTV to account for patient movement and set-up inaccuracies, generating a final planning target volume (PTV) to be treated [16]. For this study, PTVs are not marked on directly, but are always generated by adding a uniform 2D margin to the initial GTV. Once generated, the initial PTV may be edited by the clinician to achieve a final PTV contour. This editing may be necessary, for example, if the software-generated PTV extends beyond the patients’ anatomy. Although a radiologist is the recognized expert in the interpretation of medical images, in many centres (the authors’ included) definition of the GTV is performed solely by the oncologist. By requiring the GTV to be defined by a radiologist and the remainder of the marking-up process (definition of the CTV and PTV) to be done by the oncologist, improvements in planning accuracy and outcome may be possible. This study addresses two main questions: (i)

Does the use of IV contrast during the acquisition of CT data for treatment planning significantly alter tumour localization and plan quality?

and (ii) Does the input of a radiologist in delineating the tumour GTV significantly alter the volume or position of the primary, or quality of the final treatment plan?

Methods A group of 13 head and neck and 12 brain cases were included in the investigation. Table 1 gives a list of sites and associated patient numbers. Patient selection criteria: all patients having radical external beam treatment with CT planning for head and neck or brain sites are considered. Each patient must be physically able to Table 1. Breakdown of the individual sites for the study patient group Primary site

Number

Brain – surgical resection Brain – no surgical resection Larynx Nasopharynx Tongue Tonsil

5 7 2 4 1 6


undergo CT scanning (weight and girth can result in some exclusions) and be able to consent. Ethics approval was obtained for non-standard administration of contrast as part of the treatment planning process. They each had pre- and post-contrast CT scans carried out in the treatment position. Our standard CT protocols were used for image acquisition: 5 mm slice thickness with 5 mm interslice gap, acquired in helical mode with pitch of 1, on our Picker PQS CT. After administering contrast, the couch was returned to the original position without any patient movement, enabling both scans (unenhanced and enhanced) to be acquired with identical scanner coordinates. GTV and PTV contours were marked-up using the AcQSIMTM virtual simulator (Philips Medical Systems, Best, The Netherlands) on the unenhanced scan and the patients were treated based on this scan. After a gap of several weeks, each patient’s unenhanced and enhanced scans were then retrospectively re-contoured by (i) the original oncologist (A) [to generate data on intraobserver precision], (ii) a second oncologist (B) [to generate oncologist interobserver precision], and (iii) a radiologist. In this way, several sets of GTVs and PTVs were generated for each patient. At each contouring session, all previous contours were removed from the image display, so the clinician was blinded to all previous work. Diagnostic films/images and patient notes were made available to the clinician marking-up at each contouring session. The enhanced studies were intrinsically registered with the unenhanced, such that contours marked on the enhanced image would be automatically transferred and stored with all previous others already marked on the unenhanced study. For data sets with identical scanner coordinates, image registration was performed automatically by the AcQSIMTM software (see Figure 1). The accuracy of this registration is dependent on negligible patient movement during scanning. Diagnostic scan data was not registered to planning CT data, but was generally available as hardcopy films and viewed on a light-box beside the AcQSIM work-station. All contours were stored on the unenhanced data set and analysis performed on this data set. The data for each patient allow us to make a number of comparisons of localization and planning: 1. To investigate the difference between marking-up by the oncologist and radiologist: a comparison of the original unenhanced and contrast enhanced scan contours and the re-contour; 2. To investigate the influence of contrast: a comparison of unenhanced and enhanced contours marked-up by (i) oncologist (A) and (ii) the radiologist; changes in contour data (volume, tolerance coverage and target isocentre displacement) are determined; 3. To investigate the influence of the radiologist: a comparison of GTV contours marked-up by oncologist (A) and the radiologist; also using the radiologists initial GTV, a PTV was generated by the oncologist and tolerance coverage determined. In comparing contour pairs, three indices were used: Volume changes: contour pairs were analysed to identify any changes in the volume of the GTV or PTV 673

M McJury, K Dyker, R Nakielny et al

Figure 1. The image fusion workspace showing marked-up contours. If the CT coordinates are the same for both scans, enhanced CT data on the left can be automatically fused to the unenhanced data on the right. Contours marked-up on the enhanced scan are then automatically transferred to the unenhanced scan for storage with previous contours. Contours shown are planning target volume (PTVunenh) (dark line) and gross tumour volume (GTVunenh) (light line).

contours. All unenhanced patient scans and contours were imported into the CADPLANTM (Dosetek and Varian Medical Systems) treatment planning system (TPS) and dose–volume histograms (DVHs) were generated to yield values for GTV and PTV volume. Geographical changes: pairs of GTVs were compared to identify any geographical shift of the re-marked contours from the position of the original GTV. On the AcQSIM virtual simulator, shifts between the geometric centres of the GTV contours were measured in three orthogonal axes, defined as lateral (L), anterior/posterior (A/P) and superior/inferior (S/I) shift. Using the Isocentre Manager, the centre of gravity of each GTV was identified automatically. Shifts between centres of GTVs under comparison were then simply found by subtraction of the coordinates in the orthogonal axes. From these shifts, an overall 3D scalar value was computed for each contour pair. The displacement of the re-marked contours can be assessed in terms of comparison with conventional treatment set-up tolerance. That is to say, we can note when the change in geographical isocentre, due to the use of additional input (contrast-enhanced data or radiologist input) is of the same order as an alteration in patient geographical 674

set-up isocentre, which would conventionally require action by staff to reposition the patient. Area coverage changes: in the TPS, the original treatment plan was applied to all subsequent sets of contours. Pairs of PTVs (original and re-marked) were then compared in terms of tolerance volumes (TV), i.e. the percentage of the target which is either below 90% prescribed dose (target under-coverage) or above 105% prescribed dose (target over-coverage). If, for example, the enhanced target contour is assumed to be the ‘‘true’’ target, the amount of under- or over-coverage the ‘‘true’’ target will experience can be measured, the original treatment plan (based on the unenhanced target contour) having been applied.

Results Interclinician data Table 2 shows a summary of the data for changes in contour volumes. The unenhanced contours marked up by the oncologist and radiologist show good agreement, with percentage volume changes in the order of 1–6%. The British Journal of Radiology, August 2006

Optimizing localization in radiotherapy Table 2. Percentage volume changes for precision data

Table 4. Tolerance coverage changes

Data comparison

Volume % differences

GTV

PTV

Data comparison

Mean SD

Mean SD

Unenhanced – unenhanced

Intraoperator (Oncologist) Interoperator (Oncologist) Intraoperator (Radiologist) RadiologistOncologist Radiologist (GTV)+ Oncologist (PTV) Intraoperator – Oncologist

–4.1

62

1.2

24

–3.2

52

5.9

33

Enhanced – unenhanced

–1.34 30.8 Enhanced – unenhanced

39.1 131 2.53 1.7

43.7

7.4

GTV, gross tumour volume; PTV, planning target volume; SD, standard deviation.

The mean percentage changes between pairs of GTVs and PTVs correlate well. The PTVs show smaller standard deviations compared with the GTVs. When re-contouring the contrast-enhanced images, the oncologist intraoperator GTV precision shows a significant (p,0.01, Wilcoxon signed rank test, two tailed) improvement with mean of 1.7 and SD of 7.4 compared with mean 4.1 and SD 61.7. Isocentre shifts between pairs of contours are shown in Table 3. For the oncologist, all shifts are within our daily setup tolerance of 5 mm for mould-immobilized patients. Our standard protocol requires no corrective action by staff for errors in geographical isocentre set-up of this level. Isocentre shift difference between contours marked by radiologist and oncologist are also higher (though not significantly) than precision data from either operator, with value 5.8 mm, which is outside the limit for daily patient setup tolerance. When the original plans were applied to the remarked-up contours (PTVs), changes in target and normal tissue coverage were calculated. Under-coverage defined as less than 90% and over-coverage as greater than 105%. These are shown in Table 4. There is good agreement with plan coverage, with maximum percentage changes being of the order of 5–6% for oncologist mark-up precision. Many consider a radiologist’s delineation of a GTV as the gold standard, due to the radiologists’ specific training [25, 26]. In this study, the radiologist was Table 3. Isocentre shifts for gross tumour volume (GTV) precision data Data comparison

GTV isocentre shifts

Mean

SD


Intraoperator (Oncologist) Interoperator (Oncologist) Intraoperator (Radiologist) RadiologistOncologist Oncologist Radiologist

3.7

2.1

4.2

2.4

4.2

1.6

5.8

2.4

4.8 5.2

2.5 2.6

Enhanced – enhanced


SD, standard deviation.


Intraoperator (Oncologist) Interoperator (Oncologist) Contrast changes

Overcoverage

Undercoverage

Mean SD

Mean

0.2

3.2

5.6

5.9

0.8

4.8

6.0

12.2

1.2

2.5

5.9

9.8

SD

SD, standard deviation.

required to mark-up GTVs only and an oncologist generated a PTV based on this. Table 2 shows percentage differences in contour volumes for the radiologist markup precision and for the radiologist compared with the oncologist. Mark-up precision for the radiologist is good, and in fact slightly higher than that of the oncologist, both in mean volume change and SD on the mean. There is a marked difference in mean percentage GTV volume (unenhanced) between radiologist and oncologist, with the radiologist marking larger volumes. This is not statistically significant for the group as a whole, but when the data are split into head and neck, and brain groups. For data split into the two groups (Table 6), the precision of radiologist mark-up is lower in the brain cases, although this is not significant. Whilst differences in (unenhanced) GTV mark-up between oncologist and radiologist for the group as a whole are not significant, when split into two, mark-up by the differing clinicians is found to be significant at p50.01 (head and neck) and p,0.02 (brain). A similar trend significance is found for PTV mark-up (p,0.05). Head and neck cases, in fact show an improved precision for the radiologist compared with the oncologist, with a lower intraoperator SD. Considering data where the oncologist marks-up a PTV based on a GTV delineated by the radiologist, again, significance change in unenhanced PTV percentage volume change is only seen when the data is split into two groups, with head and neck cases showing a significant difference (p,0.02, Wilcoxon rank sum test).

Use of contrast Data for differences in volume between contrastenhanced and unenhanced scan contours is shown in Table 5. For the group as a whole, the use of contrast in CT scanning and subsequent mark-up by the oncologist, leads to little (non-significant) change in GTVs and PTVs, similar in value to the typical precision differences seen in Table 2. Neither of the percentage GTV volume changes (for oncologist or radiologist) was statistically significant. However, for the GTVs, a comparison of the impact of contrast for the radiologist and oncologist, is significant (p,0.01, Wilcoxon signed rank test). Obviously, contrast-enhanced data has a differing impact on tumour delineation depending on which clinician is using it. This result can also be compared with the difference between radiologist-oncologist 675

M McJury, K Dyker, R Nakielny et al Table 5. Percentage volume changes for data involving contrast-enhanced data or input from a radiologist Data comparison

Enhanced – unenhanced Enhanced – enhanced

Volume % differences

GTV

PTV

Mean SD

Mean

Oncologist –5.9 Radiologist 98.8 Radioologist GTV, Oncologist PTV

53 172

SD

2.30 46 –4.1

24.2

GTV, gross tumour volume; PTV, planning target volume; SD, standard deviation.

mark-up of unenhanced scans (see Table 2). The use of contrast seems to have a greater difference in the marking-up of the radiologist than that of the oncologist. Isocentre shifts, which result from use of enhanced data (Table 3), are similar to those of precision data. Whilst there may be changes in volume, the geographical position of the structure has not changed. In agreement with data on volume changes, differences in coverage for contrast-enhanced data also show little change to precision values (see Table 4). When comparing the difference between enhanced and non-enhanced PTV mark-up, there is a statistical difference (p,0.01) between these for the contours generated from the radiologist’s GTV compared with those generated solely by the oncologist, see Table 5. Percentage volume changes for the radiologist mark-up is significantly larger with contrast than without (p,0.01, Wilcoxon signed rank test), so contrast is obviously a much greater influence on mark-up for the radiologist than for the oncologist. When investigating the data split into two groups, the mean percentage difference in GTV for head and neck cases is larger, at mean –20.2% compared with 8.5% for the brain group (although not statistically significant). There is little difference in PTV changes between the two groups, and similarly in isocentre shifts and tolerance coverage (data not shown).

Discussion These results highlight some difficulties inherent in this study and in the marking-up process more generally. Other authors have reported similar variations in marking-up targets in radiotherapy [2, 3, 15, 17–24] due to several factors.

Imaging The first step in the process is the assimilation of information from the available radiological images and clinical data. Some authors have noted significant variation amongst clinicians at this first step [4, 19]. When first marking-up for treatment, the oncologist may have additional information, which he/she will not have when re-contouring months later. They may have seen and examined the patient recently or have discussed the case with the surgeon regarding tumour 676

extent and geographic infiltration. Excluding the effect of these factors is obviously difficult. Clinicians when marking-up will use input from other diagnostic scans which may not have been performed with the patient in the same position as the planning scan. The scans may be pre- or post-operative, and thus be anatomically different. There may also be inconsistencies in the timing and modalities of imaging available for different patients. There is obviously a role here for image registration methods. Image registration will obviously increase the accuracy of combining diagnostic data in the planning process. Furthermore, the application of more advanced methods, such as elastic rather than rigid-body registration methods, will improve accuracy, especially for patients with data acquired on different patient couches and patients in slightly different positions. It is also difficult, over the course of a long study such as this, to monitor and control the use of notes and files for additional information used in marking-up as they are in constant clinical demand and so may not have been available (this was not recorded).

Contouring When contouring, the clinician will therefore have to mentally translate visual information on certain planes into contours in a different 3D plane. This complex process leads to increased inaccuracies in the final contour volume and may contribute to the large standard deviations we see in this study. Other authors have noted potential errors introduced in this step in the overall process and report large variations in target volumes [2, 4]. The CT images assist in delineating only the gross tumour volume. Assessing microscopic tumour involvement is difficult, highly subjective, and sometimes controversial, as noted by some reports [23]. Subjectivity will also be introduced in margin growing and editing when going from CTV to PTV [4, 22]. Indeed, Yamamoto et al [2] report on reduced mark up precision for CTVs compared with GTVs highlighting this problem. Although they mark up a GTV in a similar fashion, different oncologists in an institution may have slightly different philosophies when growing and editing margins. There can even be differences in interpretation of the ICRU [16] criteria for marking between clinicians [22]. PTV contouring precision should be better than GTV precision (differences in marking-up at the upper and lower extent of the target will have a smaller impact on the larger PTV volume), but may not be due to large variations in the size of margin added. The difference could be in the order of 0.5–1.0 cm in 2D, making a large difference to volume. Our results show little difference between intraoperator and interoperator volume precision for oncologists. Volume differences are presented as a percentage change from the original contour. In the case of GTVs with small numbers of slices, adding or removing a small numbers of slices will have a much greater effect on the percentage change than for PTVs which have 2–3 times the number of slices. At the The British Journal of Radiology, August 2006

Optimizing localization in radiotherapy Table 6. Data for the whole group split into head and neck, and brain cases Data comparison

Head and neckBrain Mean


Intraoperator (Oncologist) Interoperator (Oncologist) Intraoperator (Radiologist) Radiologist – Oncologist

SD

Mean

9.2 69

–16.5

54

–25.3 48

17.0

48

5.4 20 –31.4 58

SD

49.7 136 191

178

extremes of the tumour volume (most superior and inferior positions) marking-up to include or exclude a slice can often be very subjective and will impact on precision results. This may explain the much improved standard deviations for the precision of PTV contour mark-up compared with GTV mark-up, which is in agreement with other authors [20].

Size and sites In this work, the data are recorded for a sample size of 25 patients, which will have an influence on the uncertainties and statistical confidence. Although not large in statistical terms, the sample size is certainly larger then several similar studies in the literature [3, 4, 15, 19, 23]. Volumes were generated from full 3D CT data using DVH algorithms on our TPS. Volume data will therefore be accurate and should be an improvement on comparative assessments used by other authors involving 2D assessment of maximal tumour extent on specific CT slices [10], limited calculations from selected slices from a volume set [3] or volumes manually calculated from hardcopy films [22]. The cases in this work consist of small numbers from several different sites as shown in Table 1. Analysis of the data as a single group could obscure any benefit that may exist for a particular site(s). In certain circumstances, therefore, the data have been further split into smaller groups for consideration. With small numbers for individual sites, it was considered statistically prudent to merely split the data into two groups, namely brain and head and neck. The difficulty in delineating the target will depend on the site in question. Some authors report a variation between different groups of clinicians (radiologists and oncologists) when dealing with more or less ‘‘difficult’’ cases [19]. In the case of some resections, the oncologist may simply be marking-up a post-surgical cavity as a GTV, although technically this should be a CTV. In others, they can mark the site of the original tumour. Many of our patients had only cavities remaining. For the brain patients, the variation in marking was already so large that any difference made by the contrast may be too small to be detected. Some individuals seemed to be marking up the tumour cavity only, and some were marking a larger volume, more like a true CTV. Consistency was lacking and it was not always clear if The British Journal of Radiology, August 2006

the marked volumes were supposed to include a margin for microscopic disease or not. Thus, we could have been comparing unlike volumes for some patients. For these patients, getting good agreement between oncologist and radiologist mark-up may prove difficult and, indeed, we do see poor agreement between these clinicians for brain cases in particular. Other authors also report significant differences in mark-up between oncologist and radiologist [4, 19]. Yamamoto et al [2] also note a variation in precision between pre-operative and post-operative cases. Without contrast, contouring is far more dependent on the diagnostic scans. This leads to further inaccuracies if tumour volume changes due to surgery or chemotherapy have occurred. The use of contrast may be very helpful for some sites, and significantly less so for others. Unfortunately, with low numbers and a large mix of sites, any large changes or improvements for a particular site may not be apparent when looking only at results for the entire group and only by running a much larger study may influences of this nature be more apparent. As an example, consider the three cases shown in Figures 2– 4. Figure 2 shows pre- and post-contrast images for a tonsil patient. The post-contrast scan offers little additional information about the GTV. The PTV was never likely to show any significant change as it includes nodal groups in the neck, and so is not solely dependent on ascertaining the exact extent of the GTV. The head and neck tumours in general were much easier to outline, however, even using the diagnostic scan while voluming, as vessels and nodes were more easily distinguishable from other soft tissues. The efficiency of contouring is likely to improve with the use of contrast. Figure 3 shows a set of pre- and post-contrast images for a nasopharynx patient. This is helpful in showing the extension into the cranial contents. This illustrates that direct tumour enhancement is sometimes useful in head and neck cases. Compare this with Figure 4, which shows images of a patient with a glioblastoma. Here, the periphery of the tumour, between tissue and oedema, is well visualized in the post-contrast image on the left, compared with the precontrast scan on the right. We observed that contrast was very helpful in identifying the GTV in patients who had had a biopsy or minimal debulking only. For patients who had maximal debulking of their brain tumours, there was little enhancement. There are several characteristics of the tumour itself which can influence the impact of contrast on imaging. Tumours in certain sites, e.g. oral cavity, are more likely to enhance with contrast due to increased vascularity compared with others, e.g. larynx. As mentioned above, surgery not only changes the anatomy, but also the vasculature and oedema can mimic tumour very well. Contrast should make contouring more consistent, and our results show that. Finally, it must also be accepted, that although an expert in interpretation of medical images, the radiologist is not as expert in radiotherapy treatment planning. In agreement with other reports [4, 19], the input of the radiologist has shown to lead to significant mark-up changes, both in volume and isocentre shift, a stronger influence than that of using contrast alone. In certain circumstances, the radiologist may mark a considerably 677

M McJury, K Dyker, R Nakielny et al

Figure 2. Pre- and post-contrast images of a tonsil patient. Enhanced scan is shown on the left. The gross tumour volume (GTV) based on the unenhanced scan is shown in red.

different contour to the oncologist, but from our data it is not possible to ascertain whether it is more or less accurate. In the study, all patients were treated using a plan based on the unenhanced scans marked-up by the oncologist. It is not possible, therefore, to compare the outcome of patients treated on plans generated with and without contrast and therefore it is not possible to say

clinically whether the use of contrast-enhanced scan data or radiologist input resulted in more accurate, and therefore, improved treatment. In essence, we cannot say which of the clinicians’ mark-ups is the ‘‘true’’ one, or most accurate. Others also note differences between radiologists and oncologists, with radiologists marking consistently smaller volumes [4, 19], which is the opposite of what

Figure 3. Pre- and post-contrast images of a nasopharynx patient. Enhanced scan is on the left. The gross tumour volume (GTV) based on unenhanced data is shown.

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Optimizing localization in radiotherapy

Figure 4. Pre- and post- contrast images of a glioblastoma patient. Enhanced scan on the left. The gross tumour volume (GTV) based on the unenhanced data is shown.

we see in this study. However, in this study we have only a single radiologist compared with other reports performed with a larger and perhaps more representative group of radiologists.

Conclusions Radiotherapy has become more accurately targeted over the last 10 years. This continues with the advent of intensity-modulated radiotherapy (IMRT). As treatment becomes more complex by requiring different dose levels to be given to different areas, depending on level of risk, our plans become more heterogeneous. This makes accurate contouring a cornerstone of these advancements, especially if we aim to dose escalate or alter fractionation schedules. It was reassuring to observe good intraclinician and interclinician precision for the GTVs and PTVs, with isocentre shifts within daily setup tolerance and plan coverage changes also acceptably small. Brain cases showed less precision than head and neck cases. The use of contrast markedly improved the intraoncologist precision. The impact of using contrast caused greater differences for the radiologist than the oncologist. In fact, this difference was greater than the variation between them. This was more marked for the head and neck cases, although it was non-significant. We found the radiologist marked significantly different volumes for both GTV and generated PTV. Also, the mean isocentre shifts for these contours were outside our daily setup tolerance. Contrast enhanced planning appears to offer benefit in planning head and neck patients and those brain tumour patients who have not had a maximal debulking surgical procedure and so still have macroscopic tumour remaining. The British Journal of Radiology, August 2006

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M McJury, K Dyker, R Nakielny et al 10. Zhou SM, Bental GC, Lee CG, Anscher MS. Differences in gross target volumes on contrast vs. non-contrast CT scans utilised for conformal radiation therapy treatment planning for prostate carcinoma. Int J Radiat Oncol Biol Phys 1998;42:73–8. 11. Chao KS, Wippold FJ, Ozyigit G, Tran BN, Dempsey JF. Determination and delineation of nodal target volumes for head and neck cancer based on patterns of failure in patients receiving definitive and postoperative IMRT. Int J Radiat Oncol Biol Phys 2002;53:1174–84. 12. Tepper JE, Padikal TN. The role of computed tomography in treatment planning, In: Bleehen NM, Glastein E, Haybittle JL, editors. Radiation therapy planning. New York, NY: Marcel Dekker, 1983:139–58. 13. McJury M, Nakielny R, Levy D, Lilley J, Conway J, Robinson MH. Improving the localisation of radiotherapy treatments in head and neck and brain cancer: some initial findings. J Radiother Practice 2001;2:125–32. 14. Sharma R, Duclos M, Chuba PJ, Sharmsa F, Foreman JD. Enhancement of prostate tumour volume definition with intravesical contrast: a three-dimensional dosimetric evaluation. Int J Radiat Oncol Biol Phys 1997;38: 575–8. 15. Valcenti RK, Sweet JW, Hauck WW, et al. Variation of clinical target volume definition in three-dimensional conformal radiation therapy for prostate cancer. Int J Radiat Oncol Biol Phys 1999;44:931–5. 16. ICRU Report 50. Prescribing recording and reporting photon beam radiotherapy. Bethesda, MD: ICRU, 1993. 17. Cazzanigna LF, Marinoni MA, Bossi A, et al. Interphysician variability in defining the planning target volume in the irradiation of prostate and seminal vesicles. Radiother Oncol 1998;47:293–6. 18. Fiorino C, Reni M, Bolognesi A, Cattanero GM, Calandrino R. Intra- and inter-observer variability in contouring

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SHORT COMMUNICATION

Evaluation of the larynx for tumour recurrence by diffusionweighted MRI after radiotherapy: initial experience in four cases 1

V VANDECAVEYE, MD, 1F DE KEYZER, MSc, 2V VANDER POORTEN, MD, PhD, 3K DERAEDT, H ALAERTS, MD, 4W LANDUYT, PhD, 5S NUYTS, MD, PhD and 1R HERMANS, MD, PhD

MD,

3

Departments of 1Radiology, 2Otorhinolaryngology, Head and Neck Surgery, 3Pathology and 5 Radiation Oncology, University Hospitals Leuven, Leuven and 4Department of Experimental Radiobiology/LEO, Katholieke Universiteit Leuven, Belgium ABSTRACT. Radiotherapy-induced changes in the soft tissues of the neck hamper the early detection of persistent or recurrent tumour by clinical examination and imaging procedures. Diffusion-weighted (DW) MRI is a non-invasive technique capable of probing tissue properties by measuring the movement of water. The purpose of the ongoing study is to examine the usefulness of DW-MRI for differentiation of persistent or recurrent tumour from post-radiotherapeutic sequelae or complications. Four patients, suspected of tumour recurrence after radiotherapy for laryngeal squamous cell carcinoma, were examined using a DW-MRI sequence on a clinical 1.5 T MR system prior to surgery. In two patients, the DW-MRI images showed an asymmetric hyperintense lesion on b1000 images with low apparent diffusion coefficient (ADC)value, compatible with tumour on histopathology. All surrounding tissue presented high ADC values and absent signal on the b1000 images, histopathologically correlating to post-radiotherapeutic changes. The images of the third and fourth patient showed absent or minimal symmetric hyperintensity of the laryngeal soft tissues on the b1000 images and high ADC-values. In these cases, the histopathological diagnosis of radionecrosis was made and no tumour was found. In all four cases, differentiation of tumoral tissue from radiotherapy-induced tissue alterations was possible with DW-MRI.

The clinical and radiological detection of persistent or recurrent head and neck cancer is difficult in the early phase after radiotherapy (RT) [1]. Also, the differentiation between tumour recurrence and laryngeal necrosis can be challenging after radiotherapy. Diffusion-weighted MRI (DW-MRI) is a non-invasive technique capable of probing the micro-environment of tissue by measuring water movement, and has not yet been reported for evaluation of head and neck lesions in a post-RT setting. Four patients are presented in whom DW-MRI was used to evaluate possible tumour recurrence in the larynx after RT for squamous cell cancer (SCC).

Imaging technique All examinations were approved by the local ethics committee. The MRI study was performed on a 1.5 T SONATA scanner (Siemens, Erlangen, Germany). A T2 weighted Address correspondence to: Dr Robert Hermans, Department of Radiology, Herestraat 49, 3000 Leuven, Belgium. This work was partly financially supported by the research grant ‘‘Prof. em. A. L. Baert, Siemens Medical Solutions’’.


Received 12 October 2005 Revised 9 February 2006 Accepted 20 February 2006 DOI: 10.1259/bjr/89661809 ’ 2006 The British Institute of Radiology

turbo spin-echo (TSE) sequence was performed in the transverse plane, with parameters: 48 slices, 4 mm slice thickness, 0.4 mm intersection gap, field of view (FOV) of 20 cm625 cm, matrix of 2916512, repetition time (TR)/echo time (TE)53080 ms/106 ms, 2 averages, an echo train length of 9 and a resulting pixel resolution of 0.7 mm60.5 mm64.0 mm. The total acquisition time was 5 min 42 s. Then, a T1 weighted TSE sequence in the transverse plane was acquired with the following parameters: 48 slices, 4 mm slice thickness, 0.4 mm intersection gap, FOV of 20 cm625 cm, matrix of 2506512, TR/TE5775 ms/8.3 ms, 3 averages, an echo train length of 19 and a resulting pixel resolution of 0.8 mm60.5 mm64.0 mm. The total acquisition time was 5 min 35 s. This sequence was performed before and after administration of 15 cm3 of gadolinium-BOPTA (Multihance; BRACCO, Milan, Italy). In one patient, a fat suppression pulse was added to the T1 weighted TSE after contrast-administration, increasing the sequence time to 6 min 59 s. Additional coronal or sagittal T1 sequences after contrast administration were used depending on tumour localization. Diffusion-weighted echo planar images (EPI) were acquired with 48 slices in the transverse plane, bandwidth of 1502 Hz/pixel, 4 mm slice thickness, 0.4 mm 681

V Vandecaveye, F de Keyzer, V Vander Poorten et al

intersection gap, FOV 20 cm625 cm, matrix of 1046128, TR/TE57100 ms/84 ms, 3 averages and a resulting pixel resolution of 2.0 mm62.0 mm64.0 mm. The images were acquired using six different b-values (b50, 50, 100, 500, 750 and 1000 s mm22). The total acquisition time for this sequence was 5 min 48 s. All diffusionsensitizing gradients were applied in three orthogonal directions and combined to create a 3-scan trace. An apparent diffusion coefficient (ADC) map was calculated automatically using the built-in manufacturer’s software. All sequences were acquired with identical geometry to allow correlation of the DW images with the TSE sequences. Image analysis was performed on an off-line workstation using dedicated software (BioMAP; Novartis Pharma AG, Basel, Switzerland). DW-MRI was analysed in a first step by visual inspection of DW images with b-value 1000 s mm22 (b1000 images) and ADC maps in correlation to the co-registered anatomical images. Afterwards, multiple regions of interest (ROIs) were placed by two authors in consensus (VV, 2 years of experience, and RH, 15 years of experience in head and neck radiology) over the larynx at the supraglottic, glottic and infraglottic levels and averaged, excluding the site of any suspect lesions. Separately, the suspect sites were delineated. For all ROIs the ADC values were calculated using all b-values. More specifically, ADC-calculation was acquired from the native DW-MR images. The ROIs were drawn on the DW-MR images for b50 s mm22 and were then copied to correct positions on all other images (b550, 100, 500, 750 and 1000 s mm22) automatically by the software allowing for correct determination of signal intensity per b-value and thus allowing for accurate ADC-calculation. Both the qualitative and quantitative analyses were correlated to the histological findings (localization of tumour on DW-MRI and histological specimen, and signal intensity versus histological tissue type). Routine imaging procedures in three patients included a CT study of the head and neck during intravenous injection of a contrast agent, using a multidectector scanner (Siemens Sensation 16). Collimation was 16 mm6 0.75 mm, feed/rotation 9.9 mm s21, tube voltage 120 kV and mAseff 250. The effective slice thickness was 1.5 mm, the reconstruction interval 0.75 mm; axial and coronal slices were reformatted with a thickness of 2 mm parallel to and perpendicular on the true vocal cords. Whole-body FDG-PET studies were performed in two patients on a CTI Siemens ECAT 931 (Siemens, Knoxville, TN) with an in-plane spatial resolution of 8 mm and a transverse FOV of 10.1 cm for each bed position. The emission scan was initiated 60 min after the intravenous injection of 6.5 MBq kg21 FDG (to a maximum of 555 MBq). The raw imaging data were reconstructed in a 1286128 matrix with the use of an iterative reconstruction algorithm.

Case presentations Patient 1 3 months after completion of RT for a T1N0 SCC of the right true vocal cord (25 daily fractions of 2.2 Gy 682

resulting in a total dose of 55 Gy), a 66-year old patient presented with progressive hoarseness. CT showed a contrast-enhancing nodular lesion in the right true and false vocal cord, suggestive of tumour recurrence (Figure 1). The MR TSE sequences confirmed the irregular, contrast-enhancing mass in the right hemilarynx. On the b1000 images the lesion appeared hyperintense and on the ADC map hypointense, with ADC value (in mm2 s21) of 0.8361023. The surrounding tissue at the supraglottic, glottic and infraglottic levels showed no clear signal on the b1000 images; this tissue was marked by diffuse hyperintensity on the ADC maps, showing an average ADC value (in mm2 s21) of 1.4361023 for all measured ROIs. A direct laryngoscopy under anaesthesia showed a thickened right true vocal cord with adjacent granulation tissue. Biopsies were taken and histological examination confirmed the presence of tumour. The patient underwent a total laryngectomy with unilateral neck dissection. Histopathology confirmed presence of SCC in the right vocal cord, characterized by densely packed cells with large cytoplasmatic content and irregular nuclei, multiple mitoses and intercellular bridging, anatomically corresponding to the suspect lesion on DW-MRI. The surrounding tissue showed post-radiotherapeutic changes, including variable amount of inflammation, fibrosis and limited necrosis.

Patient 2 6 months after completion of radiotherapy for a T2N0 SCC of the left true vocal cord (35 fractions of 2 Gy resulting in a total dose of 70 Gy), a 54-year-old patient showed on indirect laryngoscopy during routine followup an irregular appearance of the left vocal cord. CT showed an irregularly thickened, slightly enhancing left true vocal cord, compatible with tumour recurrence; the soft tissue infiltration extended into the false vocal cord, subglottic region, anterior and posterior commissure, and involvement of the cricoid cartilage was suspected (Figure 2). The PET scan showed moderate FDG-uptake at the level of the larynx; this finding was interpreted as tumour recurrence. On MRI, the TSE sequences confirmed the presence of a contrast-enhancing mass in the left hemilarynx. Despite the small dimensions, the b1000 image showed an asymmetric hyperintense lesion corresponding to the lesion seen on CT scan. On the ADC map, this lesion was hypointense, with an ADC value (in mm2 s21) of 0.9661023. The signal of the surrounding tissue was suppressed on the b1000 images and the ADC maps showed diffuse hyperintensity, with an average ADC value (in mm2 s21) of 1.4161023 for all measured ROIs. Direct laryngoscopy was performed at the time of the laryngectomy and showed the presence of a left sided laryngeal mass. No biopsies were taken and surgery was performed immediately. Histopathology confirmed SCC in the left vocal cord corresponding to the suspect lesion on DW-MRI (Figure 2). Typical neoplastic features were present, showing densely packed cells with large cytoplasmatic content and irregular nuclei, multiple The British Journal of Radiology, August 2006

Short communication: Evaluation of the larynx using DW-MRI after radiotherapy

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Figure 1. (a) Transverse CT image and (b) transverse contrast-enhanced T1 weighted MR image show an ulcerated lesion in the right false vocal cord (arrows), corresponding to (c) a hyperintense rim on the b1000 image (arrows) and (d) a hypointense rim on the apparent diffusion coefficient (ADC) map (arrows). (e) This rim corresponds histologically to hypercellular cancer tissue (T) in the lateral part of the false vocal cord; medially, a necrotic ulcer (NU) is present. (f) A more detailed view of the tumour shows multiple nests of densely packed squamous cell carcinoma (SCC) cells.

mitoses and intercellular bridging. The surrounding tissue showed post-radiotherapeutic changes, including variable amounts of inflammation, fibrosis and limited necrosis.

Patient 3 A 45-year old patient presented with progressive pain, dysphagia and dyspnoea, 4 months after completion of RT for a T2N0 SCC of the right true vocal cord (35 fractions of 2 Gy resulting in a total dose of 70 Gy). Clinical examination showed diffuse laryngeal oedema. Progressive dyspnoea required placement of a tracheostomy. CT showed diffuse and pronounced thickening of the laryngeal soft tissues, with obliteration of the laryngeal lumen. No focal soft tissue mass was discerned. The presence of some small gas bubbles in the crico-arytenoid joints was interpreted as indicating laryngeal necrosis (Figure 3). However, PET showed a moderately hypermetabolic focus in the larynx; this was reported as being suspect for tumour recurrence. Similar to the CT findings, the TSE sequences on MRI (Figure 3) showed diffuse soft tissue thickening in the larynx, without a focal lesion. The b1000 images showed no asymmetric hyperintense signal and the ADC map showed diffuse hyperintensity with an ADC The British Journal of Radiology, August 2006

value (in mm2 s21) of 1.8461023. A direct laryngoscopy was not performed. Because of a worsening situation, the patient underwent total laryngectomy. Histological examination of the entire larynx showed severe radiotherapy-induced changes with necrosis and inflammation, including purulent infiltration; no neoplastic tissue was found.

Patient 4 A 53-year-old patient presented with progressive pain and dysphagia 6 months after completion of chemoradiotherapy for a SCC of the left true vocal cord (35 fractions of 2 Gy resulting in a total dose of 70 Gy). T-stage before treatment was unknown. The TSE-sequences on MRI showed diffuse laryngeal soft tissue thickening, more pronounced in the left true vocal cord, and a soft tissue defect in the posterior part of the right true vocal cord, suggesting laryngeal necrosis. However, based on conventional MRI findings a tumoral lesion could not be excluded on the left glottic level. Histological examination on multiple laryngeal biopsies taken during panendoscopy suggested the presence of tumoral recurrence at the level of the true and false vocal cords. DW-MRI showed only slight symmetric 683


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Figure 2. (a) Transverse CT image and (b) transverse contrast-enhanced T1 weighted MR image show thickening and increased contrast enhancement of the left vocal cord (arrows), corresponding to a hypermetabolic spot on (c) coronal FDG-PET image (arrow). The lesion is detected on (d) the b1000 DW-MR image (arrows) as a hyperintensity rim showing low intensity on (e) the apparent diffusion coefficient (ADC) map (arrows). (f) Histological section shows corresponding tumoral infiltration (arrows).

hyperintensity on the b1000 images in the laryngeal soft tissues, no focal lesion with low ADC-value could be detected. The laryngeal soft tissues appeared hyperintense on the ADC map, with ADC value (in mm2 s21) of 1.8761023. The patient underwent total laryngectomy. Histological examination of the entire larynx showed severe radiotherapy-induced changes with ulceration, necrosis and inflammation but no neoplastic tissue was found.

Discussion DW-MRI is a non-invasive technique able to depict the extent of random movement of water protons in biological tissues; the addition of two opposed magnetic field gradients makes the signal intensity dependent on the mobility of water molecules [2]. The amount of signal loss over the range of b-values correlates with the mobility of protons and is quantified by means of the ADC. The ADC value is mainly influenced by the size of the extracellular extravascular space (EES) containing free moving protons, with additive contributions from bulk water movement (e.g. intravascular flow) in the low b-value images and minor contributions from intracellular diffusion and transmembraneous transport [3]. Thus, any tissue architectural change causing structural barriers or influencing 684

the proportion of the extracellular versus the intracellular compartment is expected to alter the ADC value. The anatomic heterogeneity of the head and neck region with numerous air–soft tissue interfaces makes DW-MRI in this region prone to susceptibility artefacts, ghosting and image distortion. However, recent technical developments make EPI-based (including DW-MRI) imaging feasible in this anatomically highly demanding region. Application of parallel imaging decreases the echo-train length, which in turn reduces off-resonance and blurring artefacts [4], while application of high bandwidth [5], thin slices and meticulous shimming reduces image distortion and chemical shift artefacts. The acquisition of images with a large range of b-values allows a more accurate calculation of ADC and improves the image quality of ADC-maps by reducing movement artefacts and noise propagation. Experimental and clinical data support the potential use of DW-MRI for in vivo characterization of tissue. Wang et al [5] showed a significantly smaller ADC for malignant lesions, including SCC, than for benign lesions in the head and neck. Furthermore, ADC measurements provide reliable information on remaining viable tumour tissue in the follow-up of human high-grade gliomas after RT [6]. For SCC in an animal model, Herneth et al show that DW-MRI differentiates viable from necrotic tumour tissue [7]. The British Journal of Radiology, August 2006


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Figure 3. (a) Transverse CT image and (b) transverse contrast-enhanced T1 weighted MR image show diffuse thickening and contrast-enhancement of the false vocal cords, without focal nodular mass. Small air bubble adjacent to the right arytenoid can be appreciated (a, arrow). (c) Coronal FDG-PET image shows tracer uptake at the laryngeal level. No asymmetric hyperintensity is revealed by (d) b1000 and (e) the ADC map shows diffuse hyperintensity of the soft tissues at the same level. Histological examination reveals necrosis, inflammation and purulent infiltration, without evidence for tumour recurrence. (f) Detailed histological image shows granulation tissue and inflammatory infiltrate.

Recent progress in non-surgical treatment of head and neck cancer, by combining multifractioned high-dose radiotherapy (RT) with radiosensitizing measures [8] or chemotherapy [9], allows us to obtain tumour control even in advanced laryngeal SCC. The diagnostic accuracy of currently used clinical and imaging follow-up procedures may be compromised when organ preservation is attempted in such advanced disease. Treatmentinduced tissue changes are anticipated to be more pronounced in such circumstances, menacing the early detection of persistent or recurrent tumour [10–12]. DWMRI may allow differentiation between neoplastic tissue and post-radiotherapy inflammatory or necrotic tissue as the differences in tissue microstructure are expected to create differences in proton mobility. In patients one and two, the appearance of recurrent tumoral tissue on DW-MRI is illustrated. The lesions were hyperintense on b1000 and hypointense on the ADC map, contrasting with the surrounding tissue. Histologically, the recurrent SCC showed densely grouped cells with large cytoplasmatic content and occasionally intercellular bridging. These tumoral characteristics are expected to restrict the movement of protons as the high cellular index and large cytoplasmatic content increase the cellular tissue fraction and reduce the EES [13]. The British Journal of Radiology, August 2006

The diffuse high ADC value and the absence of any focal restrictive signal on the b1000 images in the laryngeal soft tissues in the third and fourth patient correlated with diffuse laryngeal necrosis and absence of tumoral recurrence. The findings on DW-MRI were contradictory to the FDG-PET findings in the third patient. The increased laryngeal uptake of FDG in this patient was presumably caused by granulation tissue and metabolically active leukocytes [14]. Indeed, the results reported on the value of FDG-PET in the postradiotherapy evaluation of laryngeal cancer are variable [11, 12, 15]. The specificity of this technique is diminished by inflammatory tissue alterations present early after radiotherapy. In all four patients, the diffuse hyperintensity on the ADC maps in the normal soft tissues of the larynx and hypopharynx correlated with expected post-radiotherapeutic alterations, such as inflammation and interstitial oedema, promoting free movement of protons. In all three patients examined with CT, a correct diagnosis could be made. This technique has a high accuracy for diagnosing recurrent laryngeal cancer after radiotherapy, but false positive and false negative results may occur [16]. Differentiation of tumour recurrence from therapy-induced laryngeal necrosis based on 685


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Figure 4. (a) Transverse non-fat saturated T1 weighted MR image before and (b) fat-saturated image after administration of contrast agent at the level of the vocal cords showing right-sided soft tissue ulceration (b, arrowhead) and tissue swelling on the left side (b, arrows), not allowing exclusion of tumour recurrence. Both (c) b1000 image and (d) apparent diffusion coefficient (ADC) map do not show a restrictive focal lesion, supporting the diagnosis of laryngeal necrosis without tumour recurrence. (e) Five and (f) 10 times magnified histopathological sections show necrotic tissue (N), and stromal tissue with reactive changes (S); multiple neutrophils and giant cells are visible, suggesting profound inflammatory reaction. No tumoral tissue was found.

anatomical findings may be problematic, although some CT-findings allow the correct diagnosis to be made [17]. The accuracy of conventional MRI-techniques has shown to be similar to CT in the post-radiotherapeutic neck [11]. As illustrated in the fourth patient, the asymmetric soft tissue thickening, visible on conventional MRI, did not allow exclusion of the presence of tumour recurrence. However, DW-MRI showed no restrictive lesion in the laryngeal soft tissues, supporting the diagnosis of postradiotherapeutic alterations or complications rather than tumour recurrence. This specific ability of DW-MRI to probe tissue microstructure is an interesting complement to the currently used imaging procedures in the evaluation of the post-radiotherapeutic neck.

Conclusion This report illustrates the ability of DW-MRI to characterize the tissue changes observed in the postradiotherapeutic larynx. The use of DW-MRI may have a complementary role in the non-invasive evaluation of the larynx after RT. Early, and preferentially noninvasive, differentiation of tumour recurrence from a treatment-induced complication is desirable, as in a 686

number of patients with the latter condition the laryngeal function may be saved by conservative measures. Further studies in a large patient population are ongoing to validate the reproducibility and diagnostic accuracy of the technique in the post-radiotherapeutic neck.

References 1. Parsons JT. The effect of radiation on normal tissues of the head and neck. In: Million RR, Cassisi NJ, editors. Management of head and neck cancer: a multidisciplinary approach. Philadelphia, PA: Lippincott, 1994:245–89. 2. Stejskal EO, Tanner J. Spin diffusion measurements: spinechoes in the presence of a time-dependent field gradient. J Chem Phys 1965;42:288–92. 3. Le Bihan D, Breton E, Lallemand D, Aubin ML, Vignaud J, Laval-Jeantet M. Separation of diffusion and perfusion in intravoxel incoherent motion MR imaging. Radiology 1988;168:497–505. 4. Taouli B, Martin AJ, Qayyum A, Merriman RB, Vigneron D, Yeh BM, et al. Parallel imaging and diffusion tensor imaging for diffusion-weighted MRI of the liver: preliminary experience in healthy volunteers. AJR Am J Roentgenol 2004;183:677–80. 5. Wang J, Takashima S, Takayama F, Kawakami S, Saito A, Matsushita T, et al. Head and neck lesions: characterization



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with diffusion-weighted echo-planar MR imaging. Radiology 2001;220:621–30. Hein PA, Eskey CJ, Dunn JF, Hug EB. Diffusion-weighted imaging in the follow-up of treated high-grade gliomas: tumor recurrence versus radiation injury. AJNR Am J Neuroradiol 2004;25:201–9. Herneth AM, Guccione S, Bednarski M. Apparent diffusion coefficient: a quantitative parameter for in vivo tumor characterization. Eur J Radiol 2003;45:208–13. Kaanders JH, Pop LA, Marres HA, Bruaset I, van den Hoogen FJ, Merkx MA, et al. ARCON: experience in 215 patients with advanced head-and-neck cancer. Int J Radiat Oncol Biol Phys 2002;52:769–78. Browman GP, Hodson DI, Mackenzie RJ, Bestic N, Zuraw L, Cancer Care Ontario Practice Guideline Initiative Head and Neck Cancer Disease Site Group. Choosing a concomitant chemotherapy and radiotherapy regimen for squamous cell head and neck cancer: a systematic review of the published literature with subgroup analysis. Head Neck 2001;23:579–89. Viani L, Stell PM, Dalby JE. Recurrence after radiotherapy for glottic carcinoma. Cancer 1991;67:577–84. Briggs RJ, Gallimore AP, Phelps PD, Howard DJ. Laryngeal imaging by computerized tomography and magnetic resonance following radiation therapy: a need for caution. J Laryngol Otol 1993;107:565–8.


12. Terhaard CH, Bongers V, van Rijk PP, Hordijk GJ. F-18fluoro-deoxy-glucose positron-emission tomography scanning in detection of local recurrence after radiotherapy for laryngeal/pharyngeal cancer. Head Neck 2001;23: 933–41. 13. Lyng H, Haraldseth O, Rofstad EK. Measurement of cell density and necrotic fraction in human melanoma xenografts by diffusion-weighted magnetic resonance imaging. Magn Reson Med 2000;43:828–36. 14. Kubota R, Yamada S, Kubota K, Ishiwata K, Tamahashi N, Ido T. Intratumoral distribution of fluorine-18-fluorodeoxyglucose in vivo: high accumulation in macrophages and granulation tissues studied by microautoradiography. J Nucl Med 1992;33:1972–80. 15. McGuirt WF, Greven KM, Keyes JW Jr, Williams DW 3rd, Watson NE Jr, Geisinger KR, et al. Positron emission tomography in the evaluation of laryngeal carcinoma. Ann Otol Rhinol Laryngol 1995;104:274–8. 16. Hermans R, Pameijer FA, Mancuso AA, Parsons JT, Mendenhall WM. Laryngeal or hypopharyngeal squamous cell carcinoma: can follow-up CT after definitive radiotherapy be used to detect local failure earlier than clinical examination alone? Radiology 2000;214:683–7. 17. Hermans R, Pameijer FA, Mancuso AA, Parsons JT, Mendenhall WM. CT findings in chondroradionecrosis of the larynx. AJNR Am J Neuroradiol 1998;19:711–8.

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REVIEW ARTICLE

Non-invasive measurement of perfusion: a critical review of arterial spin labelling techniques 1

E T PETERSEN,

MSc,

1

I ZIMINE,

PhD,

1

Y-C L HO,

MSc

and

1,2

X GOLAY,

PhD

1

Department of Neuroradiology, National Neuroscience Institute and 2Singapore Bioimaging Consortium, A*STAR, Singapore

ABSTRACT. The non-invasive nature of arterial spin labelling (ASL) has opened a unique window into human brain function and perfusion physiology. High spatial and temporal resolution makes the technique very appealing not only for the diagnosis of vascular diseases, but also in basic neuroscience where the aim is to develop a more comprehensive picture of the physiological events accompanying neuronal activation. However, low signal-to-noise ratio and the complexity of flow quantification make ASL one of the more demanding disciplines within MRI. In this review, the theoretical background and main implementations of ASL are revisited. In particular, the perfusion quantification methods, including the problems and pitfalls involved, are thoroughly discussed in this article. Finally, a brief summary of applications is provided.

Perfusion, or the steady state nutritive delivery of blood to the tissue capillary bed, is vital for the homeostasis, and thereby survival, of an organ. Accurate perfusion measurement can provide important diagnostic information on pathological conditions, e.g. whether an ischaemic organ is viable or not. Among other methods, MRI has emerged as a powerful tool for assessing tissue perfusion and possesses strong diagnostic and prognostic capabilities, especially when combined with additional MRI modalities such as T1, T2 and diffusion-weighted images. Historically, radiology has principally offered morphological imaging techniques; however, these new MRI perfusion techniques will add routine physiological imaging to its portfolio. This information is not only valuable in clinical settings, but is of great importance for basic physiological science and drug development, where, for example, perfusion changes due to a pharmacological stimulus might be monitored over time. There exist two main MRI perfusion methods: bolus tracking after the injection of an exogenous endovascular tracer and arterial spin labelling (ASL), which uses magnetically labelled water protons as an endogenous tracer. The complete non-invasiveness of ASL makes it very suitable for perfusion studies of healthy volunteers and in patient groups requiring repetitive follow-ups. This is especially important in patients with particular conditions, such as kidney failure, or in paediatric populations where the use of radioactive tracers or exogenous contrast agents may be restricted. Recently, ASL has gained more clinical acceptance partly due to Address correspondence to: Xavier Golay, Singapore Bioimaging Consortium, Agency for Science, Technology and Research (A*STAR), Singapore, 138667. E-mail: [email protected]. edu.sg.

688

Received 15 December 2005 Accepted 29 March 2006 DOI: 10.1259/bjr/67705974 ’ 2006 The British Institute of Radiology

these advantages, but also as a result of the recently introduced technique for imaging the perfusion territory of individual blood vessels in the brain. In addition, the increased availability of high-field (> 3 T) clinical scanners has moved ASL from the research and development stage towards the clinics. The present article will focus on arterial spin labelling approaches. Different implementations and their advantages and disadvantages will be briefly reviewed. More attention will be given to the different ways of modelling the ASL data for perfusion quantification and the problems and pitfalls involved. Finally, research and clinical applications will be revisited.

Methodology Terminology The term ‘‘perfusion’’ refers to the process involved in the nutritive blood delivery to the tissue’s capillary bed. The physiology behind this blood delivery can be described by many different parameters, such as the blood flow itself, the volume of the blood vessels, the time it takes a particle, such as a red blood cell to traverse the vasculature. Finally, it can be described by the velocity of these particles. The results of 2D and 3D brain perfusion imaging techniques are commonly expressed as cerebral blood flow (CBF), which has a typical unit of [ml 100 ml21 min21]. As noted, this is a rate [s21] rather than a volume-flux [m3s21] measurement, which one would associate, for instance, with the flow from a pipe. This is, however, convenient, as the average flow rate is measured in a voxel of arbitrary volume. For historical reasons, CBF is often stated as [ml 100 g21min21], a The British Journal of Radiology, August 2006

Review article: Non-invasive measurement of perfusion

rather misleading unit because no information can be obtained about the tissue mass of the individual voxels and, in practice, a mean brain density of 1 g ml21 is assigned to all voxels. The cerebral blood volume (CBV) describes the fraction of a voxel that contains blood vessels and is therefore dimensionless and usually expressed as [%] or [ml 100 g21]. Finally, the mean transit time (MTT) is the average time it takes a particle to traverse the vasculature [s]. These are the main parameters in general use, although measures such as time to peak concentration of the tracer (TTP) and flow heterogeneity (FH) can also be found in the perfusion imaging literature. All of them provide information about perfusion and are often referred to as perfusion maps.

Quantitative perfusion In order to measure tissue perfusion, one needs to follow the course of blood flow through the organ, and for this there exist two methods, one based on freely diffusible and the other on intravascular tracers. As their names suggest, freely diffusible tracers can leave the intravascular space without restriction and be distributed throughout the entire tissue volume, whereas intravascular tracers remain in the vasculature, which constitutes only a fraction of the full volume. Early perfusion experiments carried out by the pioneers Kety and Schmidt [1] used inhalation of free diffusible nitrous oxide (N2O) as a tracer. This tracer distributes throughout the entire tissue volume, having a MTT in the order of minutes at normal physiological flow values. The relationship between the flow, distribution volume and the MTT is described by the central volume theorem, which states that the ratio between volume and flow is equal to the MTT. This rather long MTT allowed them to sample both arterial and venous blood, as arterial and venous tracer concentrations reached equilibrium after a few minutes. In this experiment, the time to reach this equilibrium is directly related to the cerebral blood flow using Fick’s principle: dCt ðtÞ ~CBFðca ðtÞ{cv ðtÞÞ dt

ð1Þ

where Ct(t), ca(t) and cv(t) are the concentrations of tracer in the tissue, arterial and venous blood, respectively. The theory behind their method provides the basis for arterial spin labelling, which uses magnetically labelled water protons as the freely diffusible tracer. However, when a freely diffusible tracer is used, there is no access to the blood volume and an intravascular tracer would be needed in order to obtain this volume information. In this case, the distribution volume is equal to the blood volume and using the terminology for the brain, then according to the central volume theorem: MTT 5 CBV/CBF. This principle is used in MRI bolus tracking experiments, where an intravascular tracer such as gadolinium-DTPA allows measurements of MTT, CBF and CBV. Here, the first pass of the bolus is monitored, an arterial input function (AIF) selected and based on the indicator-dilution theory of Meier and Zierler [2] and, a The British Journal of Radiology, August 2006

model-independent perfusion estimation can be obtained [3]. Time-domain impulse functions are employed to describe and compute the tissue perfusion CBF, which is calculated using the deconvolution of the tissue concentration curve Ct(t) by the measured arterial input function Ca(t): ðt

Ct ðtÞ~CBF:Ca ðtÞ6RðtÞ~CBF: Ca ðtÞRðt{tÞdt

ð2Þ

0

R(t) is the residue function describing the fraction of contrast remaining in the system after a given time, t. If, alternatively, a steady state experiment is used, i.e. letting the contrast distribute to the total body blood volume, one can measure CBV by comparing images from before and after contrast. For recent reviews on these methodologies, see Østergaard et al, Grandin, Grenier et al, and Barbier et al [4–7]. The theory behind the two methods described above provides the basis for all applied perfusion measurement techniques used today. They range from MRI to CT and nuclear medicine imaging, all capable of measuring perfusion and/or blood volume in different parts of the body, depending on the selected tracer and imaging technique. Perfusion imaging encompasses physiological mass transport and exchange mechanisms, where the system is required to be stationary, linear and time invariant in order to satisfy the underlying flow quantification theory. This means that no major physiological alterations are allowed during the acquisition scheme in order to obtain quantitative perfusion. This should be kept in mind when planning an experiment, especially when dealing with functional studies for instance, where the perfusion rate is manipulated over time.

Basic arterial spin labelling The overall goal of all existing ASL techniques is to produce a flow-sensitized image or ‘‘labelled’’ image and a ‘‘control’’ image in which the static tissue signals are identical, but where the magnetization of the inflowing blood differs. The subtraction control-label yields a signal difference DM that directly reflects local perfusion because the signal from stationary tissue is completely eliminated (Figure 1). The label is usually performed by inverting or saturating the water molecules of the blood supplying the imaged region. By adding a delay between labelling and image acquisition, called inversion delay (TI), the labelled blood spins are allowed to reach the capillaries where they exchange with tissue water and thereby give rise to the perfusion signal. The signal difference, which is only 0.5–1.5% of the full signal, depends on many parameters such as the flow, T1 of blood and tissue, as well as the time it takes blood to travel from the labelling to the imaging region. Multiple repetitions are needed for ensuring sufficient signal-tonoise, and a model of the perfusion signal is usually used in order to quantify the perfusion. There exist two main classes of ASL techniques: continuous ASL (CASL) and pulsed ASL (PASL). In 689

E T Petersen, I Zimine, Y-C L Ho and X Golay

Figure 1. Schematic description of a perfusion weighted image (DM) obtained by subtraction of the labelled images from the control images.

CASL, the supplying blood is continuously labelled below the imaging slab, until the tissue magnetization reaches a steady state (Figure 2a). The PASL approach labels a thick slab of arterial blood at a single instance in time, and the imaging is performed after a time long enough for that spatially labelled blood to reach the tissue and exchange at the region of interest (Figure 2b). Both methods need a control experiment in order to visualize and quantify the perfusion.

Continuous arterial spin labelling The original ASL method proposed by Williams et al [8, 9] in 1992 used a continuous flow-driven adiabatic inversion scheme, a method that was previously used for angiography [10]. This type of adiabatic inversion of the arterial magnetization is realized using a 2–4 s continuous radiofrequency (RF) pulse while applying a magnetic field gradient in the flow direction. The moving arterial spins will therefore experience a slow variation of the resonance frequency, which will result in their inversion, while static (tissue) spins will just be saturated. Typically, the inversion ‘‘slice’’ will be selected just proximal to the circle of Willis, near the medullospinal junction or at the level of the common carotid, and the spins in blood that flows through this plane will be inverted. The inversion efficiency a, which is important for further quantification, depends on factors like the mean velocity of the blood, angulations of the vessels to the plane and the selected RF amplitude and gradient strength. Typical labelling efficiency is in the range of 80– 95% [8, 11–19]. Among the confounding factors of these long lasting inversion pulses are magnetization transfer (MT) effects [20]. When using a single coil for labelling and imaging, the off-resonance labelling pulse (with respect to imaging slice) will act as a powerful MT pulse in a way similar to an MT-weighted technique. The resulting saturation effect of the macromolecular pool will result in a reduced signal of the free water pool from the tissue of interest [21]. This is a very important issue, as the perfusion weighted images are calculated by subtracting 690

Figure 2. (a) Continuous arterial spin labelling (ASL) multislice experiment, using double adiabatic inversion for the control experiment, where labels get inverted during the passage of the first plane and returned to equilibrium during the subsequent passage of the second plane. (b) The EPISTAR pulsed ASL sequence, which labels everything at once and uses two 180 ˚ + 180 ˚ 5 0 ˚ pulses for the control images.

a labelled from a control acquisition, and if this MT effect is present only during the labelling scheme, it will lead to overestimated perfusion. In the first implementation, these MT-effects were compensated for by applying a distal labelling during the control experiment. This produces identical saturation effects but, due to the applied gradients during labelling, this is unfortunately valid only for a single slice. For multislice acquisition, Alsop et al [22] proposed the use of two closely spaced inversion planes, also called double adiabatic inversion (DAI). In the control experiment, the magnetization gets inverted while traversing the first plane and returns theoretically to its original state during the passage through the second plane (the CASL experiment is shown in Figure 2a). Double inversion is achieved by applying a sinusoidal modulation of the RF waveform. Global control of the MT-effects is obtained by matching the RF power and the location of the planes. Figure 3 shows a perfusion map obtained using this method. Another method was also proposed, called simultaneously proximal and distal RF The British Journal of Radiology, August 2006


Figure 3. Full brain continuous arterial spin labelling (CASL) DM images acquired using double adiabatic inversion [22] in a healthy 27year-old female. These images were acquired on a 1.5 T scanner of the F.M. Kirby Research Center for Functional Brain Imaging at Kennedy Krieger Institute.

irradiation (SPDI) for multislice acquisition [23]. In this later scheme, the RF power on the control scan is distributed evenly on both sides of the acquisition volume. However, a big limitation of these approaches is the doubled RF deposition, resulting in higher specific absorption rates (SAR). This is particularly important at higher field strength, and must be carefully considered for human studies at 3 Tesla and above. Using two coils is another way of avoiding MT-effects and reducing the RF deposition [24–26]. In such methods, a small dedicated coil is used for labelling the carotid arteries and, due to the small physical extent of the applied RF field, no saturation occurs in the imaging region. Another advantage is that selective labelling of each carotid artery is possible, allowing independent mapping of the left- and right-internal carotid perfusion territories [27–29]. The main disadvantage of this approach is linked to the need for non-standard hardware, such as a separate transmit channel and a detuneable RF coil, which are not usually available on commercial clinical scanners. Another disadvantage of this approach is that the labelling takes place further away from the imaging slices, resulting in increased relaxation of the label before entering the imaging region. This approach has recently been applied for full brain coverage in humans at 3 Tesla without exceeding current SAR limits [30].

for the label experiment and two 180 ˚ + 180 ˚ pulses of half the power for the control experiment at the same proximal location (Figure 2b) [31, 32]. Figure 4 shows a multislice experiment using this technique. Shortly afterwards, an alternative to this asymmetric method of labelling was proposed by Kwong et al [33] and independently by Kim [34], who named it flow alternating inversion recovery (FAIR). Here, the label is applied using a non-selective inversion pulse, while the control employs a concomitant slice selective gradient pulse. The symmetric nature of this sequence automatically compensates for MT-effects. PASL allows inversion of the arterial spins closer to the image slices and the inversion efficiency a is improved; however, due to imperfect inversion profiles, a gap between the labelling region and the image slices of typically 1–2 cm is needed, depending on the type of RF pulse used. This increases transit time ta from the labelling slab to the imaging slices leading to decreased efficiency. In addition, T1 relaxation of all the inverted spins will also result in a theoretically lower signal difference. Nevertheless, the ease of implementation and reduced practical problems, as compared with CASL, have made PASL a popular choice for perfusion imaging. This is reflected in the wide range of sequences available today, the common ones being listed in Table 1. For further in depth explanation of these sequences, see recent ASL reviews [4, 35].

Pulsed arterial spin labelling In 1994, Edelman et al [31] proposed the first pulsed ASL scheme. Contrary to CASL, the labelling is performed once in a 10–15 cm slab proximal to the image slices. For the PASL sequences, MT-effects have to be considered as well, although these are much smaller compared with CASL. In this first version of the ‘‘EchoPlanar MR Imaging and Signal Targeting with Alternating Radio frequency’’ (EPISTAR) sequence, inversion was performed distal to the image slice during the control experiment to induce identical MT effects in both cases. Again, this truly compensates for a single slice only and, therefore, the sequence was modified for multislice acquisition using a single 180 ˚ adiabatic pulse The British Journal of Radiology, August 2006

ASL perfusion quantification Having acquired the data using either technique, the subtracted control-label images will be perfusion weighted (Figure 1). The relationship between the DM signal and the actual CBF depends mainly on proton density and T1 relaxation rates of tissue and inflowing blood, the and their respective differences. In addition, the label transit time from the inversion slab to the observed region in the images is also an important factor. Traditionally, quantitative CBF estimation is carried out using the tracer clearance theory originally proposed by Kety and Schmidt [1], which was first adapted to ASL experiments by Detre and Williams et al [8, 9]. In the 691


Figure 4. Full brain pulsed arterial spin labelling (PASL) using EPISTAR/ PULSAR [32]. Images are the average of 30 control-label pairs acquired in 3 min using a TR of 3 s and TI of 1.7 s.

original model, it is assumed that the labelled arterial blood water is a free diffusible tracer, implying that the exchange of blood water with tissue water happens instantaneously upon its arrival to the parenchyma. Therefore, this model corresponds to a single compartment tracer kinetic, described by a mono-exponential tissue response function. The modified Bloch equation, including the flow dependent exchange term, becomes: dMt ðtÞ Mt,0 {Mt ðtÞ Mt ðtÞ zCBF Ma ðtÞ{ ~ ð3Þ dt T1t l

where Mt, Mt,0 and Ma are the tissue-, equilibrium- and arterial-magnetizations, respectively, l is the blood–brain partition coefficient, and T1t is the longitudinal relaxation rate of the tissue. In the original quantification model, further assumptions about uniform plug flow and equal T1 relaxation of both tissue and arterial blood were made [8, 9]. Many derivatives and improved versions of this first solution exist for both CASL and PASL experiments. Calamante et al [36] took the difference in R1 relaxation rate for tissue and arterial blood into account as well as

Table 1. Common arterial spin labelling sequences Method

Sequence name

Reference

CASL (Asymmetric)

The original continuous arterial spin labelling DAI (Double Adiabatic Inversion) SPDI (Simultaneously Proximal and Distal RF Irradiation) Two coil methods EPISTAR (Echo-Planar MR Imaging and Signal Targeting with Alternating Radio frequency) PICORE (Proximal Inversion with Control for Off-Resonance Effects) TILT (Transfer Insensitive Labelling Technique) DIPLOMA (Double Inversion with Proximal Labelling of bOth tagged and control iMAges) STAR-HASTE (Signal Targeting with Alternating Radio frequency - HAlf-fourier Single shot Turbo spin-Echo) PULSAR (PULsed Star labelling of Arterial Regions) QUASAR (QUAntitative Star labelling of Arterial Regions) FAIR (Flow Alternating Inversion Recovery) UNFAIR (UNinverted Flow Alternating Inversion Recovery) FAIRER (Flow Alternating Inversion Recovery Extra Radiofrequency pulse) FAIRER (Flow Alternating Inversion Recovery Excluding Radiation damping) BASE (unprepared BAsis and SElective inversion)

[8, 9, 120] [22] [23] [24, 25] [31] [121] [44, 52] [122] [123]

(Symmetric) (Two coil) PASL (Asymmetric)

(Symmetric)

[32] [124] [33, 34, 125] [126] [127, 128] [129, 130] [52, 131]

CASL, continuous arterial spin labelling; PASL, pulsed arterial spin labelling.

692



the trailing edge td of the bolus. In the work of Kwong et al [33], the transit time ta was also considered. Buxton et al [37] proposed a general kinetic model where all the above mentioned parameters can be taken into account. Here, the magnetization difference between labelled and control measurements is described using the convolution integral in a way similar to Equation (2): ðt : : : DM~2 Ma,0 CBF cðtÞ:rðt{tÞ:mðt{tÞdt

ð4Þ

0

where M0,a is the equilibrium magnetization in a blood filled arterial voxel, c(t) is the delivery function or fractional arterial input function (AIF). The residue function r(t2t) describes the washout of labelled spins from a voxel, and m(t2t) includes the longitudinal magnetization relaxation effects. The possibility to choose a particular arterial input function and to consider a certain exchange mechanism (single or multicompartment) allows greater flexibility for data analysis of both CASL and PASL experiments. However, analytical solutions are only possible with simple assumptions, and the most widely used ‘‘Standard Model’’ can be summarized using the following terms:

cðtÞ~

8 0, > > > > > < a:e{t:R1a ,ðPASLÞ > a:e > > > > : 0,

r(t)~e{

{ta :R1a

,ðCASLÞ

tvta ta ƒtvtd ta ƒtvtd ð5Þ

t§td

: CBF t l

Figure 5. Diagram on an arterial spin labelling (ASL) time course. In this graph, an example of two voxels having the same flow (60 ml 100 g21 min21) and relaxation characteristics, but different arrival time, is presented. The measured signal at a single inversion time of 1.5 s would result in a 12% lower DM signal in a voxel with delayed arrival of 500 ms (open square) as compared with a voxel with an arrival time of 200 ms (closed square).

Different issues relating to the actual perfusion quantification, using a single or multiple inversion time points, and the possible errors introduced using the standard model vs more advanced methods, will be discussed next.

Quantification errors

:

m(t)~e{t R1t

Transit time Again, the assumptions are: uniform plug flow and fast exchange, which is equivalent to single compartment kinetics. Solving Equation (5) using Equation (4) in the case of a PASL experiment gives a stepwise defined equation: DM ðtÞ~ 8 0, > > < {2:a:M

a,0

:CBF

dR > > : {2:a:Ma,0 :CBF dR

{R1a :t

e

dR:ðtta Þ

1e

, R1a :td {dR:ðtta Þ : {R1app :ðttd Þ e 1{e , e

tvta ta ƒtvtd

ð6Þ

(6)

t§td

where, dR 5 R1a – R1app and R1app 5 R1t + CBF/l, also called the apparent tissue relaxation rate. A similar set of equations can be obtained for the CASL experiments, with the only difference being that c(t) in Equation (5) is constant for CASL, whereas it is subject to T1a decay for PASL, since the latter is not a steady-state method. As can be seen, various parameters like the transit time ta, blood–tissue partition coefficient l, Ma,0, R1a and R1t need to be estimated or measured in order to obtain quantitative CBF values. The difference between the many ASL sequences is mainly in the measurement (or not) of these parameters. The British Journal of Radiology, August 2006

The major source of error in the quantitative estimate of cerebral perfusion is the arterial transit time, ta, which even in healthy subjects differs across the brain, being longest in distal branches, especially in the regions between perfusion territories also known as border-zone areas. In most PASL approaches, information on perfusion is assessed at a single inversion time point (Figure 5), and therefore without information about the transit time. Quantification is then based on the second step in Equation (6), either where ta is assumed equal all over, or simply set to zero. The problem in doing this is that not only is the quantification of CBF impossible, but relative perfusion values between regions are not valid either. Sequences like QUIPSS II and Q2-TIPS [38, 39] were developed to render ASL more transit-time insensitive. Constant bolus duration is assured by saturating the part of the label that remains within the labelling slab at a time delay short enough to be able to intersect the trailing edge of the fastest blood. Appropriate timing of this saturation and the following image acquisition makes these sequences insensitive to the transit time. This is really an advantage in volunteers and patients without vascular diseases where the difference in transit time is small (,1.5 s) [40], as well as for quantitative functional studies where it is known 693


that the transit time changes are small (0.1–0.2 s) between baseline and activation conditions [41, 42]. However, these methods will fail in patients with atherosclerosis, where the transit time can be long in affected areas (.2.0 s) due to low perfusion velocity and, in some cases, extensive collateral perfusion. In these cases, the problem can be solved by acquiring images at multiple inversion times and therefore measuring the entire DM curve (Figure 6). The standard sequences, in which a single time point is acquired at a time (2–4 min per time point) are generally too lengthy to be suitable for clinical examinations [37, 41]. Gu¨nther et al [43] introduced an elegant solution to this problem using a Look-Locker-like readout to measure the ASL signal at multiple inversion times in a single scan (Figure 6). Because of multiple low flip angle readouts, the general model needs to be modified by substituting R1app in Equation (6) with R1app,eff 5 R1t + CBF/l - ln(cosw)/DTI, where w is the flip angle and DTI is the interval between the excitation pulses. A similar scheme was recently implemented using the transfer insensitive labelling technique (TILT) [44] by Hendrikse et al [45]. CASL is also sensitive to the transit time, but because of the steady state behaviour of this type of sequences, the effect is smaller than in PASL. Further improvements can be achieved using a pre-delay of typically 0.9–1.5 s between the continuous labelling and the readout, which renders CASL methods almost insensitive to transit time differences [46]. However, the duration of the delay should be chosen according to the subject’s condition; in a healthy volunteer, a delay of 1 s would be suitable,

while in patients with cerebrovascular diseases, longer delays are necessary [46].

Vascular artefacts Vascular artefacts, associated with the inflow of labelled arterial blood into the arteries, can introduce important errors in CBF quantification. Strictly speaking, for Equation (5) to be valid, none of the DM signal should originate from within the arterial vasculature. This assumption is often violated in voxels containing feeding vessels or traversing arteries, resulting in a substantial overestimation of perfusion values. Ye et al [47] proposed the use of bipolar crusher gradients to eliminate the signal from the large feeding arteries in CASL experiments. This was also adapted for PASL sequences. Another solution is to choose a sufficiently long inversion time, allowing the feeding vessels to empty before acquisition [40, 46–48].

Inversion pulse shape and efficiency In PASL sequences, where a spatially defined label is used, the shape of the inversion pulse is of great importance. Ideally, the profile should be truly rectangular, allowing zero spacing between the labelling slab and the imaging region; due to finite duration of the RF pulse, this is not realizable in practice. Imperfect profile can reduce labelling efficiency, but, more importantly, there can be contamination in the imaging region from the labelling slab. To avoid this, a gap between the inversion and image plane is often introduced, with increased transit times as a result. In order to minimize this gap, longer adiabatic pulses like the hyperbolic secant (HS) [49], frequency offset corrected inversion (FOCI) [50] or bandwidth-modulated adiabatic selective saturation and inversion (BASSI) [51] pulses are often used. Another approach, as implemented in the TILT sequence [44], is the self-refocusing concatenated 90 ˚ Shinar-Leroux pulse combination that maintains the profile efficiency of the 90 ˚ pulses [52]. For CASL labelling, where labelling is performed in general at a distance from the imaging slices, the ‘‘profile’’ is less of a concern. It is the fulfilment of the adiabatic condition ensuring proper inversion that can be problematic [9, 15, 16, 22, 53].

Signal to noise issues Figure 6. Dynamic perfusion characterization. In this timecourse diagram, the black triangles show an example of repeated acquisition at multiple inversion time points. Usually, a three-parameter fit model is applied [37] to estimate cerebral blood flow (CBF) when multiple inversion times arterial spin labelling (ASL) sequences are used, which reveal additional information about arrival time and bolus duration. The second time-course (white triangles) has been calculated with identical perfusion parameters, while taking into account a Look-Locker readout method [43, 45, 124] (flip angle 5 30 ˚). The advantage of this method is a higher signal to noise ratio (SNR) as per acquisition time.

694

At normal perfusion rates of 40–100 ml 100 g21 min21 the signal change DM is in the order 0.5–1.5% of the full signal. Therefore, an average from typically 30 to 40 pairs of subtracted control and labelled images are required to get the desired signal to noise ratio (SNR) in the perfusion-weighted maps. A total scan time of 3–4 min is needed in order to acquire these data, which makes the technique very sensitive to motion artefacts. Proper head fixation and collaboration of the subject are necessary in order to obtain good results. Fast imaging techniques like single shot EPI and spiral sequences are often used to reduce the scan time between successive control-label The British Journal of Radiology, August 2006


pairs as well as overall scan time. Also, prior saturation of the image plane reduces the sensitivity to motion. In case of limited motion, realignment of the images can improve quantification, but in general it is compromised because of low resolution, small number of slices, and small DM signal. Finally, the running state of the scanner is important. Having a perfusion signal DM at only a fraction of the full signal, a slight drop in the scanner performance can preclude acquisition of clinically relevant ASL images, even if no effects can be detected in the standard clinical images.

Fast vs intermediate water exchange The complex nature of the brain vasculature might not act like a single compartment to water based tracers. In particular, the blood–brain barrier seals the vasculature from the extravascular space, allowing a dynamic interface that protects the brain from toxic substances while allowing nutrients and other essential compounds to pass and thereby maintain its homeostasis. The ‘‘free diffusible’’ water is known to be limited to ‘‘free’’ diffusion through the lipid membrane of the endothelial cells as well as through dedicated water channels [54, 55], whereas diffusion in between adjacent endothelial cells is impossible due to tight junctional complexes [56]. The speed of the exchange in comparison with the 2– 5 s duration of the ‘‘control’’ or ‘‘label’’ experiment still remains controversial. At higher flow rates, the exchange appears restricted [54], as though the water channels have been saturated. In order to deal with these effects, a few groups have suggested more elaborate multicompartmental approaches [57–59]. In general, the conclusion is that these effects are negligible at normal human perfusion rates at 1.5 Tesla, whereas the effect might be more pronounced when going toward higher field strengths. However, the intrinsic low SNR of the ASL techniques combined with the problems related to fitting all these additional exchange parameters will keep these models within the animal experimental world for at least another few years.

Bolus dispersion Even when using vascular crushing and modelling assuming multicompartmental behaviour, it still leaves another possible error source, namely the dispersion of the labelled bolus. In a PASL experiment the label is close to being rectangular at the time of labelling and if the blood does not experience any resistance, this shape would persist all the way to the image slices. In reality, the blood friction with the vessel wall, resistance in bifurcations and the pulsatile behaviour of the flow make the flow profile look like something between a parabolic and ideal plug flow profile [60]. This results in dispersion of the bolus while it travels from the labelling slab to the imaging region, which will lead to underestimation of the perfusion. Various attempts to incorporate these effects in the modelling have been made, and Hrabe et al [61] recently provided two analytical solutions that might reflect the The British Journal of Radiology, August 2006

physiology better. While the dispersion might be fairly homogeneous in healthy volunteers, patients with vascular diseases can be expected to have completely different dispersion in affected areas, and the flow estimation using these advanced models would still be non-quantitative.

Blood–brain partition coefficient When a tracer consists of water or is dissolved in water, it is necessary to know the ratio of water in the tissue of interest and the feeding blood water in order to correct the distribution volume and thereby obtain the correct perfusion values. In the ASL technique, this is more related to the proton density, which tells us about how much signal we can expect from different tissues. This ratio is called the blood–brain partition coefficient l, and it was first defined by Kety [62] and corrected by Herscovitch et al [63] to the commonly used values today: whole brain l 5 0.9 ml g21, grey matter lg 5 0.98 ml g21 and white matter lw 5 0.82 ml g21. The relevance of this coefficient is controversial; first of all because the valid use of it requires instantaneous exchange of the tracer, which, as earlier mentioned, can be questionable, but also due to the fact that Roberts et al [64] demonstrated that this coefficient not only varies between different tissues, but also from region to region. Nevertheless, a whole brain value of 0.9 ml g21 is commonly used for the calculations. Notice the unit of ml g21, which means that the unit of CBF/l in Equation (5) becomes s21 when CBF is expressed as ml g21 s21.

Blood equilibrium magnetization The inflowing blood has a proton density different from the tissue. For absolute quantification, the equilibrium magnetization of the arterial blood Ma,0 is needed, i.e. the available longitudinal magnetization from a fully relaxed blood filled voxel. This could be obtained from a partial volume-free arterial voxel, but with a standard inplane resolution of 3–4 mm this is not possible in practice. Alternatively, this information can be taken from the much larger sagittal sinus. However, the T2* of the deoxygenated venous blood, especially at higher field strength, is shorter than that of arterial blood. This results in a underestimated Ma,0. A different approach is to measure the equilibrium magnetization in a grey or white matter region and estimate Ma,0 using the blood– brain partition coefficient. For single inversion time experiments, this is often done on a voxel by voxel basis using the control experiment as Mt,0-map after correction for TR and T1. Note that Ma,0 is a direct scaling factor of the perfusion (Equation (6)) and therefore an error in the product a?Ma,0 will directly change the calculated CBF value.

Functional studies and non-steady state perfusion Finally, a requirement for using the various models and formulae for perfusion quantification is that tissue perfusion is a stationary, linear and time invariant 695


system. This would most often be satisfied in standard perfusion scans where the subject is resting. However, ASL has gained more and more popularity in functional studies where perfusion is monitored during various stimulation paradigms. Here, a typical experiment would have a repetition time of 2–3 s, while 4–6 s would be needed in order to acquire both control and label. Now, in functional experiments, the haemodynamic response time is often in the same range, which results in a violation of the steady state requirement [65, 66]. Not only will the flow change in between the control and label experiment, but also in between labelling and acquisition. Two main problems arise from this; first, the models do not assume increasing or decreasing perfusion during the experiment and calculation errors will occur, and second, and probably the most important, is that T2* will change due to the BOLD effect, which will reduce the signal intensity changes when acquiring at typical echo times of 15–25 ms. In the case of a standard block paradigm, where the rising edges are discarded in the statistical analysis, this is not a problem. With basic research moving towards the characterization of the haemodynamic response on the other hand, the information in the rising and the falling edges are especially important. Recently, Lu et al [67] proposed a new subtraction method for reducing the T2* effect, which is very suitable for this type of experiments, however, without addressing the problem of non-steady-state regimen.

Comparison with other imaging modalities At this point, it should be kept in mind that related issues such as the ones highlighted in the previous paragraphs apply to other perfusion modalities like DSC-MRI, CT-perfusion and positron emission tomography (PET) as well. Dynamic susceptibility contrast using tracers such as gadolinium-DTPA for instance, suffers from the fact that the relationship between measured signal and contrast is non-linear and depends on parameters such as the field strength, shim of the magnet and the constitution of the vessels [68]. In addition, correct scaling of the AIF is influenced by partial volume effects in the voxels from where it is measured. Altogether, this makes quantification troublesome, and in the general clinical practice only a relative perfusion measure is possible. Similar problems exist in CTperfusion, mainly related to the extraction of the global AIF [69]. PET can be said to be a more ‘‘pure’’ method for measuring perfusion due to the use of free diffusible tracers and experimental durations that ensure a steady state. However, a relatively low resolution of typically 6– 10 mm introduces partial volume effects, i.e. a mixture of grey-matter, white-matter and CSF will be present in almost all voxels, making direct comparisons to the abovementioned methods questionable.

Applications Despite the problems related to the quantification of perfusion using ASL, this technique has been used in 696

numerous applications, ranging from basic neuroscience using animal models and human volunteers to clinical perfusion measurement in pathologies such as stroke and brain tumours. Furthermore, as long as the same procedure and parameters are consistently used, reproducible results can be achieved using ASL. Also the fact that most clinical decisions can be based on relative differences in perfusion rather than absolute measures could render this method useful in the daily clinical practice. Although most publications on ASL to date have focused on brain perfusion, other organs like the lungs, kidneys and the heart have recently gained attention with improved techniques and hardware available. In particular, the move toward 3 T high field scanners in the standard clinical environments seems to push this method from the research and development stage towards clinical applications. The various applications of ASL have recently been extensively covered elsewhere [4, 35] and only a short summary of the active fields and most recent applications will be listed here. Two main categories exist – the basic science, mainly neuroscience, and the general clinical use.

Neuroscience The ability to measure CBF is very important for the assessment of tissue metabolism and function. The complete non-invasiveness of ASL, which allows prolonged functional studies to be performed on any volunteer, makes it a preferable choice in many neuroscience applications. Following a neuronal activation paradigm, the classical blood oxygen level dependent (BOLD) contrast [70] is a result of T2* changes due to alterations of CBF and CBV, as well as the cerebral metabolic rate of oxygen uptake (CMRO2). The ASL signal, on the other hand, is an absolute measure of CBF changes, which makes this technique more reproducible over time, as well as in between subjects [71–73]. In addition, perfusion functional MRI (fMRI) is believed to localize regions of activation more accurately [74], contrary to BOLD, which is affected by the change in deoxygenated blood in draining venous vessels resulting in additional signal from ‘‘down stream’’ areas. In addition to flow information, the BOLD signal can also be extracted from an ASL [75, 76]. Recent research using these techniques includes: Lu et al [77], who combined ASL, BOLD and vascular space occupancy (VASO) and Obata et al [78], who applied the ‘‘balloon model’’ to simultaneously acquire BOLD and ASL data. Finally, Hoge et al [79] combined optical and ASL imaging methods. All three studies were aimed at developing a more comprehensive picture of the physiological events accompanying activation. For pharmaceutical validation, the suitability of ASL for prolonged functional studies is advantageous for studying perfusion-altering drugs. Alternative techniques often require the injection of a tracer, which cannot be repeated more than a few times and would possibly interfere with the drug itself. Among investigated pharmaceutical agents are the vasodilators acetazolamide [80, 81] and 2-chloradonosine [82, 83], as well as The British Journal of Radiology, August 2006


the vasoconstrictor indomethacin [84, 85] and finally isoflurane, which is used for anaesthesia [86, 87]. This field can be expected to expand with broader awareness of ASL and availability of the sequences used.

Clinical In the clinics, an important issue is how well ASL performs when compared with established ‘‘gold-standard’’ methods. Validation studies have been carried out in animals [57, 88, 89] using radioactive microspheres and 14C-iodoantipyrine autoradiography, as well as in humans [90] comparing ASL and PET. In general, good correlation is observed for grey matter, whereas whitematter often shows an underestimation of CBF. This is mainly attributed the prolonged arrival times and low perfusion values of white matter, which is in the lower range of measurable flow values using ASL. The benefits of ASL particularly suit the requirements of the paediatric population, where there is a need to avoid ionizing radiation as procedures are likely to be repeated for disease monitoring and venous cannulation can be highly traumatic. In addition, the general higher blood flow seen in children increases the difference signal DM and thereby the perfusion SNR [91], making ASL a very promising tool for paediatric perfusion studies [92]. Another particularly appropriate clinical arena is that of cerebrovascular diseases. In fact, stroke is the third cause of death behind cardiovascular diseases and cancer. Consequently, a lot of resources are put into the research of this disease – how to prevent it, the immediate therapy following stroke and post-stroke rehabilitation procedures. In recent years, ASL techniques have become an alternative tool to CT and bolus tracking MRI for studying the mechanisms of stroke and the underlying processes of brain damage resulting from ischaemia. Studies of the evolution of acute stroke and validation of outcome predictors for the identification of potentially salvageable tissue have been performed in

animal models [93–96] and patients [97, 98] using both PASL and CASL. For the assessment of the different perfusion territories and eventually collateralized flow in steno-occlusive patients, a range of ASL sequences has been made capable of labelling individual perfusion territories of the major feeding vessels one at a time [24, 27, 28, 32, 99–102]. An example of regional perfusion imaging of a healthy volunteer is shown in Figure 7. These techniques are very promising in the evaluation of the successful recruitment of subsidiary blood vessels, which are believed to play an important role in the clinical outcome of patients with cerebral artery occlusion. Until recently, diagnostic strategies to evaluate the collateral circulation could be divided into those that directly visualized collateral blood vessels, such as conventional X–ray angiography and indirect methods that assess tissue perfusion, including acetazolamidechallenge tests of cerebrovascular reserve. This new class of ASL techniques makes the combination of both methods possible for the first time and quantitative information of the flow territories supplied by each major brain vessel is obtainable. As such, these methods could become alternatives to conventional X–ray based subtraction angiography, which is so far the only modality giving temporal as well as spatial regional blood flow information. Using these techniques, Van Laar et al [103] showed a wide interindividual variability in the perfusion territories caused mainly by anatomical variations of the circle of Willis in a population study of 115 healthy ageing volunteers. ASL has also been used in the evaluation of cancer. The perfusion change in tumours depends on their aggressiveness or grade, which makes perfusion information important when selecting and evaluating therapies. Repeated ASL perfusion measurements can be used to monitor the effect of blood flow regulating agents and other anticancer therapeutics such as anti–angiogenic agents [104–107]. Finally, ASL has also been applied to organs other than the brain, although the majority of research and applications have focused on the latter. Applications range from

Figure 7. Regional perfusion image (RPI) acquired using QUASAR [124] on a 23-year-old healthy female subject.


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combined lung perfusion using ASL and ventilation scans using hyperpolarized helium-3 for evaluation of pulmonary diseases [108–111], to cardiac [112–114] and renal [6, 115–119] perfusion imaging. However, these techniques are very challenging, mainly due to nonrigidity and increased movement of these organs as compared with the brain.

Conclusion The evolution of ASL since its invention by Williams et al [8, 9] in 1992 is remarkable. Numerous sequences have been developed, each solving problematic issues such as magnetization transfer effects, insensitivity to transit times etc. in their own way. Along the way, the quantitative aspect has been the leading factor and various models have been proposed for absolute CBF quantification. However, there are still improvements to be made and, with the move from traditional brain perfusion imaging to abdominal organs, there are certainly more challenges ahead. Despite the remarkable progress, ASL has still not overtaken traditional invasive methods in the clinics. One reason for this is the intrinsically low SNR, making averaging necessary, and therefore good cooperation from the subject is essential. This is not always the case with, for instance, stroke patients or children. Nevertheless, the great improvements in hardware and the increasing availability of high field clinical scanners pre-installed with ASL sequences seem to put more focus on these techniques. Robust sequences for regional perfusion imaging and functional studies of either cognitive or pharmacological nature are also gaining more and more interest, and the complete non-invasiveness of these techniques will always be an attractive asset.

Acknowledgments The authors would like to thank Prof. Alan Jackson for the opportunity to write this review article. This work was supported in part by Philips Medical Systems, and the following grants: # NMRC/0855/2004, NMRC/ CPG/009/2004, NMRC/0919/2004 and NHGA-RPR/ 04012.

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Review article: Non-invasive measurement of perfusion 99. Davies NP, Jezzard P. Selective arterial spin labeling (SASL): perfusion territory mapping of selected feeding arteries tagged using two-dimensional radiofrequency pulses. Magn Reson Med 2003;49:1133–42. 100. Hendrikse J, van der Grond J, Lu H, van Zijl PC, Golay X. Flow territory mapping of the cerebral arteries with regional perfusion MRI. Stroke 2004;35:882–7. 101. Trampel R, Mildner T, Goerke U, Schaefer A, Driesel W, Norris DG. Continuous arterial spin labeling using a local magnetic field gradient coil. Magn Reson Med 2002;48:543–6. 102. Werner R, Norris DG, Alfke K, Mehdorn HM, Jansen O. Continuous artery-selective spin labeling (CASSL). Magn Reson Med 2005;53:1006–12. 103. van Laar PJ, Hendrikse J, Golay X, Lu H, van Osch MJ, van der GJ. In-vivo flow territory mapping of major brain feeding arteries. Neuroimage 2006;29:136–44. 104. Brown SL, Ewing JR, Kolozsvary A, Butt S, Cao Y, Kim JH. Magnetic resonance imaging of perfusion in rat cerebral 9L tumor after nicotinamide administration. Int J Radiat Oncol Biol Phys 1999;43:627–33. 105. Gaa J, Warach S, Wen P, Thangaraj V, Wielopolski P, Edelman RR. Noninvasive perfusion imaging of human brain tumors with EPISTAR. Eur Radiol 1996;6:518–22. 106. Silva AC, Kim SG, Garwood M. Imaging blood flow in brain tumors using arterial spin labeling. Magn Reson Med 2000;44:169–73. 107. Warmuth C, Gunther M, Zimmer C. Quantification of blood flow in brain tumors: comparison of arterial spin labeling and dynamic susceptibility-weighted contrastenhanced MR imaging. Radiology 2003;228:523–32. 108. Altes TA, Mai VM, Munger TM, Brookeman JR, Hagspiel KD. Pulmonary embolism: comprehensive evaluation with MR ventilation and perfusion scanning with hyperpolarized helium-3, arterial spin tagging, and contrastenhanced MRA. J Vasc Interv Radiol 2005;16:999–1005. 109. Lipson DA, Roberts DA, Hansen-Flaschen J, Gentile TR, Jones G, Thompson A, et al. Pulmonary ventilation and perfusion scanning using hyperpolarized helium-3 MRI and arterial spin tagging in healthy normal subjects and in pulmonary embolism and orthotopic lung transplant patients. Magn Reson Med 2002;47:1073–6. 110. Mai VM, Hagspiel KD, Altes T, Goode AR, Williams MB, Berr SS. Detection of regional pulmonary perfusion deficit of the occluded lung using arterial spin labeling in magnetic resonance imaging. J Magn Reson Imaging 2000;11:97–102. 111. Uematsu H, Levin DL, Hatabu H. Quantification of pulmonary perfusion with MR imaging: recent advances. Eur J Radiol 2001;37:155–63. 112. Belle V, Kahler E, Waller C, Rommel E, Voll S, Hiller KH, et al. In vivo quantitative mapping of cardiac perfusion in rats using a noninvasive MR spin-labeling method. J Magn Reson Imaging 1998;8:1240–5. 113. Floyd TF, McGarvey M, Ochroch EA, Cheung AT, Augoustides JA, Bavaria JE, et al. Perioperative changes in cerebral blood flow after cardiac surgery: influence of anemia and aging. Ann Thorac Surg 2003;76:2037–42. 114. Reeder SB, Atalay MK, McVeigh ER, Zerhouni EA, Forder JR. Quantitative cardiac perfusion: a noninvasive spinlabeling method that exploits coronary vessel geometry. Radiology 1996;200:177–84. 115. Berr SS, Hagspiel KD, Mai VM, Keilholz-George S, KnightScott J, Christopher JM, et al. Perfusion of the kidney using


116.

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123.

124.

125.

126.

127.

128.

129. 130. 131.

extraslice spin tagging (EST) magnetic resonance imaging. J Magn Reson Imaging 1999;10:886–91. De Bazelaire C, Rofsky NM, Duhamel G, Michaelson MD, George D, Alsop DC. Arterial spin labeling blood flow magnetic resonance imaging for the characterization of metastatic renal cell carcinoma(1). Acad Radiol 2005;12: 347–57. Karger N, Biederer J, Lusse S, Grimm J, Steffens J, Heller M, et al. Quantitation of renal perfusion using arterial spin labeling with FAIR-UFLARE. Magn Reson Imaging 2000;18:641–7. Prasad PV, Kim D, Kaiser AM, Chavez D, Gladstone S, Li W, et al. Noninvasive comprehensive characterization of renal artery stenosis by combination of STAR angiography and EPISTAR perfusion imaging. Magn Reson Med 1997;38:776–87. Wang JJ, Hendrich KS, Jackson EK, Ildstad ST, Williams DS, Ho C. Perfusion quantitation in transplanted rat kidney by MRI with arterial spin labeling. Kidney Int 1998;53:1783–91. Detre JA, Zhang W, Roberts DA, Silva AC, Williams DS, Grandis DJ, et al. Tissue specific perfusion imaging using arterial spin labeling. NMR Biomed 1994;7:75–82. Wong EC, Buxton RB, Frank LR. Implementation of quantitative perfusion imaging techniques for functional brain mapping using pulsed arterial spin labeling. NMR Biomed 1997;10:237–49. Jahng GH, Zhu XP, Matson GB, Weiner MW, Schuff N. Improved perfusion-weighted MRI by a novel double inversion with proximal labeling of both tagged and control acquisitions. Magn Reson Med 2003;49:307–14. Chen Q, Siewert B, Bly BM, Warach S, Edelman RR. STAR–HASTE: perfusion imaging without magnetic susceptibility artifact. Magn Reson Med 1997;38:404–8. Petersen ET, Lim TC, Golay X. A model-free arterial spin labeling quantitative approach for perfusion MRI. Magn Reson Med 2006;55:219–32. Schwarzbauer C, Morrissey SP, Haase A. Quantitative magnetic resonance imaging of perfusion using magnetic labeling of water proton spins within the detection slice. Magn Reson Med 1996;35:540–6. Tanabe JL, Yongbi M, Branch C, Hrabe J, Johnson G, Helpern JA. MR perfusion imaging in human brain using the UNFAIR technique. Un-inverted flow-sensitive alternating inversion recovery. J Magn Reson Imaging 1999;9:761–7. Mai VM, Berr SS. MR perfusion imaging of pulmonary parenchyma using pulsed arterial spin labeling techniques: FAIRER and FAIR. J Magn Reson Imaging 1999;9:483–7. Mai VM, Hagspiel KD, Christopher JM, Do HM, Altes T, Knight-Scott J, et al. Perfusion imaging of the human lung using flow-sensitive alternating inversion recovery with an extra radiofrequency pulse (FAIRER). Magn Reson Imaging 1999;17:355–61. Zhou J, van Zijl PC. Perfusion imaging using FAIR with a short predelay. Magn Reson Med 1999;41:1099–107. Zhou J, Mori S, van Zijl PC. FAIR excluding radiation damping (FAIRER). Magn Reson Med 1998;40:712–9. Schwarzbauer C, Heinke W. BASE imaging: a new spin labeling technique for measuring absolute perfusion changes. Magn Reson Med 1998;39:717–22.

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Book review Michael Faraday and the electrical century. By I R Morus. pp. 192, 2004 (Icon Books, Cambridge, UK) £9.99 ISBN 1-84046-540-9 At the beginning of the 20th century the subject of electricity continued to fascinate many medical men including the early pioneering radiologists who believed electricity might have a major therapeutic role. One such early pioneer was John Mcintyre from Glasgow who was initially appointed to Glasgow Royal Infirmary in 1885 as medical electrician before turning his enthusiasm to Xrays 10 years later on learning of Roentgen’s discovery. Indeed the precursor of the Radiology Section of the Royal Society of Medicine was known as the British Electrotherapeutic Society which was formed in 1902 in the rooms of the Medical Society of London as a breakaway group from the Roentgen Society. A journal called Medical Electrology and Radiology existed and was the Electrotherapeutic Society’s official publication. Radiologists in those days were interested in the therapeutic role of electricity. Indeed in 1907, at the inauguration of the Electrotherapeutic Section of the Royal Society of Medicine, Dr Dean Butcher gave the presidential address entitled ‘‘The Future of Electricity in Medicine’’. Thus we can see that the subject of electricity proved to be of fascination to the early pioneering radiologists soon after the advent of Roentgen’s discovery. Michael Faraday of course was greatly interested in the subject of electricity and became one of the greatest figures in 19th century science. His experiments transformed the science of electricity and he is credited with, among other things, inventing the electric motor. He also investigated the relationship between electricity and magnetism and showed that currents of electricity could be produced by a moving magnet, which led to the eventual development of the electrical power industry. In this splendid biography of Faraday by Morus an expert on the scientific study of electricity in the 19th

702

century, Faraday’s seminal contribution to the subject of electricity is discussed and placed within the context of the scientific activities of the Victorian age. We are taken on a tour through his early humble upbringing and given an insight into the scientific milieu of early Victorian London in which Farraday worked and would eventually climb up to reach the highest echelons of the scientific establishment of that era. The relationship between Sir Humphrey Davy and Faraday is covered and his activities at the Royal Institution are elegantly described. The biography of Faraday is set against a backdrop of the scientific activities of London in the Victorian era. Faraday’s meticulous experiments and rapid scientific progress are elegantly described. It was interesting to read of the petty jealousies and quibbles between the leading scientific personalities of the time and smile at how little things have changed in the intervening 150 years regarding scientific rivalries. Faraday was a scientific genius who rose from humble beginnings to reach the pinnacles of the scientific establishment of his day. His research transformed the subject of electricity and laid the foundations for the future development of the early electrical industry in the early 20th century. He influenced a generation of subsequent scientists including James Clarke Maxwell, Professor of Physics at the University of Cambridge. Faraday became a Fellow of the Royal Society and became a member of several Societies in Europe and accumulated numerous honours. He was even offered the presidency of the Royal Society in 1855 but refused. He died in 1867. This is a splendid short biography of Faraday beautifully written and a delight to read. It is written for a wide audience and I am sure many radiologists and other doctors will find it of particular interest to learn a little about our scientific heritage.

A K BANERJEE



CASE OF THE MONTH

A young female with galactorrhoea and sudden onset chest pain 1

C J DAS,

MD, DNB,

1

R SHARMA,

MD,

1

A SEITH,

MD

and 2M P BARUAH,

DM

Departments of 1Radiodiagnosis and 2Endocrinology, All India Institute of Medical Sciences, New Delhi-29, India

Received 7 March 2005 Revised 12 July 2005 Accepted 16 August 2005 DOI: 10.1259/bjr/80615044 ’ 2006 The British Institute of Radiology

A 27-year-old married woman presented to the gynaecology out-patient department with a 3 month history of galactorrhoea and irregular menstrual cycle. She also complained of polyuria and polydipsia. No history of visual problems could be elicited. After initial evaluation, she was referred to the endocrinology department for further investigation of her hormonal status. Clinical examination was essentially within normal limits except for the galactorrhoea. Her pertinent laboratory investigations included high serum prolactin level (150 ng ml21), low levels of gonadotropins (undetectable/ELISA), cortisol (3.8 mg ml21) and thyroid stimulating hormone (3.8 mg ml21), suggesting panhypopituitarism. A high resolution contrast enhanced MRI (CEMRI) of the sella was performed. MRI revealed a soft tissue mass in the suprasellar region involving hypothalamus, tuber cinerium and extending into the pituitary stalk causing its expansion. The mass was hypointense on T1 weighted images (Figure 1a) and hyperintense on T2 weighted images (Figure 1b) showing intense

Figure 2. Chest radiograph posteroanterior view revealed reticular opacities and pneumothorax on the left side.

(a)

(b)

(c)

Figure 1. Contrast enhanced MRI sella showing soft tissue mass in suprasellar region involving hypothalamus, tuber cinerium extending into the pituitary stalk, which is hypointense on (a) T1 weighted images and hyperintense on (b) T2 weighted images, showing intense homogeneous enhancement on (c) post-gadolinium images.


703

C J Das, R Sharma, A Seith and M P Baruah

(a)

(b)

Figure 3. (a,b) Follow-up contrast enhanced MRI after anti-tubercular treatment does not show any change in the size of the suprasellar mass.

homogeneous enhancement on post-gadolinium images (Figure 1c). 2 months later, the patient presented with one episode of sudden onset dyspnoea and severe nonproductive cough. A posteroanterior (PA) chest radiograph revealed reticular opacities and pneumothorax on the left side (Figure 2). What is the possible diagnosis with this clinical presentation and radiological appearance?

Discussion A provisional diagnosis of tuberculoma was suggested, based on the MR appearance and in view of the high prevalence of this disease in India. The patient was put on anti-tubercular treatment, but her clinical condition did not improve even after 12 months and a repeat CEMRI (Figure 3) at this time did not show any change in the size of the suprasellar mass. In view of the history of sudden onset dyspnoea and pneumothorax seen on the chest radiograph, a high resolution CT (HRCT) scan of chest was performed which revealed multiple bizarrely-shaped cystic lesions in both lungs (Figure 4). Skeletal survey did not reveal any bony lesions. The diagnosis was reviewed and, based on the clinical presentation as well as MR and HRCT appearance, a final diagnosis of pulmonary Langerhans’ cell histiocytosis (PLCH) with involvement of the hypothalamicpituitary axis causing panhypopituitarism and diabetes insipidus was made. A thoracoscopic lung biopsy was performed and histopathology confirmed the diagnosis of PLCH. The patient was treated with hormone replacement therapy for panhypopituitarism and Carbamazapine for diabetes insipidus. PLCH was treated with a combination of chemotherapy and 704

radiotherapy to which she responded. On clinical follow-up the patient is doing well. PLCH is an uncommon disorder of unknown aetiology that typically affects young adults who are smokers [1]. Previously called histiocytosis X, this is a group of diseases characterized by proliferation of histiocytes [1]. There are unifocal and multifocal forms of this disease. The unifocal disease shows solitary lytic bone lesion. This form spares the hypothalamic pituitary axis. Adult PLCH occurs most commonly as a disease isolated to the lungs. However, adult PLCH can occur as a part of the Hand-Schuller-Christian syndrome – a triad of diabetic insipidus, exophthalmos and lytic bone lesions which is

Figure 4. High resolution CT (HRCT) scan of chest shows multiple bizarrely-shaped cystic lesions in both lungs. The British Journal of Radiology, August 2006

Case of the month: A young female with galactorrhoea

commonly seen in children [1]. Taking LCH as a whole, about one-third of cases are seen in adults and 16–34% of these have pulmonary involvement [2]. In the past, PLCH was regarded as a disease with male predominance, but is now considered to be equally prevalent in both sexes [3], a change that possibly reflects altered smoking habits. Diabetes insipidus is the most common endocrinopathy reported with adult PLCH with the prevalence being about 5–9% [2]. Other endocrine abnormalities include growth hormone deficiency, hypothyroidism and abnormal glucose tolerance [4, 5]. The characteristic radiographic appearance of PLCH consists of diffuse, symmetric nodular/reticulonodular opacities predominantly in the mid and upper zones with relative sparing of the lung bases [3, 6]. For pulmonary involvement, HRCT is the investigation of choice [7, 8]. The principal findings are cysts with irregular margins and nodules, often in combination, giving a characteristic appearance. These cysts are thin walled and are usually less than 10 mm in diameter. Although most of the cysts appear round, they can also have bizarre shapes, being bilobed, clover-leaf shaped, or branching in appearance. The nodules are usually solid, centrilobular in location and measure less than 5 mm in diameter [7, 8]. The main differential diagnosis of PLCH on HRCT is lymphangioleiomymatosis (LAM). The most useful differentiating features are: the distribution of cysts, which are wide spread in PLCH, decreasing towards the lower zone and sparing the costophrenic angles, but evenly diffuse in LAM; the possible presence of nodules in PLCH and their rarity in LAM; the tendency for LAM cysts to be uniformly rounded, whereas some of the cysts in PLCH are irregular [9]. CEMRI is the ideal investigation for delineation of intracranial involvement. CT and MRI show thickening and enhancement of the infundibulum and hypothalamus. The posterior pituitary bright spot may be absent. The lesions are hyperintense on T2 weighted images [1]. Differential diagnoses include tuberculosis, sarcoidosis, hypothalamic hamartoma and chiasmatic glioma. Characteristic bone lesions and a classical clinical


picture aid in the diagnosis of PLCH. Treatment of PLCH is problematic because of the unpredictable and variable outcome. Abstaining from smoking may halt the disease progression. Radiotherapy and chemotherapy may be needed for the intraocular involvement. Lung transplantation for terminal respiratory insufficiency has also been advocated [10].

References 1. Lum C, Kucharczyk W, Montanera WJ, Becker LE. Sella Turcica and parasellar region. In: Scott Atlas, editor. Magnetic resonance imaging of the brain and spine, 3rd edn. Lippincott Williams and Wilkins, 2002:1340. 2. Malpas JS. Langerhans’ cell histiocytosis in adults. Hematol/Oncol Clin North Am 1998;12:259–68. 3. Friedman PJ, Liebow AA, Sokoloff J. Eosinophilic granuloma of lung. Clinical aspects of primary pulmonary histiocytosis in adult. Medicine 1981;60:385–96. 4. Soliman AT, Alsalmi I, Banna NE, Asfour M. Endocrine aspects of Langerhans cell histiocytosis. Indian J Pediatr 1996;63:402–6. 5. Rami B, Schneider U, Wandl-Vergesslich K, Frisch H, Schober E. Primary hypothyroidism, central diabetes insipidus and growth hormone deficiency in multisystem Langerhans cell histiocytosis: a case report. Acta Paediatr 1998;87:112–4. 6. Lewis JG. Eosinophilic granuloma and its variants with special reference to lung involvement. J Med 1964;33:337–59. 7. Grenier P, Valeyre D, Cluzel P, et al. Chronic diffuse interstitial lung disease: diagnostic value of chest radiography and high resolution CT. Radiology 1991;179: 123–32. 8. Brauner MW, Grenier P, Mouelhi MM, Mompoint D, Lenoir S. Pulmonary histiocytosis X: evaluation with high resolution CT. Radiology 1989;172:255–8. 9. Bonelli FS, Hartman TE, Swensen SJ, et al. Accuracy of high resolution CT in diagnosing lung diseases. AJR Am J Roentgenol 1988;170:1507–12. 10. Wilson AG. Pulmonary diseases of unknown origin and miscellaneous lung disorders in Peter Armstrong’s Imaging of diseases of the chest, 3rd edn. Mosby, 2000:655–60.

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September 2006 Volume 79 Issue 945

September 2006, Volume 79, Issue 945 Quantitative optical coherence tomographic elastography: method for assessing arterial mechanical properties ●

Low-carbohydrate diet induced reduction of hepatic lipid content observed with a rapid non-invasive MRI technique ●

●

Translational research in radiotherapy trials

Oesophageal dysmotility in systemic sclerosis: comparison of HRCT and scintigraphy ●

Relationship between the growth pattern of nasopharyngeal cancer and the cervical lymph nodes based on MRI findings: can the cervical radiation field be reduced in patients with nasopharyngeal cancer? ●

Six years experience in intracoronary brachytherapy procedures: patient doses from fluoroscopy ●

What is the optimum breast plan: a study based on the START trial plans ●

Feasibility of automated matching of supine and prone CT-colonography examinations ●

Improved motion compensation in 3D-CT using respiratorycorrelated segment reconstruction: diagnostic and radiotherapy applications ●

A mathematical model for patient skin dose assessment in cardiac catheterization procedures ●

●

Nurse-led central venous catheter service: Christie experience

Breast movement during normal and deep breathing, respiratory training and set up errors: implications for external beam partial breast irradiation ●

●

Book review

●

A case of spinal cord compression of unknown cause


SHORT COMMUNICATION

Quantitative optical coherence tomographic elastography: method for assessing arterial mechanical properties 1,2

J ROGOWSKA,

PhD,

1

N PATEL,

BS,

1

S PLUMMER,

BS

and

1,2

M E BREZINSKI,

MD, PhD

1

Brigham and Women’s Hospital, 75 Francis Street, Boston, MA 02115 and 2Harvard Medical School, Boston, Massachusetts, USA ABSTRACT. Optical coherence tomography elastography represents a potentially attractive new technique for measuring elastic properties of tissues on a micron scale. In this study, the feasibility of optical coherence tomography (OCT) to study the mechanical properties of phantoms and atherosclerotic arterial samples is reported. The elastic modulus of tissue-mimicking phantoms was measured using OCT and correlated with mechanical measurements. The results indicate that elastography based on OCT represents an attractive technique for evaluating the mechanical properties of tissues.

Knowledge of the mechanical properties of vascular tissue could provide diagnostic information about a range of vascular diseases, from pulmonary hypertension to coronary atherosclerosis. Therefore, it is important to develop and evaluate new quantitative methods for measuring the elastic modulus of normal and abnormal arterial tissue. The term ultrasound elastography was introduced for the first time by Ophir and colleagues in 1991 [1]. The elastography technique was based on applying a pressure on the examined tissue and in estimating the induced strain distribution by tracking the tissue motion [1–3]. An ultrasound elastography is limited by a resolution between 80 mm and 100 mm, which is likely insufficient for vascular assessments. A technology that has shown considerable promise as a method of high resolution intravascular imaging and elastography is optical coherence tomography (OCT) [4–9]. OCT is a micrometre scale imaging technology analogous to ultrasound, measuring the back-reflection of near-infrared light rather than sound. Current OCT systems that can be used in vivo have a resolution between 10 mm and 20 mm. Recently, several semi-quantitative approaches to OCT elastography have been demonstrated using speckle tracking [10–12]. However, in order for the technique to be used clinically, the quantitative accuracy of tissue elasticity measurements must be established, which is the focus of this paper. In this paper, we first describe and implement an OCT elastography technique on tissue phantoms, measuring Address correspondence to: Mark E Brezinski, Brigham and Women’s Hospital, MRB 106, 75 Francis Street, Boston, MA 02115, USA. This research is sponsored in part by the National Institutes of Health, Contracts NIH-RO1-AR44812 (MEB), NIH R01 AR46996 (MEB), NIH R01- HL63953 (MEB), NIH-1-R01-HL55686 (MEB), NIH R01 EB000419 (MEB) and NIH-1-R29-HL55686 (MEB).

The British Journal of Radiology, September 2006

Received 23 November 2005 Revised 9 March 2006 Accepted 12 April 2006 DOI: 10.1259/bjr/22522280 ’ 2006 The British Institute of Radiology

deformation of samples to obtain stress-strain curves that characterize the linear elastic properties. Second, these results of OCT elasticity measurements are validated and quantitated with non-imaging mechanical measurements of the same phantoms. Finally, this validated OCT elastography approach was applied to atherosclerotic arterial samples.

Materials and methods Experimental elastography system OCT is analogous to ultrasound, measuring the intensity of back-reflected infrared light [4, 5]. Ultrashort light pulses or low coherent light is generated at the sample. The time for the light to be reflected back or echo delay time is used to measure distances. The intensity of back-reflection is plotted as a function of depth. The beam is then scanned across the sample to produce two- and three-dimensional data sets.

Phantom and aorta images Phantoms were made by mixing different amounts of agar, gelatin and water. Using a fixed amount of gelatin (8 g), phantoms with different hardness were made by varying the amount of agar: 0.5 g for 1% phantom, 1.0 g for 2% phantom, 3 g for 3% phantom. Gelatin and agar were mixed together and dissolved in 50 ml of boiling water. The 0.25 g of activated charcoal particles were added as scatterers. After refrigeration, the gelatin phantoms were cut into 2.5 cm62.5 cm61 cm thick block shaped samples, covered with a cover slip, and scanned. The elastic properties of the phantoms were confirmed by applying weights of 49.81 g, 102.88 g and 707

J Rogowska, N Patel, S Plummer and M E Brezinski

133.03 g. The changes in tissue width were measured both with callipers and OCT imaging. All measurements were repeated three times for each of the phantoms, thus producing three sets of data. The OCT measurements were corrected for refractive index. The human atherosclerotic aorta samples were obtained post-mortem and stored at 0 ˚C with 0.1% sodium azide. The samples were placed under the OCT system and scanned before and after compression, using a procedure similar to phantom scanning. The weights applied to the aorta samples were 2.29 g, 5.60 g and 8.21 g. The OCT scanning was performed in the axial cross-sections of the samples, covering 2 mm62 mm areas in the centre, far away from the sample edges. The OCT image resolution was 400 by 400 pixels, where each pixel was 0.005 mm60.005 mm.

Calculation of elastic modulus of phantoms Our goal was to calculate the local values of elastic modulus in the phantoms. If we assume that the phantoms are elastic, isotropic and incompressible, and the stress is applied uniformly in axial direction, then the Young’s modulus (or elastic modulus) E is defined by [13]: E~

s e

ð1Þ

where s is the axial (or normal) stress (defined as the force F perpendicular to the cross sectional area, divided by the cross sectional area A): s~

F A

ð2Þ

and e is the axial strain (defined as the fractional change in sample thickness L): e~

DL L

ð3Þ

By applying weights, samples were compressed, and their thicknesses before and after compression were measured using callipers and OCT displacement vectors. The image data were processed pixel-by-pixel and the total axial displacement vectors were calculated using cross-correlation [10, 11]. For strains up to 5%, the relationship between force and displacement is usually linear [14]. Based on Equation (2), the equivalent stress applied to phantoms was computed by dividing the force by the surface dimensions of the samples, which yielded stress values of 0.77 kPa, 1.60 kPa and 2.06 kPa for three compressions. The optical coherence tomographic elastography (OCTE) displacement values in pixels were converted to millimetres and corrected using refraction index. By measuring the strain (Equation (3))

(a)

(b)

(c)

(d)

Figure 1. A typical optical coherence tomography (OCT) 2% phantom images after applying stress of (b) 0.772 kPa, (c) 1.595 kPa and (d) 2.062 kPa to original phantom (a). The displacement vectors are shown in red. 708


Short communication: Quantitative optical coherence tomographic elastography

Figure 2. Young’s modulus for 1%, 2%, and 3% phantoms estimated using two techniques: callipers’ measurements and OCT elastography. The bars indicate standard deviation. The differences between calliper and OCTE measurements were found to be non-significant (p50.6, p50.25 and p50.38 for 1%, 2% and 3% phantoms, respectively).

(a)

(b)

(c)

(d)

Figure 3. Aorta images after applying stress of (b) 0.0355 kPa, (c) 0.0868 kPa and (d) 0.1273 kPa to original image (a). Figure 3a indicates location of intima (I), media (M) and glass compression plate (G). The displacement vectors are shown in red. The British Journal of Radiology, September 2006

709

J Rogowska, N Patel, S Plummer and M E Brezinski

(a)

(b)

(c) Figure 4. Axial displacement maps for compressed aorta images shown on Figure 3.

for three different applied stresses, the stress-strain relationship was calculated using a linear regression. The Young’s modulus was estimated from the slope of the strain-stress linear function. All measurements were repeated three times for three sets of samples, and then averaged.

to Young’s modulus using a conversion factor C [13, 15]:

Estimation of the elastic modulus of aorta

Results

Once the stress-strain measurements are obtained on the phantoms, the slopes of these stress-strain curves are used to determine the elastic modulus values. Since the Young’s moduli (E) of the phantoms are known, and the slope values (S) of the stress-displacement curves for the phantoms can be measured, the conversion factor C relating the stress-displacement slope values to the Young’s modulus could be determined by:

The phantoms underwent compression with weights and the changes in phantom’s thickness were measured both with callipers and OCT imaging. The image data were processed pixel-by-pixel and the total axial displacement vectors were calculated. An example of the 2% phantom is shown on Figure 1. For all phantoms, stress-strain relationships using calliper and OCT measurements were calculated and Young’s modulus was estimated (Figure 2). The differences between calliper and OCT measurements were found to be non-significant (paired t-test; p50.6 for 1% phantom, p50.25 for 2% phantom and p50.38 for 3% phantom). The relationship between two Young’s modulus estimators (using callipers and OCTE) was found to be highly correlated, R250.92. Aorta samples also underwent compression with weights (Figure 3). The changes in tissue thickness were measured using OCT. The image data were processed pixel-by-pixel and the axial displacement vectors (Figure 3b–d) and displacement maps (Figure 4) were

C~

1 E1% E2% E3% z z 3 S1% S2% S3%

ð4Þ

where E1%, E2%, and E3% are elastic moduli for 1%, 2% and 3% phantoms, respectively, and S1%, S2%, and S3% are slope values from the corresponding stress-displacement curves. In order to calculate the estimated elastic modulus of the aorta (rEAorta), the slope value of the OCTE stressdisplacement curve for the aorta (SAorta) was transformed 710

rEAorta ~C SAorta

ð5Þ


Short communication: Quantitative optical coherence tomographic elastography

calculated. In order to calculate the relative elastic modulus of the aorta, we used the conversion factor C512.14, which was derived from stress-displacement phantom curves. The average estimated Young’s modulus for the aorta sample shown on Figure 4 was equal to 6.11 kPa.

Discussion This paper attempts to quantify OCT elastography as a method for assessing Young’s modulus, and therefore evaluating tissue mechanical properties. The OCT elastography system was calibrated with tissue-mimicking phantoms and the elastic modulus of the atherosclerotic arterial samples was estimated. A strong correlation existed between the mechanical measurements and those performed with OCT elastography, with no significant difference existing between the two techniques. OCT elastography was then used to estimate Young’s modulus through the use of the conversion factor. According to other reports [15], our estimated elastic modulus of 6.11 kPa for the aorta was within the range of Young’s modulus values calculated for smooth muscle (0–100 kPa) and soft tissue (0–200 kPa). Several obstacles need to be overcome for OCT elastography application in human arteries in vivo. First, controlled transmural pressures need to be applied to make OCT measurements in vivo. This can potentially be achieved with an intravascular pressure balloon or saline flushes with measured pressures. Second, it is assumed that the tissue under investigation is elastic, uniform and nearly incompressible. Based on work with ultrasound, the approximation of vascular tissue being elastic and incompressible is reasonable. However, with atherosclerotic or severely hypertensive arteries, the tensile properties will not necessarily be completely uniform. The importance of this has to be examined with future studies. Third, for correct conversion of stressdisplacement data to an elastic modulus, it is essential that the calibration sample and tissue sample have similar shape [13]. This needs to be taken into account when calibrating the technique for imaging in vivo arteries. Sample height and sample compressibility are non-critical factors for system accuracy. However, the size of the area where the stress is applied to the sample is a very critical parameter and needs to be well controlled to obtain accurate results. In summary, OCT elastography represents an attractive technique for evaluating the mechanical properties of tissue due to its micrometre scale resolution. Other OCT elastography studies have essentially focused on qualitative analysis, not on the accuracy of measured mechanical parameters. The focus of this work was to


establish quantification of OCT elastography. The study reported here was also the first attempt to calculate the elastic modulus of atherosclerotic tissue using OCT.

References 1. Ophir J, Cespedes I, Ponnekanti H, Yanzi Y, Li X. Elastography: a quantitative method for imaging the elasticity of biological tissues. Ultrason Imaging 1991;13:111–34. 2. de Korte CL, Cespedes EI, van der Steen AF, Lancee CT. Intravascular elasticity imaging using ultrasound. Ultrasound Med Biol 1997;23:725–46. 3. de Korte CL, van der Steen AF. Intravascular ultrasound elastography: an overview. Ultrasonics 2002;40:859–65. 4. Brezinski ME, Tearney GJ, Bouma BE, Izatt JA, Hee MR, Swanson EA, et al. Optical coherence tomography for optical biopsy: properties and demonstration of vascular pathology. Circulation 1996;93:1206–13. 5. Boppart SA, Bouma BE, Pitris C, Tearney GJ, Southern JF, Brezinski ME, et al. Intraoperative assessment of microsurgery with three-dimensional optical coherence tomography. Radiology 1998;208:81. 6. Jesser CA, Boppart SA, Pitris C, Stamper DL, Nielsen GP, Brezinski ME, et al. High resolution imaging of transitional cell carcinoma with optical coherence tomography: feasibility for the evaluation of bladder pathology. Br J Radiol 1999;72:1170–6. 7. Patwari P, Weissman NJ, Boppart SA, Jesser C, Stamper D, Fujimoto JG, et al. Assessment of coronary plaque with optical coherence tomography and high frequency ultrasound. Am J Cardiol 2000;85:641–4. 8. Podoleanu AGh. Optical coherence tomography. Br J Radiol 2005;78:976–88. 9. Tearney GJ, Brezinski ME, Bouma BE, Boppart SA, Pitris C, Southern JF, et al. In vivo endoscopic optical biopsy with optical coherence tomography. Science 1997;276: 2037–9. 10. Rogowska J, Patel NA, Fujimoto JG, Brezinski ME. OCT elastography of the vascular tissue – importance of crosscorrelation kernel size. OSA Biomedical Topical Meetings, OSA Technical Digest, (Optical Society of America, Washington, DC, 2002), pp. PD20.1-20.3. 11. Rogowska J, Patel NA, Fujimoto JG, Brezinski ME. Quantitative OCT elastography technique for measuring deformation and strain of the atherosclerotic tissues. Heart 2004;90:556–62. 12. Schmitt JM. OCT elastography: imaging microscopic deformation and strain in tissue. Opt Express 1998;3:199–211. 13. Erkamp RQ, Wiggins P, Skovoroda AR, Emelianov SY, O’Donnell M. Measuring the elastic modulus of small tissue samples. Ultrason Imaging 1998;20:17–28. 14. Chen EJ, Novakofski J, Jenkins WK, O’Brien WD. Young’s modulus measurements of soft tissues with application to elasticity imaging. IEEE Trans Ultrason Ferroelectr Freq Control 1996;43:191–4. 15. Samani A, Bishop J, Luginbuhl C, Plewes DB. Measuring the elastic modulus of ex vivo small tissue samples. Phys Med Biol 2003;48:2183–98.

711


SHORT COMMUNICATION

Low-carbohydrate diet induced reduction of hepatic lipid content observed with a rapid non-invasive MRI technique 1

K G HOLLINGSWORTH, D J LOMAS, FRCR

PhD,

1

M Z ABUBACKER,

FRCR,

1

I JOUBERT, 2M E D ALLISON,

FRCP

and

1

Departments of 1Radiology and 2Medicine University of Cambridge and Addenbrookes Hospital, Level 5, Box 219, Addenbrookes Hospital, Cambridge CB2 2QQ, UK ABSTRACT. Low carbohydrate diets are currently fashionable for inducing weight loss, but the metabolic effects at organ level are not well understood, especially the effect on liver fat storage. Such studies require serial hepatic fat measurements, for which liver biopsy is impractical. In 10 healthy volunteers we demonstrate the use of rapid (total 2 min acquisition time, 10 min magnet room time), non-invasive, quantitative MRI to serially measure hepatic fat changes induced by following a low carbohydrate diet for 10 days. A significant (p,0.01) reduction in hepatic fat after 3 days of dieting was observed in 5 subjects. All subjects demonstrated significant (p,0.01) reductions in hepatic fat by day 10. A strong correlation (k50.81) existed between the initial fat content and the percentage fat content reduction in the first 3 days of the diet. All subjects lost weight (average 1.7 kg at day 3 and 3.0 kg at day 10), but this was not correlated with hepatic fat loss after 3 days or 10 days of dieting. The results presented illustrate the potential value of MR hepatic fat quantification in longitudinal studies of hepatic fat content.

Hepatic steatosis is of particular interest in the western world owing to the increasing prevalence of the insulin resistance or ‘‘metabolic’’ syndrome. Hepatic steatosis and non-alcoholic steatohepatitis (NASH) are now considered a part of this condition and there is emerging evidence that, rather than being a consequence of systemic insulin resistance, they may have a causative role [1, 2]. As a result, therapeutic or dietary interventions to reduce hepatic steatosis may be more appropriate for treating insulin-resistance, rather than by treating the different aspects separately. It is estimated that 20–25% of the US population is obese and that among the obese group there is a high prevalence of nonalcoholic fatty liver disease (NAFLD). The risk of serious sequelae such as type II diabetes mellitus or progression to chronic liver disease means that this is an area requiring urgent research and assessment [3]. These types of studies require accurate serial quantification of hepatic steatosis, but currently liver biopsy (and related biochemical analysis) is the established method. As an invasive procedure, even with a minimal related morbidity and mortality, this is difficult to justify in healthy volunteers and impractical for serial measurements. Additionally the method samples only small volumes of tissue, which may lead to problems with regional fat variation and give unrepresentative results. An alternative approach is to use a validated rapid imaging method that allows absolute hepatic fat (mainly Funded by the Fund and Friends of Addenbrookes.

712

Received 6 January 2006 Revised 3 April 2006 Accepted 11 April 2006 DOI: 10.1259/bjr/23166141 ’ 2006 The British Institute of Radiology

intracellular triglyceride) estimation through fat specific chemical shift imaging corrected for T2* variation [4, 5]. Previous studies have shown a good correlation between steatosis assessed by liver biopsy and by MRI methods [6], and excellent correlation exists between hepatic fat measured by liver histology, MRI and CT methods [7, 8]. The effect of low carbohydrate diets on weight loss, insulin resistance and serum triglyceride markers has recently been the focus of intensive research efforts [9, 10], but to the best of the authors’ knowledge, the effect of such diets on hepatic fat has never been directly measured. The aim of this work was to monitor the hepatic fat response of healthy volunteers during the induction phase of a low-carbohydrate diet using the MRI method described above.

Materials and methods The study was approved by the local ethics committee and informed consent was obtained from the volunteers after the procedure was fully explained. Participants received no monetary incentive. Given the lack of consensus on the normal hepatic lipid percentage range measured by MRI we used an arbitrary limit of a pre-diet hepatic fat estimate of 7.0% or above as the entry threshold for the study. This threshold was chosen as it was the median fat percentage observed in our previous work on healthy volunteers (n525); a similar result has also been shown with this technique by Fishbein [4]. 10 The British Journal of Radiology, September 2006

Short communication: Hepatic lipid content observed with a rapid non-invasive MRI technique

healthy volunteers (3 male, 7 female, age 32–56 years) were recruited from the community: the volunteers were not taking regular medication and had no history of hepatic or biliary disease. Five further volunteers were excluded from the study after an initial MRI examination showed their hepatic fat estimate to be below 7.0%. The participants were asked to follow a low carbohydrate diet at home for 10 days, restricting carbohydrate intake to less than 20 g carbohydrate per day in the form of green salad or vegetables, but with no other restriction on total energy intake or food choice. The volunteers kept a diet diary and abstained from alcohol. Hepatic fat measurement by MRI was performed at 4 time points; immediately pre-diet, at 3 days and 10 days on the diet and 7 days after reverting to their normal diet. The volunteers were examined at the same time of day on each occasion. The initial body mass index (BMI) was recorded and the weight of the volunteers was assessed on each visit to the MRI unit. Weight was measured with the subjects changed for the MRI examination and without shoes. Adherence to the diet was monitored by the patients maintaining a food record sheet and by urinary ketone assessment to ensure that ketosis was initiated and maintained. Ketones are usually not detectable in healthy volunteers following a balanced diet. Examinations were performed on a 1.5 T whole body MRI (Excite, GEHT, Milwaukee) with an 8-channel body array. In and out of phase gradient echo scans (matrix 2566128, section (slice) thickness 10 mm, gap 1.5 mm, repetition time (TR)/echo time (TE)5180/2.2 ms (out of phase)/4.4 ms (in phase)) were acquired axially at two different flip angles (20 ˚ and 70 ˚) and a T2* map of the liver was obtained using a location-matched, multisection, multiecho gradient sequence (TR5120 ms, 16 equally spaced echoes, TE152.2 ms, TE254.4 ms). These required a total of three 20 s breath-holds in addition to an initial 20 s breath-hold study for checking the positioning. The T2* data was used to correct the inphase and out-of-phase images intensities for T2* relaxation. The fat percentage was calculated by: FP~100(Sin Sout )=2Sin

image (both values corrected for T2* relaxation). The acquisition of in-phase and out-of-phase images with different flip angles (and hence T1 weighting) allows us to distinguish whether fat or water is the majority species [4, 5]: this resolves the ambiguity that occurs in Equation (1) for fat percentage greater than 50%. Four sections centrally placed in the liver were analysed by a single operator and three circular regions of interest (ROI) with fixed area (5 cm2) were positioned over the liver parenchyma on each section, avoiding large vessels and the gallbladder. The 12 ROIs for each individual were then averaged to give a result for that time point. Significant changes between time points were assessed using a paired t-test (SPSS 12.0). Hepatic fat percentage was the primary outcome measure, as assessed from the MR images. Correlation coefficients (Pearson) were calculated between the initial fat measurement and the percentage reduction in fat after 31 days and 10 days of the diet, and also between the percentage weight reduction and the percentage reduction in hepatic fat at days 3 and 10 of the diet.

Results The diet records indicated that all the volunteers maintained the diet successfully and this was confirmed by evidence of ketosis on urine testing. All volunteers had ‘‘negative’’ ketone readings before the diet, and these all increased to at least ‘‘moderate’’ by day 3 and for the duration of the diet. The initial BMI of the volunteers (range 23–32, median 28) demonstrated that the volunteers were, at worst, moderately overweight. Table 1 shows the changes in hepatic fat percentage for the 10 individuals, at all the time-points in the trial. Figure 1 shows the relative change (compared with the pre-diet measurement) in the hepatic fat percentage graphically. Figure 2 illustrates the time course of results and error bars for subjects 1 and 2. There was no significant variation of the measured fat percentage with the location of the ROI. Five of the subjects experienced a significant (p,0.01) decrease in hepatic fat in the first 3 days of dieting. Two of these subjects displayed a further significant change between day 3 and day 10 of the diet. All 10 subjects experienced a significant

ð1Þ

where FP is the fat percentage, Sin is the intensity from the in-phase image, Sout is the intensity from the out-of-phase

Table 1. Selected BMI, hepatic fat and weight data for the volunteers Subject

Initial BMI (kg m22)

Initial hepatic fat (%)

3 day 10 day 17 day Weight hepatic fat (%) hepatic fat (%) hepatic fat (%) loss (kg) on completion of diet (day 10)

Weight change (kg) 1 week after diet (day 17)

1 2 3 4 5 6 7 8 9 10

30 32 29 27 23 28 26 32 26 28

12.1 8.5 11.8 7.0 9.0 7.2 11.2 8.8 7.7 10.2

10.0a 7.1a 9.0a 7.0 7.9a 6.9 9.8 8.2 7.1 9.1a

+0.5 +2.0 21.5 21.5 0 – +1.0 +0.5 +0.5 22

9.4a 7.4a 6.7a 5.6a 7.5a 6.4a 9.0a 7.0a 5.0a 6.1a

9.9 6.7 6.4 8.4b 7.7 – 9.9 8.6b 6.4b 5.0b

4 4 4.5 2 3 3 3 3 1 2

a

p,0.01 compared with initial measurement. p,0.01 compared with previous measurement (17 days measurement only). BMI, body mass index.

b


713

K G Hollingsworth, M Z Abubacker, I Joubert et al

Figure 1. Relative change in hepatic fat for all 10 healthy volunteers compared with hepatic fat percentage at day 0.

Figure 2. Plot of individual hepatic fat measurements for subjects 1 and 2 with error bars.

(p,0.01) decrease in hepatic fat within 10 days of starting the diet. It was noted that the volunteers with the highest initial fat content had the greatest percentage decrease during the first 3 days of the diet (k50.81, range 0–24%), although the correlation at day 10 was weaker (k50.42, range 11–43%). All the subjects lost weight (mean weight loss: 1.7 kg at day 3 and 3.0 kg at day 10) but this was not correlated with changes in hepatic fat percentage. 1 week after stopping the diet, five subjects gained weight, three lost weight and one maintained the same weight (one was lost to follow up after cessation of the diet), though there was no correlation with hepatic fat change in the week after cessation of the diet.

Discussion This study evaluates hepatic fat response to a low carbohydrate diet in 10 healthy volunteers and demonstrates a significant reduction of hepatic fat levels. To the best of our knowledge this is the first study to use quantitative image-based MR to demonstrate a significant reduction in hepatic fat during a low carbohydrate diet intervention. The dominance of fat and protein in this diet means that a decrease in liver fat reflects mobilization of hepatic lipid stores as an energy source and a contributor to the related ketosis. All subjects lost 714

weight, demonstrating that either the diet itself or its appetite suppressant effect is hypocaloric. The limitations of this pilot study were the small number of subjects studied and the lack of biochemical blood correlates, in particular the evaluation of insulin resistance by clamping methods. Although all subjects experienced reduced hepatic fat levels during the study, there was variation in the percentage of initial fat lost by day 10 (range 11–43%). Although there was an empirical correlation between initial fat percentage and the percentage of hepatic fat lost by day 3, evaluation of the insulin resistance changes of the subjects may elucidate the reason for the differing degrees of response, particularly since there was no correlation between the percentage hepatic lipid reduction and the reduction in weight. This would also provide insight into whether a similar response could be expected from the different patient groups with hepatic steatosis. A non-dieting (control) group was not assessed in this study but previous work on observing liver fat changes in healthy, non-dieting volunteers has shown weekly changes of no more than approximately 1% [11]. There was also no detailed study of the subjects dietary and alcohol habits for more than 3 days before commencing the diet: such a standardization step may be important in drawing full quantitative conclusions from a diet study of short duration. Future work in larger volunteer and patient groups will address these limitations. There have been few comparable studies as the majority of metabolic studies do not measure hepatic fat content directly (presumably owing to the ethical difficulty of performing serial liver biopsies), choosing rather to measure serum triglycerides. In a study of overweight subjects fed a high fat (56% of calorie intake) diet (with carbohydrate) for a period of 2 weeks the liver accumulated fat (average 35% of the initial fat content) [12] compared with an isocaloric, low fat (16% of calorie intake) where there was a reduction in liver fat (average 20% of the initial fat content). A study of the effect of a hypocaloric low carbohydrate diet followed for 14 days by patients with type II diabetes found that there was an improvement in insulin resistance and a decrease in plasma triglycerides, though liver fat was not directly measured [10]. Other authors have reported evidence that low carbohydrate diets can alter body fat composition [13]. This is of importance given the emerging literature on the role of hepatic fat in the development of systemic insulin resistance, leading to type II diabetes mellitus, hyperlipidaemia, hypertension and increased artherosclerotic risk. This pilot study suggests that a low carbohydrate diet may have a role in modifying hepatic fat and hence insulin resistance. Alternative non-invasive methods such as ultrasound and CT have been proposed, but ultrasound is limited by lack of specificity and CT by the use of ionizing radiation, which is difficult to justify in healthy individuals. The method used in this study can be implemented on the majority of currently installed MRI systems, and is simpler to implement than fat measurement by proton MR spectroscopy, which requires longer acquisition times and specialist analysis [6, 12]: spectroscopy methods cannot easily be used to sample the entire liver and cannot detect focal fat variation in the liver in a short examination. One study [14] finds that carbon-13 The British Journal of Radiology, September 2006

Short communication: Hepatic lipid content observed with a rapid non-invasive MRI technique

spectroscopy of hepatic lipids is in excellent agreement with morphometric analysis of biopsy specimens. However, this is principally a research technique and is not available in the vast majority of clinical settings. This pilot study demonstrates the practical implementation and utility of MRI fat quantification as a tool in serial studies of hepatic fat content. In healthy volunteers the method demonstrated significant hepatic fat reduction resulting from a low carbohydrate dietary intervention.

Acknowledgments This study was funded by the Fund and Friends of Addenbrookes. We thank the staff of the Magnetic Resonance Imaging and Spectroscopy Unit for support in running the study

References 1. Samuel VT, Liu ZX, Qu X, Elder BD, Bilz S, Befroy D, et al. Mechanism of hepatic insulin resistance in non-alcoholic fatty liver disease. J Biol Chem 2004;279:32345–53. 2. Tiikainen M, Hakkinen AM, Korsheninnikova E, Nyman T, Makimattila S, Yki-Jarvinen H. Effects of rosiglitazone and metformin on liver fat content, hepatic insulin resistance, insulin clearance, and gene expression in adipose tissue in patients with type 2 diabetes. Diabetes 2004;53:2169–76. 3. Saadeh S, Younossi ZM, Remer EM, Gramlich T, Ong JP, Hurley M, et al. The utility of radiological imaging in nonalcoholic fatty liver disease. Gastroenterology 2002;123:745–50. 4. Fishbein MH, Gardner KG, Potter CJ, Schmalbrock P, Smith MA. Introduction of fast MR imaging in the assessment of hepatic steatosis. Magn Reson Imag 1997;15:287–93.


5. Hussain HK, Chenevert TL, Londy FJ, Gulani V, Swanson SD, McKenna BJ, et al. Hepatic fat fraction: MR imaging for quantitative measurement and display – early experience. Radiology 2005;237:1048–55. 6. Thomsen C, Becker U, Winkler K, Christoffersen P, Jensen M, Henriksen O. Quantification of liver fat using magnetic resonance spectroscopy. Magn Reson Imag 1994;12:487–95. 7. Longo R, Ricci C, Masutti F, Vidimari R, Croce LS, Bercich L, et al. Fatty infiltration of the liver: quantification by 1H localized magnetic resonance spectroscopy and comparison with computed tomography. Invest Radiol 1993;28:297–302. 8. Joy D, Thava VR, Scott BB. Diagnosis of fatty liver disease: is biopsy necessary? Eur J Gastro Hepatol 2003;15:539–43. 9. Boden G, Sargrad, K, Homko C, Mozzoli M, Peter Stein T. Effect of a low-carbohydrate diet on appetite, blood glucose levels, and insulin resistance in obese patients with Type 2 diabetes. Ann Intern Med 2005;142:403–11. 10. Yancy WS, Olsen, MK, Guyton JR, Bakst RP, Westman EC. A low-carbohydrate, ketogenic diet versus a low-fat diet to treat obesity and hyperlipidemia. Ann Int Med 2004;140:769–77. 11. Lomas DJ, Black RT, Pinney J. Hepatic steatosis quantification by MRI: serial measurement and normal variation. Proc Intl Soc Magn Reson Med 2003;11:1424. 12. Westerbacka J, Lammi K, Hakkinen AM, Rissanen A, Salminen I, Aro A, et al. Dietary fat content modifies liver fat in overweight nondiabetic subjects. J Clin Endocrin Metab 2005;90:2804–9. 13. Dansinger ML, Gleason JA, Griffith JL, Selker HP, Schaefer EJ. Comparison of the Atkins, Ornish, Weight Watchers, and Zone diets for weight loss and heart disease risk reduction: a randomized trial. JAMA 2005;293:43–53. 14. Petersen KF, West AB, Reuben A, Rothman DL, Schulman GI. Non-invasive assessment of hepatic triglyceride content in humans with carbon-13 nuclear magnetic resonance spectroscopy. Hepatology 1996;24:114–7.

715


COMMENTARY

Translational research in radiotherapy trials 1

C WEST,

BA, PhD

and 2S McKEOWN,

MA, PhD

1

Academic Department of Radiation Oncology, The University of Manchester, Christie Hospital, Manchester M20 4BX and 2School of Biomedical Sciences, University of Ulster at Coleraine BT52 1SA, Northern Ireland, UK

Received 7 April 2006 Accepted 13 April 2006 DOI: 10.1259/bjr/62971945 ’ 2006 The British Institute of Radiology

In the last 10 years, a range of powerful laboratory techniques have become available to measure genes at the DNA, RNA and protein level. This progress in assay methods is associated with a huge increase in the amount of data produced and the development of computational methods for analysing the biological information. These advances have the potential to provide the means of assigning molecular signatures that describe intrinsic differences in how patients respond to radiotherapy. Currently, the methods are being validated as tools that provide prognostic or predictive information on cancer pre-disposition and treatment outcome on an individual basis. In order to raise awareness of recent developments in the area, the British Institute of Radiology’s (BIR’s) Radiation and Cancer Biology Committee organized a meeting at the BIR on 10 March 2006 to overview the methods and to discuss potential applications for advancing translational research in radiotherapy trials. The greatly improved opportunities for translational research in radiotherapy trials in the UK were summarized by Professor Peter Hoskin (Mount Vernon Hospital, Northwood). In order to improve the design and quality of trials involving radiotherapy within the UK, trials are now reviewed by the NCRI Radiotherapy Clinical Studies Group. The central review encourages multicentre involvement to facilitate patient accrual. It also allows the identification of potential translational research opportunities at an early stage. There are currently 44 radiotherapy trials open in the UK with more planned, giving widening opportunities for translational questions to be addressed. High throughput genotyping has the potential to find common genetic variants reported as single nucleotide polymorphisms (SNPs) conferring increased radiosensitivity (Dr Alison Dunning, University of Cambridge). In terms of application in clinical trials, one of the key advantages of genotyping is the relative ease of sample collection. A single blood sample taken in EDTA is required, with some flexibility of the timing of sample 716

collection. An important caveat being that DNA quality and uniformity matters. Although samples are relatively stable compared with those collected for RNA or proteomic analyses, there can be problems with collections involving DNA extracted from mixed types of samples, which can yield unusable data. The HapMap (SNP haplotype map) project has already published its phase 2 results and most of the .2 million documented SNPs in the human genome are now tagged and can therefore be studied. The good news is the technology involved for analysing SNPs has progressed in recent years and high-throughput methods incorporating robotics enable the simultaneous processing of many thousands of samples within a single day. SNP-tagging methods, for reducing the number of SNPs that need to be investigated, are also developing rapidly. Studies can focus on candidate genes likely to be involved in determining an individual’s sensitivity to radiation, e.g. genes involved in the recognition and repair of DNA damage. Although each gene of interest might contain several hundred SNPs, the number needed to be studied directly can be reduced by using tagged SNPs. ‘‘Tagging’’ involves identifying regions of the genome (haplotype blocks) where SNPs have evolved together and then finding the SNPs within each block that provide information on most of the other SNPs. Nevertheless, even with candidate genes and SNP tagging, there is an obvious need to collaborate and pool samples obtained in several trials to enable exploration of the full human genome. Since there are many variants, each with probably only a small individual effect, large samples (.2000 patients) are required to have any statistical power to detect real effects due to the variants – the field of association studies has been plagued with false positive results due to small study sizes. For radiotherapy, high-throughput genotyping offers the opportunity to identify a range of radiosensitivity phenotypes, which in the future should allow more accurate tailoring of treatment protocols to a patient’s predicted normal tissue radiosensitivity. This The British Journal of Radiology, September 2006

Commentary: Translational research in radiotherapy trials

could improve treatment planning for all groups of patients from the most radiosensitive to radioresistant. The potential of cDNA and oligonucleotide microarrays was described by Dr Wendy Allen (Queen’s University Belfast). The techniques are now well advanced with evaluation of 60 000 genes per sample possible. The methods are capable of comparing the expressed genes in tumour samples with patient outcome following radiotherapy. A disadvantage of the technology is the requirement for fresh material that is not always readily obtainable within a multicentre trial setting. Methods are now becoming available, however, to extract good quality RNA from wax embedded material which will increase the applicability of the approach, since this type of material is easier to store and much more widely available. Microarrays yield huge amounts of information, and so the bioinformatic handling of the data is crucial for identifying potentially important genes and molecular signatures. Dr Allen emphasised the need for validation studies to verify molecular profiles identified in a training set with a second blinded set of samples. The potential applications of the technique in identifying predictive/prognostic gene signatures was also discussed. Taking the step from expressed gene to translated protein, Professor Tony Whetton (The University of Manchester) described the current state of the proteomics revolution. Although the technology is continually improving, the capabilities are several years behind the high-throughput DNA and RNA approaches. The techniques required are technically demanding, combining highly sophisticated protein separation methods (two-dimensional liquid chromatography) with mass spectrometry. This is partly improved by new tagging techniques and other approaches which allow relative quantification from several samples as well as protein/ peptide identification. One of the biggest problems, however, is the dynamic range issue. The separation of low abundance proteins of interest from high abundance proteins remains critical, although major progress has been made in this area with major protein removal columns. The preliminary removal of these proteins makes analysis of low abundance, potentially diagnostic proteins much more likely. As gene expression array and proteomic data can be linked, there is also potential to compare transcribed and translated genes in the same samples. Procedures for blood sampling for proteomic studies must be carefully standardized, and the use of proteomic techniques for tissue samples remains problematic and is probably 5 years from being a realistic proposition. Currently, the development of large-scale proteomic studies in radiotherapy trials is not realistic. However, exploration of proteomic techniques in smallscale studies would be of value. Dr Tim Helliwell from Liverpool University discussed the importance of tissue microarrays (TMAs) and their ability to increase evaluation of multiple samples. With TMAs, several hundred samples can be examined on a single slide. There are several advantages of this approach, although it must be seen as a population screening tool and should not be used for individual patient diagnosis. TMAs allow the relative frequency of a target molecule to be assessed in relation to outcome. It The British Journal of Radiology, September 2006

can be used for visualizing proteins (using specific antibodies) or gene amplification/expression (using fluorescent in situ hybridization). The approach is limited by a potential sampling error (which can be minimized by increasing the number of cores per sample), the ability to select the correct probes and also probe availability. However, with more probes becoming available, there is much than can now be done. Again, with the great increase in information there is a need for bioinformatic analysis of the results. TMAs have great potential for obtaining molecular markers associated with prognosis in past, current and future trails. Dr Francesca Buffa (Gray Cancer Institute, Northwood) highlighted the problem of processing the information that can be produced using the new techniques. The analysis of a large number of genes/ proteins and a relatively small number of patients poses several problems regarding variable selection, accuracy of prediction on future data and interpretation of the derived model. Analyses must account for multiple testing and false positives are likely. There are different approaches for handling the data. Supervised methods classify according to a pre-selected clinical end-point – such as locoregional control or disease-specific survival – or biological function. Unsupervised techniques do not require a pre-determined outcome or biological knowledge but cluster genes and/or samples that are similar. A knowledge-based approach can also be used, based on our understanding of biology to look at the expression of genes related to a particular phenotype, such as hypoxia. She discussed the advantages and disadvantages of these approaches from a statistical point of view and raised issues that needed to be considered when designing a trial so that data analysis can be more informative; such as reproducibility and variability of the data collected and standardization of laboratory procedures between centres. Dr Rob Bristow (Princess Margaret Hospital in Toronto, Canada) emphasised the need for a focused tumour type specific approach to the development of prognostic and predictive biomarker profiles. He described his work on prostate cancer and the evaluation of markers related to radiobiologically relevant phenotypes: hypoxia and activated DNA repair processes. A considerable amount of work is clearly required in evaluating markers in different tumours and those useful in one disease site might have no relevance in another. The potential was highlighted for using phase I or II trial data to develop and explore mechanistic hypothesis applicable to later phase trials, and to establish appropriate times for sampling. There is also a need to evaluate markers not just in clinical trial material but also in relation to the carcinogenic process of different tumours. As some markers might prove to be targets for novel drug development, there is a need to show their expression in tumours compared with surrounding normal tissue. The importance of being aware of any underlying clinical heterogeneity in biomarker studies was also highlighted (Mr Priy Silva, The University of Manchester). Cancers of the head and neck comprise a heterogeneous group. The multiple sub-sites involved are associated with differences in radioresponse. A study in head and neck squamous cell carcinoma from patients 717

C West and S McKeown

who underwent radical radiotherapy highlighted the potential confounding influence of underlying clinical heterogeneity. Even within a very clinically homogeneous, single sub-site group – oropharyngeal patients who all received radiotherapy to their primary tumour – there was variation in outcome probability in relation to clinical and biomarker data. Differences were shown in the behaviour of tonsil and base of tongue tumours. For example, one marker provided highly significant prognostic information in tonsil tumours and another in base of tongue. Evaluation of potential biomarkers of response must allow for any underlying clinical heterogeneity. Where possible, the powering of studies to enable meaningful subgroup analyses would be useful. The quality of samples and reproducibility of data obtained in both single and multicentre studies is an important consideration for the design of translational research in clinical trials. The need to establish detailed protocols for sample collection particularly in multicentre studies was emphasised (Dr Catharine West, The University of Manchester). A two-centre study was described involving the collection of head and neck cancer samples during surgery for RNA profiling. All samples taken in RNA later yielded high quality RNA, with significantly greater variation between than within samples. Using a knowledge-based approach for data

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analysis the derivation of a radiobiologically relevant phenotype was described: a hypoxia metagene, which was shown to yield prognostic information in independent datasets. The meeting provided no obvious quick-fit answers for what endpoints or approaches to use in the planning of translational aspects of clinical trials. Indeed all of the methods have their advantages and disadvantages and, although they can provide informative data, none can provide all of the answers at present. What did emerge from the meeting was a need for a multidisciplinary and collaborative approach. The development of translational research protocols will benefit from inclusion, at the planning stage, of not just the clinical oncologists involved in running a trial, but also scientists, statisticians, bioinformaticians and pathologists. A strong theme for all the various approaches described was the old adage of ‘‘rubbish-in-rubbish-out’’. Prospective sample collection is clearly important, but must be well thought out with translational research requirements discussed at an early stage in the clinical trial development process. The current funding opportunities for clinical trials and translational research should be exploited by the Clinical Oncology and Radiobiology community and, towards this goal, communication and collaboration are essential.



Oesophageal dysmotility in systemic sclerosis: comparison of HRCT and scintigraphy 1

E H PITREZ, MD, 2M BREDEMEIER, MD, MSc, 2R M XAVIER, MD, PhD, 2K G CAPOBIANCO, MD, 2 V G RESTELLI, MD, 1M V VIEIRA, MD, PhD, 3D H C LUDWIG, MD, 2J C T BRENOL, MD, PhD, 1 A P A FURTADO, MD, 4L M B FONSECA, MD, PhD and 4B GUTFILEN, MD, PhD

MSc,

1

Serviço de Radiologia, 2Serviço de Reumatologia and 3Serviço de Medicina Nuclear, Hospital de Clıńicas de Porto Alegre, Faculdade de Medicina, Universidade Federal do Rio Grande do Sul, Rua Ramiro Barcelos 2350, Porto Alegre, RS, Brazil, 90.035-003 and 4Departamento de Radiologia, Hospital Clementino Fraga Filho, Universidade Federal do Rio de Janeiro, Brigadeiro Trompovsky, Ilha do Fundaõ, CEP 21941-590, Rio de Janeiro, RJ, Brazil ABSTRACT. The aim of this study was to compare oesophageal abnormalities observed in high-resolution CT with radionuclide transit in patients with systemic sclerosis. 76 patients with systemic sclerosis were evaluated by high-resolution CT and oesophageal transit scintigraphy. Residual activity >20% (in relation to peak activity) at 15 s after the beginning of the swallow of the labelled liquid (in supine position) was considered indicative of oesophageal dysfunction. Supra-aortic and infra-aortic oesophageal coronal diameters were measured in high-resolution CT. Oesophageal dilatation was deemed present when the diameters exceeded 10 mm. 19 patients (25%) had supraaortic oesophageal dilatation and 48 patients (63.1%) had infra-aortic dilatation. The prevalence of radionuclide transit delay was 77.6%. All patients (19/19) with supraaortic dilatation had oesophageal dysfunction, compared with 70.2% (40/57) of the patients with no supra-aortic dilatation (p 5 0.004). Oesophageal dysfunction was present in 97.9% (47/48) of patients with infra-aortic dilatation, compared with 42.9% (12/28) in patients without it (p , 0.001). Receiver operating characteristic (ROC) curves have demonstrated that the supra-aortic and infra-aortic diameters had good discriminatory capacity for oesophageal dysfunction in systemic sclerosis (area under the curve, 95% confidence interval: 0.80, 0.70–0.89 and 0.92, 0.86–0.98, respectively). There is a clinically significant association between oesophageal dysmotility and highresolution CT findings of oesophageal coronal dilatation. The evaluation of infra-aortic oesophageal coronal diameter can provide additional useful information about the functional and anatomic conditions of the oesophagus in systemic sclerosis.

Systemic sclerosis (SSc) is a connective tissue disease characterized by fibrosis in multiple organs, especially in the skin, lungs and gastrointestinal system [1, 2]. The oesophagus is the most frequently involved internal organ [3–5], and atrophy and fibrosis of the smooth muscle are probably responsible for the impaired oesophageal motility that is common in this disease [2, 5]. Diminished lower oesophageal sphincter pressure and associated gastro-oesophageal reflux (GER) occur frequently, leading to oesophagitis, strictures and ulcerations [4, 6]. Early recognition of the oesophageal dysfunction is important because patients may be asymptomatic for a long time despite the presence of oesophageal motor abnormalities [1, 3, 4]. Manometry has been considered the gold-standard examination to detect oesophageal dysfunction in early stages of SSc, but it gives no information on morphology and has low patient acceptance [1–3]. Radionuclide oesophageal scintigraphy(RES)standsoutasafrequentlyusedalternative to manometry, since it is a safe, non-invasive, and sensitive method that allows quantitative assessment [1, 2, 7–10]. The British Journal of Radiology, September 2006

Received 7 November 2005 Revised 4 January 2006 Accepted 30 January 2006 DOI: 10.1259/bjr/17000205 ’ 2006 The British Institute of Radiology

Thorax high-resolution computed tomography (HRCT) is the preferred radiological examination in the evaluation of interstitial lung disease in SSc [11]. Oesophageal abnormalities in thorax CT, including dilatation, air–fluid levels, and food retention have already been documented in SSc patients [12]. As far as we are aware, a comparison of the HRCT and RES results in the evaluation of oesophageal dysfunction has not yet been performed. Therefore, our objective in this study was to compare both methods and to test the utility of HRCT in the diagnosis of oesophageal dysfunction in SSc using RES as the gold standard.

Material and methods Patients 76 patients with the diagnosis of SSc (made according to the clinical evaluation of experienced rheumatologists) were prospectively studied. The patients met the 719

E H Pitrez, M Bredemeier, R M Xavier et al

American College of Rheumatology criteria for SSc [13] or the criteria suggested by LeRoy and Medsger for diagnosis of early forms of SSc [14]. Patients with overlapping syndromes, active chronic or acute infections, insulin-dependent diabetes mellitus or with longstanding diabetes (more than 5 years since diagnosis) were excluded. Patients with definite diagnosis of SSc (according to the American College of Rheumatology criteria) who developed inflammatory myopathy were not excluded from the analysis. All patients signed written informed consent before entry in the study. The study was approved by the Research Ethics Committee of the Hospital de Clıńicas de Porto Alegre.

Clinical evaluation All patients were interviewed and examined by a rheumatologist (MB). The interview instrument was an extensive questionnaire directed to the evaluation of end-organ damage. Duration of disease was defined as the period of time between the onset of Raynaud’s phenomenon or skin symptoms (whatever came first) and the moment of the interview [15]. The sample included 67 women and 9 men, and the age range varied from 26 years to 78 years (mean 51 years, standard deviation 11 years). The duration of symptoms ranged from 2–25 years (median 13 years).

Radionuclide oesophageal transit scintigraphy All subjects were examined after 4 h of fasting. The study was performed in supine position with liquid ingestion. The liquid phase was prepared with 6 ml of water labelled with 1 mCi of 99mtechnetium-phytate for each swallow. Imaging was performed with a large fieldof-view gamma camera (high sensitivity parallel-hole collimator, with the camera underneath. The field-ofview extended from the pharynx to the upper part of the stomach. A practice run with unlabelled water preceded the study. The labelled liquid was placed in the mouth and swallowed on command. The computer acquired a dynamic study at a rate of 1 s frame21 for 60 s with a 64 6 64 matrix. No dry swallows were allowed. A region of interest over the whole oesophagus was defined and a time–activity curve was generated. The percentage of emptying of the oesophagus was analysed in 15 s and 60 s after swallow. Residual activity >20% (in relation to peak activity) at 15 s after the beginning of the swallow in the supine liquid phase was considered abnormal and indicative of oesophageal dysfunction [16].

150 mAs (Elscint), 140 kVp and 130 mAs (Siemens). The scanning time was 1.0 s and the examination was performed from the apices of the lungs to the lung bases. Prone scans were obtained whenever subpleural lesions were observed. HRCT was reviewed with lung (window width, 2000 H; window level, 2700 H) and soft-tissue windows (window width, 400 H; window level, 40 H). The thoracic oesophagus was divided in supra-aortic and infra-aortic levels (above and below the superior limit of aortic arch, respectively). The largest coronal diameters of the supra-aortic and infra-aortic levels observed in the HRCT pictures were measured and recorded, considering the internal limits of inner oesophageal mucosa (Figure 1). Oesophageal dilatation was diagnosed if the luminal coronal diameter of the oesophagus exceeded 10 mm, whatever the content (air, liquid or solid) [12]. Intraobserver agreement for the oesophageal coronal diameter measurement in HRCT was tested in 30 examinations. The second reading was performed 12 months apart from the initial evaluation. There was a good agreement between the two observations for supraaortic oesophageal coronal diameter (Fleiss weighted Kappa 5 0.65) and a moderate agreement for infra-aortic oesophageal coronal diameter (Fleiss weighted Kappa 5 0.56). Interobserver agreement for oesophageal coronal diameter measurement in HRCT was tested using another 30 examinations. The measurements obtained by another radiologist demonstrated a good agreement for the supra-aortic (Fleiss weighted Kappa 5 0.64) and for the infra-aortic oesophageal coronal diameters (Fleiss weighted Kappa 5 0.60). All clinical, radiological, and scintigraphic examinations were performed within 6 months. No observer was informed about the clinical details of the SSc patients or about the results of the other examinations.

Statistical analysis Data were analysed using EPI-INFO version 6 and SPSS for Windows version 11.0. Categorical variables were presented as numbers and proportions. Quantitative variables were tested graphically (with

High-resolution CT The HRCT was performed with an Elscint Twin scanner (Elscint, Haifa, Israel) and with a Somatom Plus 4 (Siemens Medical Systems, Iselin, NJ). HRCT scans were obtained in supine position during full inspiration, with the following parameters being used: 1-mm sections at 10-mm intervals, a high-spatialfrequency algorithm, 512 6 512 matrix, 140 kVp, and 720

Figure 1. A schematic drawing of how the coronal oesophageal diameter was measured.


Oesophageal dysmotility in systemic sclerosis

normal probability plots) and statistically (with Kolmogorov-Smirnov goodness-of-fit test) for the normality of distribution. Quantitative variable with normal distribution was presented as mean and standard deviation (SD). Non-normal quantitative variables were presented as median and interquartile range (IQR). The associations between categorical variables were tested using Fisher’s exact test. Paired comparisons involving non-normal quantitative variables were performed with the Wilcoxon signed ranks test. A two-tailed p-value # 0.05 was considered statistically significant. Receiver operating characteristic (ROC) curves were used to test the ability of HRCT to differentiate patients with and without oesophageal radionuclide transit delay. The area under the ROC curve is a suitable measure to summarize the discrimination power of a diagnostic model (representing the accuracy of the model) and can range from 0.5 (no discrimination) to 1.0 (perfect discrimination). 95% confidence intervals for the areas under the curves were also calculated. Sensitivity and specificity values were estimated, along with 95% confidence intervals (95% CI).

Results 59 out of 76 patients (77.6%) had scintigraphic evidence of oesophageal dysfunction. The prevalence of 20% or more retention of liquid in supine position at 60 s (75%) was similar to the evaluation at 15 s. The thorax HRCT demonstrated ground-glass opacities or reticular pattern/honeycombing in 43 patients (56.6% of the sample). Out of these patients, 24 presented ground-glass opacities and reticular pattern/honeycombing, 6 presented reticular pattern/honeycombing without ground-glass opacities, and 13 presented only ground-glass opacities. Supra-aortic oesophageal dilatation (Figure 2a) was present in 19 patients (25.0%). Infra-aortic oesophageal dilatation (Figure 2b) was detected in 48 patients (63.1%). Supra-aortic and infra-aortic oesophageal dilatations in HRCT were significantly associated with oesophageal

dysmotility (Tables 1 and 2, respectively). All oesophageal wall calibres were smaller than 3 mm (Figure 3). As the selection of cut-off values of coronal diameters is rather arbitrary, we additionally used ROC curves to test the ability of HRCT to discriminate patients with and without oesophageal dysfunction on scintigraphy. The areas under the ROC curves suggest that the oesophageal coronal diameters (particularly the infra-aortic) have good discriminatory ability for oesophageal dysmotility (Figure 4). An infra-aortic oesophageal diameter >9 mm would provide a sensitivity of 83.1% (95% CI: 71.0–91.5) and a specificity of 94.1% (95% CI: 71.2–99.0) for oesophageal dysmotility. Supra-aortic oesophageal diameter >4 mm would provide a sensitivity of 64.4% (95% CI: 50.9–76.4) and a specificity 94.1% (95% CI: 71.2–99.0) for oesophageal dysmotility.

Discussion SSc affects the skin in the form of thickening, tightness, pitting oedema, digital ulcers and telangiectasias, but also affects visceral organs like the lungs, the kidneys and the gastrointestinal tract. Although skin changes are the most remarkable abnormalities, the visceral disease (especially pulmonary) is responsible for the prognosis [12]. The gastrointestinal tract is involved in up to 80–90% of SSc patients [12, 17–19], with the oesophagus being the internal organ most frequently involved (50–80%) in this disease [5, 6]. SSc leads to atrophy of smooth muscle, to impaired peristalsis, and to reduction of the lower oesophageal sphincter (LES) tone. Reduction in LES tone leads to gastro-oesophageal reflux and oesophagitis [5–7, 17]. Symptoms include dysphagia, odynophagia, heartburn and regurgitation [5, 7, 12, 17], but the absence of symptoms does not exclude advanced affliction of the oesophagus [6, 19–21]. The lower two-thirds of the oesophagus are usually affected, commonly appearing dilated, with absent peristalsis, and sometimes with strictures and mucosal changes due to oesophagitis [4–7, 12].

Figure 2. A 63-year-old man. High resolution CT demonstrates (a) supra-aortic oesophageal dilatation and (b) infra-aortic dilatation (black arrows). Note peripheral ground-glass opacities (white arrowheads). The British Journal of Radiology, September 2006

721

E H Pitrez, M Bredemeier, R M Xavier et al Table 1. Association between supra-aortic coronal oesopha-

Table 2. Association between infra-aortic coronal oesopha-

geal diameter and oesophageal dysmotility assessed by radionuclide oesophageal scintigraphy*

geal diameter and oesophageal dysmotility assessed by radionuclide oesophageal scintigraphy

Supra-aortic coronal oesophageal diameter

Oesophageal dysmotility

Yes

No

n (%)*

n (%)*

Infra-aortic coronal oesophageal diameter

>10 mm ,10 mm

19 (100) 40 (70.2)

0 (0) 17 (29.8)

> 10 mm , 10 mm

Oesophageal dysmotility

p**

0.004

Yes

No

n (%)*

n (%)*

47 (97.9) 12 (42.9)

1 (2.1) 16 (57.1)

p**

,0.001

*The percent values were calculated for the rows. ** Fisher’s exact test.

*The percent values were calculated for the rows. **Fisher’s exact test.

The sclerodermic oesophagus has been traditionally evaluated by different diagnostic methods, including manometry [1, 4, 22, 23], endoscopy [5], scintigraphy [1, 3, 4, 7, 9, 10, 16, 24–27], barium-oesophagogram [5, 19, 28], cine-oesophagography [23, 29], ultrasound [29] and oesophageal pH monitoring [30]. Since the introduction of the scintigraphic method in the oesophageal evaluation by Kazem, in 1972 [25], several authors have demonstrated that the radionuclide oesophageal transit examination is a safe and non-invasive method that could be used as an alternative to oesophageal manometry in SSc patients [1, 3, 4, 24, 26]. Oesophageal transit scintigraphy has been demonstrated to be sensitive (91–100%) and specific (88%) to detect motor function abnormalities diagnosed by manometry [1, 26, 31]. Although scintigraphy has been considered a good method for the evaluation of oesophageal motility disorders, there is not a uniformly accepted protocol to guide the execution and the interpretation of the examination. This compromises the applicability and reproducibility of the results obtained by different groups of researchers. To illustrate the variability in the methods used, Akesson et al [3] applied a single swallow protocol with the ingestion of a pineapple puree in sitting position, and oesophageal transit delay was defined as a transit time longer than 300 s. Carette et al [26] used a single swallow of 10 ml of labelled water in supine position, followed by a dry swallow 30 s later. Davidson et al [24] used a supine colloid ingestion protocol, defining RES abnormality as the retention of colloid >5% at 15 s after ingestion. In the present study, in which oesophageal transit was evaluated after a single swallow of labelled liquid in supine position, the prevalence of oesophageal dysfunction was 77.6%. In previous studies, the reported prevalence of oesophageal transit abnormalities on scintigraphy varied from 82% to 100% [1, 3, 7, 24, 26] what demonstrates a relative homogeneity of results despite methodological variations. In the present report we used an evaluation of oesophageal transit after a single swallow. Previous studies have suggested that multiple swallows may be more appropriate in the evaluation of oesophageal dysfunction, since the variability observed in single swallow studies may be reduced [16]. However, as far as we are concerned, the diagnostic accuracy of single and multiple swallows (in relation to manometry) has not been directly compared in previous studies. Ham et al [32] suggested that the oesophageal transit patterns observed after multiple swallowing were different from those observed after a single swallow. Tolin et al [8] demonstrated that the oesophageal emptying may be

incomplete in individuals with normal manometry after multiple swallows. Therefore, a definite answer about the ideal protocol to be used in RES has not yet been found. The lungs are the second most commonly affected internal organ in scleroderma. The patients often have diffuse interstitial lung disease (fibrosing alveolitis), which is clinically and prognostically relevant. Thorax HRCT is the established non-invasive gold-standard technique for the diagnosis of fibrosing alveolitis [11]. HRCT abnormalities in scleroderma lung disease are well documented: ground-glass opacification in isolation or with reticular patterns are associated with inflammatory alveolitis and a reticular pattern – with or without honeycomb changes – is associated with a fibrotic histology [11, 12]. Further than diagnosis, patients with scleroderma are also evaluated with HRCT to assess the activity and progression of pulmonary disease. Thorax CT has been previously used to evaluate the oesophagus in SSc patients. Bhalla et al [12] demonstrated oesophageal asymptomatic dilatation (12– 40 mm) in 20/25 patients (80%) with SSc. The authors suggested that oesophageal evaluation by CT could be used to narrow the differential diagnosis of patients with interstitial lung disease. In the present study, infra-aortic and supra-aortic dilatation (>10 mm) were observed in 63.1% and 25% of the cases, respectively. The

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Figure 3. A 55-year-old woman. Infra-aortic slice demonstrates patulous thin-wall (white arrow) oesophageal dilatation. The British Journal of Radiology, September 2006

Oesophageal dysmotility in systemic sclerosis

Figure 4. Receiver operating characteristic curves of (a) supra-aortic and (b) infra-aortic coronal diameters for oesophageal dysfunction detected in scintigraphy.

predominance of dilatation in the infra-aortic region (lower two-thirds) has been largely demonstrated [3, 6, 7, 12] and is related to the greater amount of smooth muscle fibres in the inferior portions of the oesophagus. These fibres are affected by extensive fibrosis and atrophy [7], leading to impaired peristalsis, dilatation, pouches of air–liquid, and even to solid contents in the oesophagus (Figure 5). There are a few studies reporting supra-aortic oesophageal dilatation in SSc patients [21, 33]. Proximal oesophageal dilatation can be explained examining the anatomic variability of the distribution of the smooth musculature in the oesophageal wall. Post-mortem studies have shown that the level of the transition zone between striated and smooth muscle is quite variable, sometimes occurring as proximally as the thoracic inlet or even in the cervical oesophagus [33]. Therefore, patients with a higher transition zone would probably have a greater likelihood of presenting supra-aortic dilatation. As far as we are aware, there have been no previous studies trying to relate the coronal oesophageal diameters with the oesophageal dysfunction in SSc. Our study demonstrated that oesophageal dilatation (particularly at the infra-aortic level) is highly associated to radionuclide transit delay. Additionally, our results also suggest that the infra-aortic coronal diameter may be

useful in the diagnosis of oesophageal dysmotility, as suggested by the ROC curve analysis. Therefore, besides the routine evaluation of interstitial lung disease on HRCT, the measurement of coronal oesophageal diameters could provide useful and easily attainable information about the oesophageal motility status in SSc. In the present study we had no control group or patients with other causes of oesophageal dysfunction to compare with SSc patients. Therefore, our results apply exclusively to patients with SSc. It would be interesting to compare HRCT and RES in other diseases characterized by oesophageal dysmotility or structural abnormalities. In no patient of this study the oesophageal wall thickness exceeded the reported normal value of 3 mm. This has been also described in a previous report [34] and probably reflects that, although histopathological oesophageal abnormalities (fibrosis and atrophy) may be striking, these findings do not correlate with oesophageal wall thickening, not even in patients with long lasting disease. Considering the results presented here, we conclude that the measurement of oesophageal coronal diameters in thorax HRCT (directed primarily to the evaluation of interstitial lung disease) provide useful information about the functional status of the oesophagus in SSc patients.

Acknowledgments We acknowledge Drs Raquel Faccioni, Marcus Franck, Tatiana Freitas Tourinho, Marcelo Maltchick, Paulo Se´rgio Thys, Patrıćia Minuzzi da Motta, Max Brenner, Tamara Mucenic, Adriano Barbiero, Tatiana Karenini Mu¨ ller, Carmen Both Schenatto, Charles Lubianca Kohem, Lilian Scussel-Lonzetti, Claiton Viegas Brenol, Sandra Helena Machado, and Iloíte Scheibel on their valuable support. We also thank to Juliana Bredemeier for the English review of this manuscript.

References Figure 5. A 61-year-old woman. Dilated infra-aortic oesophagus after overnight fasting demonstrates ingested contents (black arrow).


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Relationship between the growth pattern of nasopharyngeal cancer and the cervical lymph nodes based on MRI findings: can the cervical radiation field be reduced in patients with nasopharyngeal cancer? 1

N FUWA, MD, PhD, 2Y ARIJI, DDS, PhD, 3T DAIMON, PhD, 4M WAKISAKA, MD, PhD, 4A MATSUMOTO, MD, 1 T KODAIRA, MD, PhD, 1H TACHIBANA, MD, 1T NAKAMUA, MD, PhD and 1Y SATOU, MD, PhD 1

Department of Radiation Oncology, Aichi Cancer Center Hospital, 1-1 Kanokoden, Chikusaku, Nagoya 464-8681, 2Department of Oral and Maxillofacial Radiology, Aichi-Gakuin University School of Dentistry, 2-11 Suemori-dori, Chikusaku, Nagoya 464-8651, 3Department of Clinical Research and Management, Translational Research Information Center, Foundation for Biomedical Research and Innovation, 1-5-4 Minatojima-nakamachi, Chuo-ku, Kobe 650-0047 and 4Department of Radiology, Oita Medical University, 1-1 Idaigaoka, Hasama-machi, Oita 879-5593, Japan ABSTRACT. To identify patients with nasopharyngeal cancer in whom the cervical radiation field can be reduced, we classified the growth patterns of nasopharyngeal cancer based on MRI findings into 4 types and performed an evaluation. Based on MRI findings, we classified the growth patterns of primary cancer in 94 patients with nasopharyngeal cancer into Type 1 (superficial type), Type 2 (lateral invasive type), Type 3 (upward invasive type), and Type 4 (anterior extension type), and further classified Type 2, based upon nasopharyngoscopic findings, into Type 2a (unilateral invasive type) and Type 2b (bilateral invasive type). The cervical lymph node metastasis areas were evaluated according to these types. Type 2 showed a significantly higher incidence of cervical lymph node metastasis only on the ipsilateral side than the other types (p50.0024). In particular, all patients with Type 2a had cervical lymph node metastasis only on the ipsilateral side (p50.0212). This study suggests that the distribution of metastasised cervical lymph nodes depends on the pattern of tumour extent of the primary site.

Cervical lymph node metastasis often occurs in patients with nasopharyngeal cancer (NPC). It has been reported that 60–87% of patients demonstrated cervical lymphadenopathy at the time of diagnosis [1–5]. Previous articles [6–8] have attributed bilateral cervical lymphadenopathy to the abundance of lymphatic tissues in the posterior wall of the nasopharynx and the presence of abundant lymphatic anastomoses crossing the midline. Therefore, bilateral cervical lymph nodes have always been included in the target volume on radical radiotherapy (RT) [6, 7]. As a result, the radiation fields were wide and complications such as the xerostomia secondary to parotid irradiation have been experienced [9–14]. Recently, MRI has been used to evaluate NPC [15–22]. Multiplanar imaging can show tumour extension in all planes, and there is improved tumour delineation and identification of the spread of the tumour to adjacent areas as a result of the better soft tissue contrast in comparison with CT. We previously evaluated the relationship between tumour growth patterns and lymph node metastasis based on MRI images in 32 patients with nasopharyngeal cancer. We showed a high incidence of cervical lymph The British Journal of Radiology, September 2006

Received 7 March 2005 Revised 22 March 2006 Accepted 27 March 2006 DOI: 10.1259/bjr/27870658 ’ 2006 The British Institute of Radiology

node metastasis only on the ipsilateral side in patients with tumour invasion only to the lateral nasopharynx (lateral invasive type), and the cervical radiation field can be reduced in such patients [23]. In this study, to identify patients in whom the cervical radiation field can be reduced, the relationship between the growth pattern of nasopharyngeal cancer and lymph node metastasis on MR images was evaluated in detail in 94 patients, obtained by adding 62 patients to the above 32 patients.

Methods and materials Patients Between April 1990 and August 2004, 94 patients (70 males and 24 females) were retrospectively reviewed. The patient age ranged from 14 years to 80 years (median 51 years). Histology revealed that 17 cases of nasopharyngeal cancer were WHO type I, 15 were type II, and 62 were type III. According to the 1997 TNM classification, the tumour stage was classified as stage I in 3 patients, stage II A in 1 patient, stage II B in 31 patients, stage III in 725

N Fuwa, Y Ariji, T Daimon et al

22 patients, stage IV A in 11 patients, stage IV B in 20 patients and stage IV C in 6 patients. As a basic treatment method, patients aged # 70 years were treated with alternating chemoradiotherapy, in which chemotherapy (CDDP, 5FU) is alternated with radiotherapy [24], and those aged > 71 years were treated by radiotherapy alone.

Evaluation items The evaluation items were the possible relationships between the growth pattern as well as the size of nasopharyngeal tumour and cervical lymph node metastasis, between the degree of histological differentiation and cervical lymph node metastasis, and between the tumour growth pattern and the degree of histological differentiation. Differences were analysed by the log-linear model without interaction terms using the contingency table data [25].

Figure 1. A 31-year-old woman presenting with Type 1 (superficial type) spread. Axial T1 weighted MR image, GdDTPA enhanced, shows no abnormal findings in the nasopharyngeal mucosal space.

MRI techniques MR studies were performed with a 1.5 T unit (Signa; General Electric Medical Systems, Milwaukee, WI). Images were obtained with 5 mm thick contiguous sections in two or three planes (axial, coronal, sagittal) depending on the extent of the tumour. All patients underwent both plain and contrast-enhanced MRI. A spin-echo (SE) multisection imaging technique was used in all examinations. A SE 600/25 image and a SE 2000/80 image were considered to be T1 weighted and T2 weighted, respectively.

Classifications of MR image To determine the pattern of tumour extent at the primary site and existence of metastasised cervical lymph nodes, MR images were evaluated by three experienced radiologists specializing in head and neck cancers working together as a team. When they did not reach consensus after an initial reading regarding the primary tumour and the cervical lymphadenopathy, the final decision was made by majority rule. The 94 patients were classified into 4 types with 2 subtypes (Figures 1–4) according to the tumour extent at the primary site demonstrated on MR images before the treatment. Type 1 (superficial type) tumours were limited to the nasopharyngeal mucosal space. Tumour locations could not be judged on MRI, but were assessed by nasopharyngoscopy from the appearance of slightly irregular surface of nasopharyngeal mucosa. Type 2 (lateral invasive type) tumours could be detected on the lateral pharyngeal wall with involvement of the parapharyngeal space laterally. This invasive type was divided into two subtypes. Type 2a (unilateral invasive type) tumours were those that remained on either side of the midline by MRI and did not extend the midline of the posterior wall of nasopharynx assessed by nasopharyngoscopy. Type 2b (bilateral invasive type) tumours were those that extended a cross the midline by MRI 726

Figure 2. A 64-year-old man presenting with Type 2a (unilateral invasive type) spread. Axial T1 weighted MR image, Gd-DTPA enhanced, shows a right nasopharyngeal tumour infiltrating the right parapharyngeal space.

and/or nasopharyngoscopy. Type 3 (upward invasive type) were tumours that mainly invaded cranially toward the skull base. Type 4 (anterior extension type) tumours extended anteriorly toward the nasal cavity, but did not invade the adjacent normal structures. Tumours that could not be classified into the above 4 types were considered to be unclassifiable.

Classification of tumour size All 94 patients were classified into three groups according to the maximal tumour diameter of axial imaging: small (S#2 cm), medium (2 cm,M,4 cm) and large (4 cm #L) tumours. Type 1 tumours that cannot be measured on MRI images were classified as small. The British Journal of Radiology, September 2006

MR analysis of nasopharyngeal cancer

previous study [26]. In this study, in addition to lymph nodes fulfilling this criterion, the following lymph nodes were also regarded as metastasis even when the minimal axis diameter was # 10 mm: necrotic lymph nodes with a visualized capsule and lymph nodes showing a definite decrease in size on MR images after treatment.

Results

Figure 3. A 35-year-old man presenting with Type 2b (bilateral invasive type) spread. Axial T1 weighted MR image, Gd-DTPA enhanced, shows a left a nasopharyngeal tumour (arrowheads) invading across the midline to the opposite side.

Classification of pathology The 94 patients were divided into two groups according to the degree of tumour differentiation, those with lymphoepithelioma, undifferentiated carcinoma and poorly differentiated squamous cell carcinoma as Group 1, and those with moderately and well differentiated squamous cell carcinoma as Group 2.

Definition of metastatic lymph nodes Metastatic cervical lymph nodes were defined as nodes with a minimal axial diameter of 10 mm in a

(a)

Of the 94 patients, 19 (20%) had type 1, 12 (13%) had type 2a, 18 (19%) had type 2b, 14 (15%) had type 3, 29 (31%) had type 4 and 2 (2%) were unclassified. 82 (87%) out of 94 patients demonstrated unilateral and/or bilateral cervical lymphadenopathy. 41 (50%) out of 82 patients presented bilateral cervical and/or contralateral cervical lymphadenopathy. The relationship between the pattern of tumour extent at the primary site and the cervical lymph node metastasis is shown in Table 1. The incidence of ipsilateral cervical lymph node metastasis was significantly higher for Type 2 than for the other types (p50.0024). In particular, all patients with Type 2a had cervical lymph node metastasis only on the ipsilateral side (p50.0148). The incidence of bilateral cervical lymph node metastases was high for Types 1 (p50.0393) and 4 (p50.0482). Table 2 shows the size of nasopharyngeal tumour and areas of cervical lymph node metastasis. No differences were observed in areas of cervical lymph node metastasis among these groups, but the incidence of lymph node metastasis was lower in large tumour groups (p50.0435). Table 3 shows areas and incidences of cervical lymph node metastasis according to the degrees of histological differentiation. No differences were observed in areas of cervical lymph node metastasis between Group 1 (differentiated type) and Group 2 (poorly differentiated type), but the incidence of lymph node metastasis was lower in Group 1 (p50.0008).

(b)

Figure 4. (a) A 60-year-old man presenting with Type 3 (upward invasive type) spread. Axial T1 weighted MR image, Gd-DTPA enhanced, shows a nasopharyngeal tumour (arrowheads) which is invading toward the left skull. (b) Coronal T1 weighted MR image, Gd-DTPA enhanced, shows a nasopharyngeal tumour (arrowheads) which is invading mainly toward the left skull base.


727

N Fuwa, Y Ariji, T Daimon et al Table 1. Relationship between the pattern of tumour extent at the primary site and areas of cervical lymph node metastasis

No lymph node metastasis (N0) Ipsilateral lymph node metastasis Bilateral and/or contralateral lymph node metastasis

Type 1 n519

Type 2a n512

Type 2b n518

Type 3 n514

Type 4 n529

2 5 12

1 11 0

2 11 5

3 5 6 (1)

3 9 17

Unclassified type (2 cases) was excluded from Table 1. ( ): Number of contralateral neck lymph node metastasis.

Table 2. Relationship between the size of nasopharyngeal tumour and areas of cervical lymph node metastasis

Table 3. Relationship between the degrees of histological

Small n519 (S#2 cm)

Group 1 n518

No lymph node metastasis Ipsilateral lymph node metastasis Bilateral and/or contralateral lymph node metastasis

Middle n541 (2 cm, M,4 cm)

Large n534 (4 cm#L)

2

3

7(1)

5

22

14

12

16

13

( ): Number of contralateral neck lymph node metastasis.

Table 4 shows the degree of histological differentiation and the growth pattern of nasopharyngeal tumour. No association was observed between Type 1 or 2 and the degree of histological differentiation. However, patients with Type 3 (upward invasive type) were frequently included in Group 1 (p50.0034), and all patients with Type 4 tumours were included in Group 2 (p50.0383).

Discussion In the 32 patients in our previous study, tumour growth patterns were classified into 3 major types (Type 1, superficial type: Type 2, invasive type; and Type 3, anterior extension type), and Type 2 was further classified into 3 subtypes (Type 2a, unilateral invasive type; Type 2b, bilateral invasive type; and Type 2c, upward invasive type) [23]. In the previous study, the lymph node metastasis pattern was similar between Types 2a and 2b, but Type 2c showed a different pattern. Therefore, in this study, Type 2 was defined as the lateral invasive type and classified into Type 2a (unilateral invasive type) and Type 2b (bilateral invasive type), while Type 2c was separated from Type 2 and changed to Type 3 (upward invasive type). In the previous study, Type 2a was differentiated from Type 2b based only on MRI findings. However, in this study, nasopharyngoscopic findings were also used, and Type 2a was defined as tumour extension not passing the midline of the nasopharynx. As Table 1 shows, the incidence of cervical lymph node metastasis only on the ipsilateral side was significantly higher for Type 2 than for the other types (p50.0024). In particular, all patients with Type 2a showed cervical lymph node metastasis only on the ipsilateral side, which suggested that the cervical radiation field can be reduced for this type. Types 1 728

differentiation and of cervical lymph node metastasis

No lymph node metastasis Ipsilateral lymph node metastasis Bilateral and/or contralateral lymph node metastasis

Group 2 n576

7

5

8

33

3

38 (1)

Group 1: differentiated squamous cell carcinoma. Group 2: undifferentiated carcinoma, poorly differentiated squamous cell carcinoma. ( ): Number of contralateral neck lymph node metastasis.

(superficial type) and 4 (anterior extension type) showed similar lymph node metastasis patterns and higher incidences of bilateral cervical lymph node metastasis than the other types, which suggested that a reduction in the radiation field is difficult for these types. No significant association was observed between tumour size and cervical lymph node metastasis. However, bilateral cervical lymph node metastasis was frequently observed in patients with small tumours (S#2 cm) rather than patients with large tumours (p50.0272). Concerning the degree of histological differentiation and cervical lymph node metastasis, the percentage of patients with no lymph node metastasis (N0) was significantly higher in Group 1 (differentiated type) than in Group 2 (poorly differentiated or undifferentiated type) (p50.0008). In Group 1, N0 was observed in 7 patients, of whom 3 showed Type 3 on MRI images. As Table 4 shows, the patients with Type 3 included 7 patients with the differentiated type. Of the 7 patients, 3 had no cervical lymph node metastasis, suggesting that the incidence of cervical lymph node metastasis is low in patients with Type 3 and the differentiated type, compared with those with the other types. Table 4. Relationship between the pattern of tumour extent at the primary site and the degrees of histological differentiation Type1 n519

Group 1 4 Group 2 15

Type 2a n512

Type 2b n518

Type 3 n514

Type 4 n529

2 10

4 14

7 7

0 29

Group 1: differentiated squamous cell carcinoma. Group 2: undifferentiated carcinoma, poorly differentiated squamous cell carcinoma. Unclassified type (2 cases) was excluded from Table 4.


MR analysis of nasopharyngeal cancer

Figure 5. A 40-year-old woman presenting with Type 4 (anterior extension type) spread. Axial T1 weighted MR image, Gd-DTPA enhanced, shows a nasopharyngeal tumour (arrowheads) extending anteriorly without invasion to adjacent areas.

When the growth pattern of nasopharyngeal cancer was classified and its association with cervical lymph node metastasis was evaluated, the incidence of bilateral cervical lymph node metastasis was high for Types 1 and 4, that of ipsilateral metastasis was high for Type 2, and intermediate findings were obtained for Type 3. The incidence of cervical lymph node metastasis was significantly low in patients with differentiated tumours. These results suggest that the radiation field and dose of the neck for nasopharyngeal cancer can be individualized. The radiation field and dose of the neck may be reduced especially for patients with Type 2a. This hypothesis will be confirmed by prospective study.

References 1. Lindberg R. Distribution of cervical lymph node metastases from squamous cell carcinoma of the upper respiratory and digestive tracts. Cancer 1972;29:1446–9. 2. Mesic JB, Fletcher GH, Goepfert H. Megavoltage irradiation of epithelial tumors of the nasopharynx. Int J Radiat Oncol Biol Phys 1981;7:447–53. 3. Bedwinek JM, Perez CA, Keys DJ. Analysis of failures after definitive irradiation for epidermoid carcinoma of the nasopharynx. Cancer 1980;45:2725–9. 4. Khoury GG, Paterson ICM. Nasopharyngeal carcinoma: a review of cases treated by radiotherapy and chemotherapy. Clin Radiol 1987;38:17–20. 5. Wang CC. Treatment of malignant tumors of the nasopharynx. Otolaryngol Clin North 1980;13:477–81. 6. Fletcher GH, Million RR. Nasopharynx. In: Fletcher GH, editor. Textbook of radiotherapy, 3rd edn. Philadelphia, PA: Lea & Febiger, 1980:364–83. 7. Moss WT. The nasopharynx. In: Moss WT, Cox JD, editors. Radiation oncology, 6th edn. St Louis, MO: The C.V. Mosby Company, 1989:198–214.


8. Perez CA. Nasopharynx. In: Perez CA, Brady LW, editors. Principles and practice of radiation oncology, 2nd edn. Philadelphia, PA: J.B. Lippincott Company, 1992:617–43. 9. Lee AW, Law SC, Ng SH, Chan DK, Poon YF, Foo W, et al. Retrospective analysis of nasopharyngeal carcinoma treated during 1976–1985: late complications following megavoltage irradiation. Br J Radiol 1992;65:918–28. 10. Zimmerman RP, Mark RJ, Tran LM, Juillard GF. Concomitant pilocarpine during head and neck irradiation is associated with decreased posttreatment xerostomia. Int J Radiat Oncol Biol Phys 1997;37:571–5. 11. Tokars RP, Griem ML. Carcinoma of the nasopharynx: an optimization of radiotherapeutic management for tumor control and spinal cord injury. Int J Radiat Oncol Biol Phys 1979;5:1741–8. 12. Hoppe RT, Goffinet DR, Bagshaw MA. Carcinoma of the nasopharynx: eighteen years’ experience with megavoltage radiation therapy. Cancer 1976;37:2605–12. 13. Bedwinek JM, Perez CA, Keys DJ. Analysis of failures after definitive irradiation for epidermoid carcinoma of the nasopharynx. Cancer 1980;45:2725–9. 14. Nishioka T, Shirato H, Arimoto T, Kaneko M, Kitahara T, Oomori K, et al. Reduction of radiation-induced xerostomia in nasopharyngeal carcinoma using CT simulation with laser patient marking and three-field irradiation technique. Int J Radiat Oncol Biol Phys 1997;38:705–12. 15. Teresi LM, Lufkin RB, Hanafee WN. Nasopharynx, oropharynx, and tongue base. In: Stark DD, Bradley WG Jr, editors. Magnetic resonance imaging, 2nd edn. St Louis, MO: Mosby-Year Book, 1992:1135–63. 16. Vogl T, Dresel S, Bilaniuk LT, Grevers G, Kang K, Lissner J. Tumors of the nasopharynx and adjacent areas: MR imaging with Gd-DTPA. AJR Am J Roentgenol 1990;154:585–92. 17. Hudgins PA, Gussack GS. MR imaging in the management of extracranial malignant tumors of the head and neck. AJR Am J Roentgenol 1992;159:161–9. 18. Olmi P, Fallai C, Colagrande S, Giannardi G. Staging and follow-up of nasopharyngeal carcinoma: magnetic resonance imaging versus computerized tomography. Int J Radiat Oncol Biol Phys 1995;32:795–800. 19. Naito Y, Honjo I, Nishimura K, Torizuka K. Magnetic resonance imaging around the eustachian tube. Am J Otolaryngol 1986;7:402–6. 20. Chong VFH, Fan YF. Detection of recurrent nasopharyngeal carcinoma: MR imaging versus CT. Radiology 1997;202:463–70. 21. Som PM. Detection of metastasis in cervical lymph nodes: CT and MR criteria and differential diagnosis. AJR Am J Roentgenol 1992;158:961–9. 22. Ng SH, Chang TC, Ko SF, Yen PS, Wan YL, Tang LM. Nasopharyngeal carcinoma: MRI and CT assessment. Neuroradiology 1997;39:741–6. 23. Wakisaka M, Mori H, Fuwa N, Matsumoto A. MR analysis of nasopharyngeal carcinoma: correlation of the pattern of tumor extent at the primary site with the distribution of metastasized cervical lymph nodes. Preliminary results. Eur Radiol 2000;10:970–7. 24. Fuwa N, Kano M, Toita T, Shikama N, Kodaira T, Matsumoto A, et al. Alternating chemoradiotherapy for nasopharyngeal cancer using cisplatin and 5-fluorouracil - a preliminary report of Phase II Study. Radiother Oncol 2001;61:257–60. 25. Alan A. Categorial data analysis, 2nd edn. New York, NY: Wiley, 2000. 26. van den Brekel MW, Stel HV, Castelijns JA, Nauta JJ, van der Waal I, Valk J, et al. Cervical lymph node metastasis: assessment of radiologic criteria. Radiology 1990;177:379–84.

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Six years experience in intracoronary brachytherapy procedures: patient doses from fluoroscopy 1

´ NDEZ, C PRIETO, BSc, 1,4E VANO, PhD, 1,4J M FERNA 4 1 L GONZALEZ, PhD and D MARTINEZ, BSc

BSc,

2,4

C GALVAN,

MD, PhD,

3

M SABATE,

MD, PhD,

1

Medical Physics Service, 2Radiotherapy Service and 3Interventional Cardiology Service, San Carlos University Hospital, 28040 Madrid and 4Radiology Department, Medicine School, Complutense University, 28040 Madrid, Spain

ABSTRACT. Typical patient dose levels during intracoronary brachytherapy (ICB) procedures using beta sources were determined across a sample of 221 treatments. Dose–area product values, fluoroscopy time and number of frames per procedure, with median values of 62 Gy cm2, 17.0 min and 1493 images, respectively, resulted in a 20% to 50% increase in the values measured for percutaneous transluminal coronary angioplasty procedures in the same medical centre (median values 41 Gy cm2, 14.3 min and 1078 images). Likely reasons for this increase include the additional complexity of ICB, the need for recording and reporting every step of the treatment, getting the essential parameters for the volume determination of the lesion and therapeutic radiation dose calculation and, finally, the learning curve for this kind of procedure. A high concentration skin dose distribution during ICB procedures was measured and in 12% of the patients peak skin doses higher than 1.5 Gy were confirmed. 10 patients were submitted to clinical follow-up and skin injuries were not identified.

Restenosis, or re-narrowing of a coronary artery after dilatation, is the main limitation of percutaneous transluminal coronary angioplasty (PTCA). Stents have substantially reduced the incidence of restenosis, but neo-intimal proliferation within the stent is still a great problem for a significant percentage of patients. Intracoronary brachytherapy (ICB) is indicated in patients with in-stent restenosis [1]. ICB includes delivery of a local radiation dose to the artery wall (target volume) after stenting. There are several potential problems in the safety aspects for the practice of ICB procedures: (a) procedures more complex than standard PTCA procedures in catheterization laboratories; (b) introduction of brachytherapy equipment not previously used, resulting in a new challenge in radiological protection; and (c) coordination among different specialists (cardiologists, radiotherapists and medical physicists working in the catheterization laboratory). These aspects affect safety of both the patient and staff involved. Complex interventional cardiology procedures can produce deterministic effects (skin injuries) due to the high radiation doses imparted to some regions of the patient’s skin [2–4]. PTCA is one of the most frequent interventional procedures in cardiology, and sometimes requires long fluoroscopy times and a large number of cine frames to document the patient’s lesion and the result of the treatment. The radiation field is usually ‘‘concentrated’’ in one or two specific areas of the skin, usually between 60 cm2 and 80 cm2 (depending on the X-ray beam projection, image intensifier field size and collimation carried out by the cardiologist). Because of the high restenosis rate, a certain number of patients 730

Received 18 August 2005 Revised 17 October 2005 Accepted 31 October 2005 DOI: 10.1259/bjr/75766147 ’ 2006 The British Institute of Radiology

require several coronary angiographies and PTCAs, contributing to the increase in skin irradiation, sometimes over intervals of several months or years [5]. ICB procedures exhibit a higher concentration of fields than other interventional cardiology procedures [6], and an increased dose–area product (DAP), fluoroscopy time (FT) and number of frames (NF) [7]. The need for recording and reporting every step of the treatment, the essential parameters for target volume determination, and therapeutic dose calculations [8, 9] are important factors to explain this increment. Therefore, the estimation of the risk of deterministic effects (skin injuries) in ICB procedures is an aspect of radiation protection that should be considered as part of the quality assurance (QA) programme. In this work, the outcomes in patient radiation protection aspects from our 6 years experience with a sample of 221 ICB procedures are shown.

Methods and materials A total of 221 patients underwent ICB procedures between November 2000 and May 2005 with beta sources from a Novoste system (www.novoste.com, Norcross, Georgia, USA) in 73 cases, and a Guidant system (www.guidant.com, Indianapolis, USA) in 148 cases. Intravascular ultrasound (IVUS) was used in a significant number of patients to quantify the lesion and to determine the clinical dosimetry parameters for the brachytherapy procedures. DAP values, together with FT and NF were recorded in 202 of the ICB patients as relevant dosimetric parameters. ICB procedures were The British Journal of Radiology, September 2006

Radiation protection experience in intracoronary brachytherapy

carried out in three dedicated X-ray interventional cardiology rooms equipped with Philips Integris 3000 and 5000 systems (Philips, Best, The Netherlands; all with high filtration fluoroscopy pulsed modes and with the routine use of wedge filters) by one expert interventional cardiologist in collaboration with a radiotherapist and a medical physicist. The X-ray systems were under a QA programme, including periodic constancy checks to evaluate incident air kerma at the entrance of the image intensifier and entrance surface air kerma for different thicknesses of polymethylmethacrylate (PMMA) and copper, following the protocol proposed by the European DIMOND consortium [10]. To evaluate the increase in dose and procedure complexity in ICB, DAP, FT and NF were also recorded for 1707 PTCAs during the same period and using the same X-ray systems. DAP was measured with the built-in calibrated ionization transmission chambers (PTW, Freiburg, Germany). Skin dose distributions were measured using slow film Kodak X-OMAT-V initially, using the procedure previously described [11] and the new EDR2 [12] introduced in 2002 (Kodak, Rochester, NY). For both types of film, thermoluminescent dosimetry (TLD) was used additionally to evaluate some high dose values and for autocalibration of the films. TLD-100 (LiF:Mg,Ti) chips and a Harshaw TLD/Bicron/NE-Technology (BICRON-NE, Solon, OH) reader were used. Digital recording of all the cine images in DICOM (Digital Imaging and Communications in Medicine) format allows, in some complex cases when saturation of the slow film occurred, a complementary analysis of the skin dose distribution using the technical parameters of all the cine series acquired during the procedures [13]. These parameters were retrieved throughout using an updated version of an ad hoc specific software [14]. DICOM header in the cine series recorded by Philips systems allows us to identify the X-ray beam projection (left–right and craniocaudal angulations), image intensifier field size, kilovolts and milliamperes per frame, and distance from the focus to the image intensifier entrance during the different cine series. Methodology for the evaluation of skin dose distribution and peak skin dose (PSD) has been developed by the authors [11, 12]. As a part of the clinical follow-up of patients undergoing ICB procedures, a specific protocol has been developed in the framework of the DIMOND III project [15, 16] to detect possible deterministic effects on the skin of patients whose slow film pattern shows densities corresponding to doses above 1.5 Gy. Clinical follow-up was initiated when PSD exceeded 2 Gy, DAP values exceeded 180 Gy cm2 or whenever it was recommended by other medical circumstances (e.g. previous procedures, special skin radiosensitivity). In these cases, the interventionist should arrange for review of the patient between 10 days and 14 days after the procedure. The use of these triggering levels for the clinical follow-up respond to the likely inability to verify whether the maximum skin dose exceeded the threshold for deterministic effects when slow film is clearly saturated and no TLD chips had been placed in the region of high dose. The purpose of this review was to The British Journal of Radiology, September 2006

identify skin effects. A record of the FT, NF and DAP is part of the documentation of each patient reviewed.

Results Dosimetric aspects of the X-ray procedure Table 1 shows a comparison between DAP, FT and NF in PTCA and ICB procedures in the period 2000–2005. Previous initial data published for the period 2000– 2001 [6, 7] were similar for PTCA, but greater for ICB. The reduction of doses during ICB procedures is shown in Figure 1 for the years 2001–2004. The years 2000 and 2005 have been excluded as few patients were treated in those years.

Patient skin dose Skin dose is generally quite widely distributed in cardiology procedures, but in some procedures a certain area of the skin could receive a higher dose if the visualization of the lesion to be treated requires maintaining a fixed orientation of the X-ray beam. The introduction of a ‘‘concentration factor’’ has been proposed in a previous paper [5] to consider this aspect. The concentration factor has been defined as: CF~

PSD PSD ~ DAP ASD S

where PSD is the ‘‘peak skin dose’’ and ASD is the ‘‘average skin dose’’, obtained as the quotient of DAP and the total irradiated area S (measured from the slow film). Figure 2 shows a mosaic of several selected examples of slow film for PTCAs and for ICB. Note the highest concentration of the irradiation (more density in some areas of the films) for the ICB procedures. Conversely, note the greater number of projections (with different C-arm angulations) for PTCA procedures. DAP greater than 180 Gy cm2 was measured in 13% of all the procedures over the whole study period, although this threshold was surpassed more frequently during the introduction of this kind of procedure (27% of the cases in the period 2000–2001). For 172 patients (78% of patients treated) skin dose distribution was measured. In 12% of the patients, peak skin doses higher than 1.5 Gy were measured (one patient with 4.6 Gy), but only 10 of these patients were effectively reviewed, as some patients belonged to other health areas and some had died. No radiation skin injuries in ICB procedures were found during follow-up examinations.

Discussion It is clear that ICB produces a higher DAP, FT and NF, due to the complexity of the procedure, the need for recording and reporting every step of the treatment and essential parameters for volume determination of the lesion, and therapeutic dose calculation to the target volume [8, 9]. IVUS is advisable [8] to evaluate in detail 731

C Prieto, E Vano, J M Fernańdez et al Table 1. DAP, fluoro time and number of frames for PTCA and ICB procedures between November 2000 and May 2005 Procedure

PTCA ICB

Sample size

DAP (Gy cm2) Median

Min./Max.

3rd quart.

Median

Min./Max.

3rd quart.

Median

Min./Max.

3rd quart.

1707 202

41 62

10/492 11/384

69 99

14.3 17.0

3.1/102.5 6.0/70.0

23.1 23.3

1078 1493

143/4887 443/3444

1536 1937

Fluoro time (min)

Number of frames

DAP, dose–area product; PTCA, percutaneous transluminal coronary angioplasty; ICB, intracoronary brachytherapy.

the lesion and to obtain the basic data for clinical dosimetry of the brachytherapy treatment. However, its use implies an increase in the fluoroscopy time. The control of the right position of the radioactive source in the lesion and the removal of the source also require extra fluoroscopy time and filming series. Sometime, when ‘‘stepping’’ (irradiation of the lesion in several steps because of the extension of the lesion plus appropriate safety margins) is required, the procedure could become even longer. The correct position of the radioactive source in the lesion is one of the basic aspects of the quality of ICB. Sabate et al [17] state that a recognized limitation of endovascular beta-radiation therapy is the development of new stenoses at the edges of the irradiated area. The term ‘‘geographic miss’’ is used to define cases in which the radiation source did not fully cover the injured area. To avoid this problem, fluoroscopy guidance throughout the procedure and additional cine series to confirm the correct position of the radioactive source are required. Images recorded allowed explanation of some unsuccessful treatments as due to geographical miss. The difficulty in performing IVUS in the vessel to be treated, the difficulty in positioning the radioactive source, and the need to use some specific and fixed projections to correctly document the position of the source, increase the ‘‘concentration factor’’ of radiation in some specific skin areas [6]. In these situations, the total DAP could not always be a good indicator of the level of risk for deterministic effects (skin injuries), as previously highlighted [5].

The preliminary European reference level [18] for PTCA is 94 Gy cm2, which is well above our median values for both PTCAs and ICBs (Table 1). The reduction of median DAP for ICB along the period studied may be due to greater experience and confidence by the staff involved, the effect of periodical training in radiation protection highlighting special aspects of ICB [19], and the use of IVUS in fewer patients (73% of patients in 2000–2001 compared with 53% in 2002–2005).

Conclusions ICB leads to an increase in patient dose during the procedure in comparison with PTCA. The patient dose and skin dose increase because of the increase in the complexity of the procedure, the need for recording and reporting every step of the treatment and essential parameters for volume determination and the therapeutic dose calculation. The use of IVUS to quantify the lesion and to determine the clinical dosimetry parameters for the brachytherapy procedures may also involve an increase in patient dose due to the increased fluoroscopy time. Finally, the learning curve with new techniques implied a greater increment with respect to conventional PTCA procedures in the first period of this study. Continuous training in radiation protection highlighting special aspects of ICB together with greater experience and ability in these kinds of procedures also contributed to shortening the difference in patient dose between ICB and PTCA procedures. The extra dose associated with ICB procedures and greater skin dose concentration factor may lead to further skin injury problems. Notwithstanding, in our institution, with the X-ray systems submitted to rigorous quality assurance programmes and with the optimized technical and clinical protocol, this extra dose has not been of special concern for patient skin injuries during the reported period. Nevertheless, in those cases in which a threshold dose is reached, a specific clinical follow-up protocol is advisable. During the study period, satisfactory outcomes and acceptably low radiation doses to the skin of the patients were confirmed.

Acknowledgments

Figure 1. Evolution of median values of dose–area product (DAP) in intracoronary brachytherapies (ICBs) during the period 2001–2004. Detail of the distribution (frequency histogram) of DAP in ICBs. 732

This study was partially funded under the European Commission DIMOND III project (FIGM-CT-200000061), and Coordination Action SENTINEL (FI6R012909). Funding was also provided by the National Program for Scientific Research, Development and Technological Innovation of the Spanish Department The British Journal of Radiology, September 2006

Radiation protection experience in intracoronary brachytherapy

Figure 2. Mosaic of several selected examples of slow films used to measure skin dose distribution for percutaneous transluminal coronary angioplasties (PTCAs) and for intracoronary brachytherapies (ICBs).

for Science and Technology (project BFI2003-09434) and by the Autonomous Community of Madrid (project GR/ SAL/0272/2004).

References 1. Waksman R, Cheneau E, Ajani AE, White L, Pinnow E, Torguson R, et al. Intracoronary radiation therapy improves the clinical and angiographic outcomes of diffuse in-stent restenostic lesions. Results of the Washington radiation for instent restenosis trial for long lesions (Long-Wrist) studies. Circulation 2003;107:1744–9. 2. International Commission on Radiological Protection. Avoidance of radiation injuries from medical interventional procedures. ICRP Publication 85. Ann ICRP 2000;30:7. 3. Koenig TR, Mettler FA, Wagner LK. Skin injuries from fluoroscopically guided procedures: Part 2, Review of 73 cases and recommendations for minimizing dose delivered to patient. AJR Am J Roentgenol 2001;177:13–20. 4. Vano E, Arranz L, Sastre JM, Moro C, Ledo A, Garate MT, et al. Dosimetric and radiation protection considerations based on some cases of patient skin injuries in interventional cardiology. Br J Radiol 1998;71:510–6. 5. Vano E, Goicolea J, Galvan C, Gonzalez L, Meiggs L, Ten JI, et al. Skin radiation injuries in patients following repeated coronary angioplasty procedures. Br J Radiol 2001;74:1023–31. 6. Vano E, Prieto C, Fernandez JM, Gonzalez L, Sabate M, Galvan C. Skin dose and dose–area product values in patients undergoing intracoronary brachytherapy. Br J Radiol 2003;76:32–8.


7. Prieto C, Vano E, Fernandez JM, Sabate M, Gonzalez L, Aviles P. Staff doses in procedures of intracoronary brachytherapy using beta sources. IAEA-CN-91/105 Contributed papers. International Conference on Occupational Radiation Protection: Protecting Workers against Exposure to Ionizing Radiation. Geneva, Switzerland 26–30 August 2002. 8. Po¨tter R, Van Limbergen E, Dries W, Popowski Y, Coen V, Fellner C, et al. Recommendations of the EVA GEC ESTRO Working Group: prescribing, recording, and reporting in endovascular brachytherapy. Quality assurance, equipment, personnel and education. Radiother Oncol 2001;59:339–60. 9. Nath R, Amols H, Coffey C, Duggan D, Jani S, Li Z, et al. Intravascular brachytherapy physics: Report of the AAPM Radiation Therapy Committee Task Group No 60. Med Phys 1999;26:119–52. 10. Faulkner K. Introduction to constancy check protocols in fluoroscopic systems. Radiat Prot Dosim 2001;94:65–8. 11. Vano E, Guibelalde E, Fernańdez JM, Gonza´lez L, Ten JI. Patient dosimetry in interventional radiology using slow films. Br J Radiol 1997;70:195–200. 12. Guibelalde E, Vano E, Gonzalez L, Prieto C, Fernańdez JM, Ten JI. Practical aspects for the evaluation of skin doses in interventional cardiology using a new slow film. Br J Radiol 2003;76:332–6. 13. Balter S, Vano E, Gonzalez L. Fluoroscopic patient dosimetry from DICOM headers. 11th International Congress of the International Radiation Protection Association. http://www. irpa11.com. Session 4c1. Madrid. Spain. 23–28 May 2004. Paper available at: www.irpa11.com/new/pdfs/4c1.pdf (Accessed 12 August 2005). 14. Vano E, Fernandez JM, Ten JI, Guibelalde E, Gonzalez L, Pedrosa C. Real-time measurement and audit of radiation dose to patients undergoing computed radiography. Radiology 2002;225:283–8. 15. European Commission. DIMOND III project. http://www. dimond3.org/ (Accessed 6 August 2005). 16. Vano E, Aviles P, Prieto C, Fernandez JM, Guibelalde E, Galvan C, et al. A dosimetric trial for the clinical follow-up of potential skin injuries on patients undergoing interventional cardiology procedures. 11th International Congress of the International Radiation Protection Association. ID: 1307. Madrid. Spain. 23–28 May 2004. Paper available at: www. irpa11.com/new/pdfs/4c11.pdf (Accessed 12 August 2005). 17. Sabate´ M, Costa MA, Kozuma K, Kay P, van der Giessen J, Coen V, et al. Geographic miss: a cause of treatment failure in radio-oncology applied to intracoronary radiation therapy. Circulation 2000;101:2467–71. 18. Neofotistou V, Vano E, Padovani R, Kotre J, Dowling A, Toivonen M, et al. Preliminary reference levels in interventional cardiology. Eur Radiol 2003;13:2259–63. 19. Prieto C, Vano E, Fernandez JM, Sabate M, Galvan C, Meiggs L, et al. Radiation protection training in intracoronary brachytherapy. Proceedings of the II International Conference. Radiation Protection Training: Future Strategies. 2003. Ed. CIEMA, Madrid, Spain. ISBN: 847834-450-0.

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What is the optimum breast plan: a study based on the START trial plans 1

K VENABLES, PhD, MIPEM, 1E A MILES, MPhil, DCR(T), 2E G A AIRD, PhD, FIPEM and 1P J HOSKIN, FCRP, FRCR,

on behalf of the START trial management group 1

Marie Curie Research Wing and 2Medical Physics Department, Mount Vernon Hospital, Rickmansworth Road, Northwood, Middlesex HA6 2RN, UK ABSTRACT. Each year thousands of women within the UK are treated with radiotherapy for breast cancer. The majority of these women are treated using a medial and lateral tangential field. This study evaluates the plans submitted to the quality assurance (QA) team of the START trial and investigates some of the differences between departments. Throughout the START trial, hardcopies of the radiotherapy dose distribution on the central slice for one in three women were submitted to the QA team for analysis. The QA team measured physical parameters including breast size and lung depth as well as noting parameters used for the radiotherapy delivery including beam energy, field size and wedge angle. Over 1400 plans from 36 centres were analysed. The mean patient separation was 19.7 cm (SD 2.7 cm) with a mean lung depth of 1.5 cm (SD 0.7 cm). The modal beam energy was 6 MV and the mean wedge angle was 23 ˚. Significant differences in the choice of wedge angle between departments were noted; however, in 90% of cases the resultant plan complied with the maximum dose gradient of 10% on the central axis specified by the trial protocol. Less than 3% (37 plans) had dose gradients of greater than 12%. This resulted in a mean dose gradient for all patients on the central axis of 5.7% (SD 2.9%).

Many women with breast cancer are treated with radiotherapy each year. Guidelines on the production of isodose distributions have been given by the ICRU [1]. The START trial is a multicentre UK trial of breast radiotherapy comparing different fractionation regimens that accrued 4451 patients between January 1999 and December 2002 [2]. There were two randomization options called A and B. Trial A was a three way randomization to either 50 Gy in 25 fractions, treated daily over 5 weeks, or one of two schedules treating five times per fortnight to a total dose of 39 Gy or 41.6 Gy in 13 fractions of 3.0 Gy or 3.2 Gy, respectively. Trial B patients were randomized to 50 Gy in 25 fractions over 5 weeks or 40 Gy in 15 fractions over 3 weeks, treated daily. Guidelines on the production of isodose distributions were provided in the trial protocol, which stated that two tangential beams, angled to remove divergence in the lung, should be used. Departments were asked to use wedges to achieve a dose gradient of less than 10% across the breast on the central axis slice. In addition, cobalt-60 was only permitted for patients with a separation of less than 18 cm. The trial has had a quality assurance (QA) programme from the outset. One of the aims of the QA team was to document the treatment technique in each centre and to ensure compliance with the protocol. The treatment of breast cancer with radiotherapy is evolving and there is an increased awareness of the need for compensation in three dimensions. Within the START trial, only two centres used compensation other than a wedge. For some women, particularly those with larger breast sizes, this may be necessary to achieve dose uniformity particularly when doses away from the 734

Received 24 October 2005 Revised 16 December 2005 Accepted 9 January 2006 DOI: 10.1259/bjr/80814021 ’ 2006 The British Institute of Radiology

central axis are considered. Areas of increased dose within the breast may be the cause of breast pain and result in poorer cosmesis [3], while irradiation of large lung volumes could potentially lead to radiation pneumonitis [4]. In current clinical practice the majority of patients are treated with two opposing tangential fields. This paper analyses the data for these patients.

Method Radiation dose distributions for 1 in 3 of the patients, assigned at the time of randomization, were collected by the QA team. Hardcopy central axis distributions were sent to the QA team after the patient had completed treatment. From these, data on the physical shape of the patient’s breast, the linac parameters used for treatment and the resulting isodose distributions were collected. Data from each plan were checked by a second member of the QA team. Universal wedges were converted to an equivalent wedge angle using the formula: tan h~

w1 tan (60) w1 zw2

where w1 is the weight of the wedge beam and w2 the weight of the corresponding open field. The wedge angle of the universal wedge was assumed to be 60 ˚ for all departments. During the analysis a further parameter of patient shape was defined: max distance from post: field edge to skin breast shape~ patient separation The British Journal of Radiology, September 2006

START breast plans

Dose gradient was defined as the difference in dose between a minimum point dose and maximum isodose encompassing an area of 2 cm2. Unless there was a cold spot in the centre of the breast, departments were asked to record the minimum dose at a point 0.5 cm above the lung boundary on the perpendicular bisector of the posterior beam edge and at least 1 cm from the 50% field edge. This point was chosen so as to be in breast tissue rather than ribs, and to reduce the differences due solely to the use of different planning systems, which vary in their ability to account for lack of lateral scatter from the lung. The QA team visited each department to observe simulation, planning and treatment of patients and to perform measurements on breast shaped phantoms. The results of the measurements in phantoms have been previously published [5, 6]. Questions relating to planning policy were asked of the staff producing the isodose distributions. Of relevance to the work presented in this paper were the criteria for their planning distribution, such as the choice of wedge angle, the resultant position and size (area and intensity) of hot spots and density correction used for lung. In a number of cases, this changed during the trial as better simulation facilities, particularly CT scanners, became available. All analyses were performed using the SPSS program (SPSS Inc., Chicago, IL). Correlation (Spearman rank coefficient used if data not normally distributed) was used to test associations between dose gradient, wedge angle, breast shape and depth, patient separation, energy and wedge angle. Differences in dose gradient by centre were compared using analysis of variance (ANOVA). Data are presented using boxplots and scatter diagrams. In boxplots, the horizontal line shows the median value of the data, the length of the box represents the interquartile range, and the length of the line indicates the range excluding outliers and extremes (outliers/ extremes are values more than 1.5/2.5 box-lengths from the 25th percentile).

Results

Table 2. Patient treatment details Number of patients (% of total)

Site Technique Use of asymmetry

Breast 1338 (90 %) Chest wall 150 (10 %) Isocentric 1035 (70%) Fixed FSD 453 (30%) None 1290 (86.7%) Posterior border 133 (8.9%) Superior border 65 (4.4%)

FSD, focus to surface distance.

the patient separation (Spearman correlation coefficient 0.21 p,0.001) (Figure 3). Of the plans received by the QA team, 748 incorporated a lung correction using either a standard value for all patients, which varied from 0.2 to 0.33 relative to water depending on centre, or bulk density corrections derived from CT data for an individual patient, which varied from 0.19 to 0.5 (53 plans). Ten patients were planned using a pixel by pixel correction for inhomogeneity. The START trial stated that the recommended dose gradient across the patient breast be less than 10% on the central axis and 1390 plans (93.4%) complied with this. A further 95 plans (6.4%) exhibited dose gradients of between 10% and 15%. The remaining 3 plans (0.2%) had a dose gradient of greater than 15%. ICRU guidance suggests that plans should comply with a dose of 95– 107% prescription dose over the target volume [1]. It was only possible to assess central sections and of these, 1232 plans (83%) complied with these recommendations. Dose gradient was found to be weakly correlated to breast depth (r50.50, p,0.001) (Figure 4) and with patient separation (r 5 0.57, p,0.001) (Figure 5). This latter figure shows data separated by energy; as expected, the dose gradients are less if higher energy machines are used. The correlation between gradient and separation also improves if each energy is considered separately, and this is shown in Table 3.

1488 plans were collected from 36 departments (1 centre input patients from 2 sites; these were counted as separate departments by the QA team since the planning systems were different on the two sites). Patient characteristics and treatment techniques are shown in Tables 1 and 2. The depths of lung incorporated into the treatment plan varied between departments and is shown in Figure 1. A range of energies from Co60 to 10 MV was used for the treatment of these patients. The majority were treated using linear accelerators with a nominal energy of 6 MV. Energy used for treatment is illustrated in Figure 2. The energy used only correlated weakly with Table 1. Patient demographics

Separation Breast depth Lung depth

Minimum (cm)

Maximum (cm)

Mean (cm)

Standard deviation

13.4 0.5 0

30.2 13.2 3.5

19.8 5.2 1.4

2.7 1.9 0.6


Figure 1. Box plot illustrating variation in lung depth vs centre.

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K Venables, E A Miles, E G A Aird et al

Figure 2. Energy used for treatment (n51488).

The mean wedge angle used was 22 ˚ for patients after lumpectomy (range 0–60 ˚) and 33 ˚ (range 0–60 ˚) for patients post-mastectomy. 690 patients were treated on linear accelerators with external fixed wedges (although a small number of these patients may have been treated using enhanced dynamic wedge (EDW). No significant difference was found between the medial and lateral wedge angles. A correlation was found between wedge angle and patient shape, which was stronger for some

departments than for others. A department with strong correlation is shown in Figure 6 and the results for all patients in the trial in Figure 7. For centres where only fixed wedge angles were available, less correlation is observed. Wedge angle and patient shape did not correlate with an increased uniformity of dose gradient for the same centre. There was a negative correlation between wedge angle and dose gradient (medial wedge r520.35, p,0.001, lateral wedge r520.34, p,0.001). Differences in wedge angles used were also seen between departments. Significant differences were seen in the dose gradients obtained from different centres (ANOVA p,0.001). Further analysis of these differences revealed that even when comparing distributions from different departments with the same planning system and same nominal energy of accelerator (6 MV), differences were evident, as shown by Figure 8. Of particular note is the difference between centres B and E, both of whom put large numbers of patients into the trial. Centre B did not apply any lung correction but allowed for the effect of lung by planning with the apex of the breast hotter than the medial or lateral edges. Centre E estimated a lung shape based on the maximum lung depth and applied a lung correction of 0.25. The dose gradient recorded is not affected by the incorporation of lung correction. The planning system was a simple beam library system and the reduction in dose at the minimum point due to the loss of lateral scatter when the lung was included would not be reflected in the calculation. The lower than average dose gradient for centre B might suggest that they were not adequately accounting for the effect of the lung and were using steeper wedge angles than would have been used on a lung corrected plan. The higher than average dose gradients recorded by centre E might suggest that their estimate of lung shape was not sufficiently accurate and that they were overestimating the volume of lung in the plan. The distribution of radiation within the breast was classified by comparing the dose at the apex of the breast (taken as 1.5 cm below the skin surface, Figure 9) with the dose in the most medial and lateral parts of the breast at least 1 cm from build up or penumbra regions. The majority of patients (70%) had an even distribution of radiation between these areas. However, in 11% of plans the medial and lateral aspects of the breast were hotter than the apex of the breast by 2–4%. When only those patients who do not have a lung correction are considered, the proportion in which the medial and lateral aspects were hotter than the breast apex was slightly less (8%).

Discussion

Figure 3. Association between energy used for treatment and average patient separation. Lines give the range of patient separations treated at each energy.

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There was no correlation between average lung depth for a centre and average separation for that centre, which may have been observed if practice in some centres placed the lateral border more posteriorly than others. The lung depth included in the treatment field is affected by the patient’s position; thus it would be necessary for some clinicians to include more of the lung in order to comply with the protocol recommendations on the positioning of the borders. In 1988, Fraass et al recommended the incorporation of lung density correction The British Journal of Radiology, September 2006

START breast plans

Figure 4. Relationship between dose gradient (%) and breast depth.

[7]. This is common practice in the UK with just over half of centres (51%) incorporating a correction and the majority of others planning to compensate for the effect of lung. The weak correlation between energy and patient size may be related to the availability of different energies. In departments where high energy machines were available for any patient, there may be greater tendency for its use – one department treated 53% of patients using 10 MV photons. The depth of the 95% isodose increases as higher energy machines are used. There will therefore be a decrease in the coverage of the superficial regions of the planning target volume (PTV) when 10 MV photons are used compared with 6 MV photons. The depth below

the skin surface at which the PTV should be drawn is usually taken to be 0.5 cm, however, compromise in this region is often preferred to decreased dose at depth. The exception is post-mastectomy patients for whom the loss of dose in the superficial tissues, when treated at high energy, represents a significant percentage of the planning target volume. Four patients in the trial who had undergone a mastectomy were treated using 10 MV photons without the application of bolus. For the majority of machines in clinical use in the UK, the wedges are housed within the treatment head. However, some patients are treated on units with externally mounted wedges. The scatter from these external wedges can increase the dose to the patient’s

Figure 5. Relationship between dose gradient and separation. Data for 6 MV have been excluded from the graph to allow the trends for other energies to be seen.


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K Venables, E A Miles, E G A Aird et al Table 3. Correlations between dose gradient and patient separation with respect to energy used for treatment Energy

Spearman correlation

Number

Co60 4 MV 5 MV 6 MV 8 MV 10 MV All patients

0.641 0.608 0.709 0.58 0.797 0.724 0.57

28 50 76 1249 21 66 1488

skin and contralateral breast and therefore it has been recommended by some authors that the wedging be applied predominantly from the lateral beam [8]. This practice was not seen in the START data where the majority of patients had the same wedge used for both medial and lateral fields. Differences were seen in the wedge angles used in different departments for the same ‘‘shape’’ of breast. The wedge angle needed will be affected by the energy of machine used. However, 75% of departments treated more than 80% of their breast patients at 6 MV and thus comparisons between departments are possible. When using fixed wedge angles there is less scope for optimization of the plan, whilst maintaining rapid treatment times. Departments using large wedge angles tend to produce plans in which there is more dose towards the medial and lateral aspect of the breast (see Figure 9 for definition of points). In addition, where lung correction has been omitted, true dose in these regions may be higher than indicated on the plan, due to the increased transmission through lung, leading to a possible increase in the risk of rib fracture. Care must be taken in comparing distributions between planning systems due to the subtle differences in the implementation of the algorithms. However, differences in average dose gradient were also evident when comparing departments using the same planning system and same nominal energy of machine. No further

Figure 7. Correlation between wedge angle and breast shape for all patients in the trial r520.39, p,0.001. Note the wide spread in breast shapes for the fixed wedge angles 0 ˚, 15 ˚, 30 ˚, 45 ˚ and 60 ˚.

statistical analysis of this aspect was performed because of the small numbers in some of the groups, but it reflects differences in planning policy. Some departments with conventional rather than CT simulators estimated the lung shape using either a rule of thumb or a lung template. Although this will usually give a more realistic estimate of the distribution than ignoring the lung correction completely, it may result in a different distribution from that which would be produced if more accurate information was available. In some instances plans are produced in which no attempt is made to estimate the lung outline. In these cases it is not possible to incorporate lung correction to compensate for the reduced attenuation of the lung relative to breast tissue. One way to overcome this is to produce a plan which indicates an artificial dose gradient across the breast that will not be present in

Figure 6. Correlation between wedge angle and breast shape (max. distance from post edge to skin/patient separation) for centre 12. This was a large centre with a wide range of planning staff, but with rigid guidelines on the final distribution (r520.81 p,0.001, n544).

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Figure 8. Box plot showing difference in dose gradient for different centres using the same planning system and same nominal machine energy. Circles represent outliers. The British Journal of Radiology, September 2006

START breast plans

should be present (lower if only a single slice is assessed) and the hotspots should be located evenly throughout the breast tissue. Ideally, patients should have CT scans to obtain accurate information on the position of the lung. If this is not possible, the effect of low density lung tissue should be carefully considered and the distribution should appear hotter at the apex of the breast. For some patients, compromise must be reached and the planner should carefully weigh the increased penetration available by using higher energy beams against the decrease in dose to superficial tumours. In particular, if post-mastectomy radiotherapy is to be given using photons with energies higher than 6 MV then the use of bolus should be considered.

Figure 9. Definitions used for points of the breast.

the patient. For these patients, it is expected that plans sent to the QA team will aim for the dose at the apex of the breast on the plan to be hotter than that at the medial and lateral border. For a small percentage of patients, the reverse of this situation was seen and the medial and lateral aspect of the plan was hotter than the apex. In a few cases, this may be necessary to compensate adequately for variations in the shape of the breast in three dimensions, but should not be normal practice if only the central axis of the breast is considered. For the majority of patients, only data on the central slice was sent to the QA team; this will mask the true dose gradients which will be present in the patient’s breast as a whole. For some women to conform with ICRU recommendations of a maximum dose gradient of 95–107%, intensity-modulated radiotherapy will be necessary to eliminate hot spots in the superior and/or inferior aspect of the breast.

Conclusion The analysis of the plans submitted to the QA team emphasises the high degree of protocol compliance that was achieved for this trial. The collection of central axis distributions by the QA team enabled protocol violations to be flagged up to centres early in the trial and thus improved the compliance with the protocol. It also enabled the QA team to discuss alternative patient positioning where larger than average lung depths were seen. The data collected represents a snapshot of clinical practice in the years the trial was accruing and will aid the analysis of the main trial once mature survival, recurrence and late effect data are available. When producing a dose distribution for a patient, a number of factors must be considered. Ideal breast plans have a 95% isodose which covers breast tissue to within 5 mm of the skin surface. No hot spots in excess of 107%


Acknowledgments The START trial management group: Edwin Aird, Jane Barrett, Peter Barrett-Lee, Judith Bliss, Jackie Brown, John Dewar, Jane Dobbs, Jo Haviland, Penny Hopwood, Peter Hoskin, Pat Lawton, Brian Magee, David Morgan, Roger Owen, Eileen Parkin (RAGE Observer), Joyce Pritchard (RAGE Observer), Val Speechely, David Spooner, Mark Sydenham, Karen Venables, Elizabeth Miles and John Yarnold.

References 1. International Commission on Radiation Units and Measurements. Prescribing, recording and reporting photon beam therapy. ICRU Report 50. Bethesda, MD: ICRU, 1993. 2. START trial management group. Standardization of Breast Radiotherapy (START) Trial. Clin Oncol 1999;11:145–7. 3. Yarnold JR, Donovan EM, Bleakley NJ, Reise SF, Regan J, Denholm E, et al. Randomised trial of standard 2D radiotherapy (RT) versus 3D intensity modulated radiotherapy (IMRT) in patients prescribed breast radiotherapy. Radiother Oncol 2002;64:S15. 4. Early Breast Cancer Trialists’ Collaborative Group. Favourable and unfavourable effects on long-term survival of radiotherapy for early breast cancer: an overview of the randomised trials. Lancet 2000;355:1757–70. 5. Venables K, Winfield EA, Deighton A, Aird EGA, Hoskin PJ. The START Trial – measurements in semi-anatomical breast and chest wall phantoms. Phys Med Biol 2001;46:1937–48. 6. Venables K, Winfield EA, Aird EGA, Hoskin PJ. Threedimensional distribution of radiation within the breast: an intercomparison of departments participating in the START trial of breast radiotherapy fractionation. Int J Radiat Oncol Biol Phys 2003;55:271–9. 7. Fraass BA, Lichter AS, McShan DL, Yanke BR, Diaz RF, Yeakel KS, et al. The influence of lung density corrections on treatment planning for primary breast cancer. Int J Radiat Oncol Biol Phys 1988;14:179–90. 8. Ikner CL, Russo R, Podgorsak MB, Proulx GM, Lee RJ. Comparison of the homogeneity of breast dose distributions with and without the medial wedge. Med Dosim 1998;23: 89–94.

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Feasibility of automated matching of supine and prone CT-colonography examinations 1

A H DE VRIES, MD, 2R TRUYEN, MSc, 2J 2 F GERRITSEN, MSc and 1J STOKER, MD

VAN DER

PEIJL,

MSc,

1

J FLORIE,

MD,

1

RE

VAN

GELDER,

MD,

1

Department of Radiology, Academic Medical Center, Amsterdam and 2Philips Medical Systems, Best, The Netherlands ABSTRACT. Matching of prone and supine positions in CT colonography may improve accuracy of polyp detection. The purpose of this study was to investigate the feasibility of automatic prone-supine matching in CT-colonography using proven polyps as fixed points of reference. The method is based on similarities in the direction of centre-lines and allows for compression and extraction of the centre-lines in both positions. To illustrate the impact of the match error of the new method in practice, the visibility of the matched polyps in a primary three-dimensional unfolded cube setting was determined as well. The method was compared with a method that relies on the normalized distance along the centre-line (NDAC method). The median absolute match error was 14 mm (range 0–59 mm, average 20 mm) either proximal or distal from the actual polyp in prone position. In the observer study, 70% (26/37) of the polyps were directly visible in prone view. The overall difference in median absolute match error between both methods was small (2 mm), although half way along the centre-line there were polyps with substantial differences in match error (larger with NDAC). We concluded that automated prone-supine matching of CT-colonography studies is feasible and has a low match error. The difference with the NDAC method was small and not significant, although half way along the centre-line some differences were seen.

Colorectal cancer is the second leading cause of cancer-related mortality in the western world. CTcolonography is considered as a potential screening tool for colorectal cancer. To improve the accuracy of polyp detection, patients are both scanned and examined in the prone and supine position. Combining information from these scans will assist the reviewer in evaluating colon segments and differentiating polyps from faeces or folds. As most faecal material is subject to gravity, combining both scans may be able to increase specificity of CTcolonography. Since reference-points (e.g. hepatic flexure) are often not fixed, manual verification of findings on supine and prone positions may be a time-consuming activity. Using an automated supine-prone matching algorithm may facilitate this process, and may lead to a more efficient interpretation of CT-colonography. The method evaluated in this article is based on similarities in direction of the centre-line and allows for compression and extraction of the centre-lines in both positions. The first aim of the study was to assess the feasibility of automated matching of supine and prone CTcolonography examinations with colonoscopically proven polyps as fixed points of reference. Therefore the match error was calculated. The visibility of matched polyps was determined as well. This was carried out in a For this research project, a grant from Philips Medical Systems was received.

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Received 12 September 2005 Revised 9 January 2006 Accepted 18 January 2006 DOI: 10.1259/bjr/55953054 ’ 2006 The British Institute of Radiology

three-dimensional setting after a match with this new method was calculated. Third, the method was compared with a method that relies on the relative location between start and end of both paths.

Materials and methods To evaluate the algorithm, CT-colonography examinations with colonoscopically proven polyps (>5 mm) were used. These examinations formed part of a comparative study of CT colonography and colonoscopy in a surveillance population [1]. All polyps could be identified three dimensionally in both prone and supine positions. The polyps in both positions were linked, based on the three criteria of size, location and morphology. Polyps were excluded in cases where there could be doubt about the accuracy of these links. After manual insufflation of colon and rectum and intravenous administration of either 20 mg butylscopolamine bromide (Buscopan; Boehringer-Ingelheim, Ingelheim, Germany) or 1 mg glucagon hydrochloride (Glucagen; Novo Nordisk A/S, Bagsvaerd, Denmark), patients were scanned with a four-slice CT scanner (Mx8000; Philips Medical Systems, Best, The Netherlands) in supine and prone positions. Scanparameters were as follows: 120 kV, collimation 4 mm 6 2.5 mm, rotation-time 0.75 s, pitch 1.25, slice-thickness 3.2 mm, reconstruction interval 1.6 mm and reconstruction filter C. The tube-current varied between 25 mAs The British Journal of Radiology, September 2006

Automated matching of supine and prone CT-colonography

and 70 mAs depending on the waist circumference of the patient. Following scanning, centre-lines were calculated semiautomatically; a ‘‘seed’’ was placed in the proximal colon and discontinuities of the colon were bridged manually. For calculation of the centre-line, software was used to reconstruct the colon three-dimensionally (Easy Vision; Philips Medical Systems, Best, The Netherlands). The matching of prone and supine was performed based on two principles: 1. maximal alignment of centre-line-directions of both prone and supine positions. 2. limited compression or expansion of the centrelines to make the maximal alignment of corresponding segments possible. In order to reach maximal alignment, the outcome of the following function needs to be minimum: qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi 2 2 2ffi fi,j ~ xdiri {xdirj z ydiri {ydirj z zdiri {zdirj Low values of fi,j imply a small difference in centre-line direction (dir) between sample points i (prone) and j (supine) in the three dimensions (x, y and z), indicating a good match. In this step, the points with the same direction in both positions are therefore matched as well as possible. Second, since corresponding directions of both centrelines are often not situated in exactly the same part of the colon, expansion and compression of the paths was applied. To avoid matching of the hepatic flexure in supine position to a curvature in the sigmoid colon in prone position, a penalty-value for expansion and compression of the paths was applied. This penalty value was proportional to the expansion or compression used. Therefore, the more the centre-line was manipulated in order to match parts with a similar direction, the higher the penalty value was. The sum of the outcome of the function of maximal alignment and the penalty value of the expansion/ compression formed the match cost. With the use of dynamic programming [2], the match cost was computed for each possible combination of points on the centre line in both positions. The combination of points was made such that the total cost (i.e. the sum of all individual costs) was minimal. The performance of the algorithm was measured using the absolute match error (in millimetres). This match error was calculated by measuring the distance from the actual polyp location to the calculated (expected) polyp position, [A] (Figure 1). This was done along the central axis in prone position after the match was made. To illustrate the impact of the match error of the new method in practice, the visibility of the matched polyps in a primary three-dimensional unfolded cube setting was determined as well. This was done by a research fellow (JF, medical doctor), who had evaluated over 100 primary three-dimensional CT-colonographies (all colonoscopically verified). The reviewer was presented a polyp in supine position that was to be indicated in the The British Journal of Radiology, September 2006

Figure 1. The performance of the algorithm was calculated by measuring the distance from the actual polyp location in supine position to the calculated (expected) polyp position in prone position after the match was made. (A, match error).

adjacent prone position, after the match was calculated (Figure 2). The visibility was measured on a four point scale: I. The corresponding polyp was clearly visible without scrolling along the central axis. II. As I, but to be certain the reviewer scrolled along the central axis. III. The corresponding polyp could not be seen instantly, but was traceable after scrolling along the central axis within a margin of 150 mm from the calculated location. IV. The corresponding polyp could not be found by the reviewer within a margin of 150 mm from the initial spot of the virtual camera. The match was considered successful if the matched polyp could be classified under category 1 or 2. This method was compared with a method based on the relative location between start and endpoint of this centre line (Normalized Distance Along the Centre-line, NDAC) [3]. Here, the beginning of the centre-line, the caecum, is located at index 0 on the path and the end, the rectum, is defined as index 1. All other points will have indices between 0 and 1. The NDAC absolute match error is then defined as: Absolute match error5[NDAC polypsupine 2 NDACpolypprone] 6 Length Central Axis prone The Wilcoxon-test was performed to test for differences in absolute match error in both methods. The null hypothesis was that there was no difference between both methods. With a p-value of less than 0.05, this hypothesis could be rejected.

Results 32 of the 249 CT-colonography studies from the former study included one or more polyps >5 mm that were visible in both supine and prone examinations. These 32 741

A H de Vries, R Truyen, J van der Peijl et al

Figure 2. User interface panel shows both the supine and prone dataset in an unfolded three-dimensional rendering, in a twodimensional axial rendering and in a volume rendered overview of the colon. On the left the supine position was displayed, on the right the prone position. Interactive navigation along the pre-calculated centre-line was possible. This panel shows a calculated match. In this example a match error of 27 mm was measured and the corresponding polyp was immediately visible (category I). This match error must not be confused with the difference in ‘‘distance along path’’. The matching was evaluated with the use of a colonoscopically proven polyp (arrow) (Prototype based on EasyVision; Philips Medical Systems, Best, The Netherlands).

examinations included 53 polyps (5 mm or larger) visible in both positions. One polyp in two patients was excluded because of invisibility on the three dimensional display. These polyps in the rectum were both hidden behind the balloon of the inserted catheter and could only be seen in a two-dimensional read. One patient was excluded because 12 of the 14 polyps were situated in the rectum and sigmoid. Here there could be doubt about the correctness of the exact linking of the polyps seen in colonoscopy and colonography since many morphologically less specific polyps were seen. In total, 16 polyps in three patients were excluded. Of the remaining 37 polyps, 26 (70%) were sessile, 8 (22%) were pedunculated and 3 (8%) were flat. 24 polyps were 5–9 mm, 11 polyps were 10–14 mm and 2 polyps were larger than 15 mm. These were both carcinomas (5%). Of the remaining polyps 14 (38%) were adenomas, 10 (27%) were non-adenomas and of the remaining polyps histology was not obtained. 742

Two polyps were situated in the descending colon and four polyps in the rectum. In the remaining four segments, 7 to 9 polyps were situated. Four of the 29 patients had undergone a hemicolectomy. The median absolute match error was 14 mm (range 0– 59 mm, average 20 mm) either proximal or distal from the actual polyp in prone position. In the observer study, 70% (26/37) of the polyps was directly visible in prone view. Of these directly visible polyps, 20 polyps (54%) were seen instantly without flying though the colon (category I), the remaining 6 polyps (16%) were also instantly visible, but the observer chose to move the virtual camera over a small distance to verify the polyp (category II). The remaining 11 polyps (30%) could not be seen instantly but all were found after a flight through the colon within 150 mm of the matching position (category III). When the NDAC method was applied to our CTcolonography examinations, this resulted in a median The British Journal of Radiology, September 2006

Automated matching of supine and prone CT-colonography

Figure 3. (a) The absolute match error of the method based on the maximal alignment of the centre-line plotted against the relative location of a polyp along the centre-line in supine position. The 0 on the x-axis corresponds to caecum and 1 to rectum. (b) The absolute match error of the NDAC method plotted against the relative location of a polyp along the centre-line in supine position. The 0 on the x-axis corresponds to caecum and 1 to rectum. The three circled arrows indicate polyps from three different patients with a relatively large absolute match error.

absolute match error of 16 mm (range 0.5–105 mm, average 26 mm), slightly larger than the other method. There was no significant difference (p50.502) in match error between both methods. Half way along the centre-line three polyps were visible, with a substantial difference in match error in favour of the method of maximal alignment of centrelines (Figure 3). These polyps were all situated in the transverse colon. None of these three patients had undergone a hemicolectomy.

Discussion This study shows that the matching of prone and supine CT-colonography examinations is feasible; the median absolute match error was 14 mm and 70% of all polyps were visible after the match was made. The overall difference in median absolute match error between both methods was small (2 mm), although half way along the centre-line there were polyps with substantial differences in match error (larger with NDAC). All these polyps were situated in the transverse colon. The transverse colon is the largest intraperitoneally situated colonic segment [4] that can move relatively freely in the abdominal cavity. We assume that this is the most mobile part of the colon, based on our observations of the three-dimensional overviews in both positions. The fact that polyps with a substantial difference in match error are situated in this segment might suggest that the method that relies on maximal alignment of centre-lines has an advantage in this mobile region. The methods of matching prone and supine positions were evaluated based on the assumption that the included polyps were (immobile) true polyps and, second, that the actual link between these polyps in both positions was correct. In case there was doubt about the status of the polyp and linking, the polyp was excluded. With these exclusion criteria a very robust reference standard was created. Consequently, using this The British Journal of Radiology, September 2006

reference standard we were able to determine most precisely the matching quality of the methods. Matching is important in determining whether a suspected lesion is a polyp or faecal material. A rule of thumb is that a suspected lesion with an unchanged position in prone and supine is most likely a polyp, while a change of position is related to stool. An accurate matching algorithm is therefore helpful to reduce the number of false positive findings in an efficient way. Pitfalls in diagnosis can be caused by adherent stool mimicking a polyp, in case faecal tagging is absent or inadequate. On the other hand, pedunculated polyps (large stalk) or polyps situated in mobile segments may move (or seem to move) when both positions are compared. Therefore, when assessing whether a suspicious finding is a polyp, it is important not only to take the location into account, but also homogeneity and morphology of the lesion. Incorporated in a CAD algorithm, this matching tool could also be used to reduce the number of false positive findings. Na¨ppi et al [5] used a region-growing scheme with distance calculations to divide the colonic lumen into overlapping segments that match in the supine and prone data sets. Polyp candidates detected by means of a CAD scheme were eliminated if they could be seen in only one of the two corresponding segments. A limiting factor for generalizing the visibility of the matched polyps is the use of a 3D unfolded cube method for reviewing the data. Using this method, 70% of the polyps could be seen instantly when the match was made. This review method may have had a positive influence on the percentage of polyps that were seen instantly, since the unfolded cube display method creates a 360 ˚ view without major distortion of the image. Although evaluated here in a 3D-setting, the method of matching prone and supine datasets can be applied in a 2Dsetting as well. This can be done on condition that a centreline is calculated through both corresponding colons. Theoretically, the matching quality of the method of maximal alignment is not altered by absence of a part of 743

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the beginning or ending of a centre-line in one position (e.g. caecum or rectum). In other words, the absence of baseline adjustment, a condition in other clinically evaluated methods [3, 6, 7], will not have a large influence on the match error in this new method. This baseline adjustment may not be possible in patients with an inadequately distended or insufficiently cleansed caecum or rectum in one position. Although this condition was not present in any of the data sets used, this proposition was confirmed in a test case. Here we removed 50 cm of the beginning of the centre-line without large consequences for the match error in the remaining colon segments. The design of this feasibility study, with selected polyps as fixed reference points, precludes an assessment of the clinical value of the matching tool. Further study should assess matching quality and time efficiency of the matching tool in a series of unselected CTcolonography examinations. From this study we conclude that automated pronesupine matching of CT-colonography studies is feasible and the difference in median absolute match error from the NDAC method was small and not significant, although half way along the centre-line (transverse colon) more pronounced differences were seen in some cases.

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References 1. van Gelder RE, Nio CY, Florie J, Bartelsman JF, Snel P, De Jager SW, et al. Computed tomographic colonography compared with colonoscopy in patients at increased risk for colorectal cancer. Gastroenterology 2004;127:41–8. 2. Rouchka E. Dynamic programming. Stockholm Bioinformatics Center 2004 June 1. Available from: http:// www.sbc.su.se/,per/molbioinfo2001/dynprog/dynamic. html [Accessed 23 March 2006]. 3. Iordanescu G, Summers RM. Automated centerline for computed tomography colonography. Acad Radiol 2003;10:1291–301. 4. Saunders BP, Fukumoto M, Halligan S, Jobling C, Moussa ME, Bartram CI, et al. Why is colonoscopy more difficult in women? Gastrointest Endosc 1996;43(2 Pt 1):124–6. 5. Nappi J, Okamura A, Frimmel H, Dachman A, Yoshida H. Region-based supine-prone correspondence for the reduction of false-positive CAD polyp candidates in CT colonography. Acad Radiol 2005;12:695–707. 6. Acar B, Napel S, Paik DS, Li P, Yee J, Beaulieu CF. Registration of supine and prone CT colonography data: Method and evaluation. Radiology 2001;221:332. 7. Li P, Napel S, Acar B, Paik DS, Jeffrey RB, Beaulieu CF. Registration of central paths and colonic polyps between supine and prone scans in computed tomography colonography: pilot study. Med Phys 2004;31:2912–23.



Improved motion compensation in 3D-CT using respiratorycorrelated segment reconstruction: diagnostic and radiotherapy applications S MORI,

PhD,

M ENDO,

PhD,

R KOHNO,

PhD

and S MINOHARA,

PhD

National Institute of Radiological Sciences, Inage-ku, Chiba 263-8555, Japan ABSTRACT. Conventional respiratory-gated CT and four-dimensional CT (4DCT) are disadvantaged by their low temporal resolution, which results in the inclusion of anatomic motion-induced artefacts. These represent a significant source of error both in radiotherapy treatment planning for the thorax and upper abdomen and in diagnostic procedures. In particular, temporal resolution and image quality are vitally important to accurate diagnosis and the minimization of planning target volume margin due to respiratory motion. To improve both temporal resolution and signal-tonoise ratio (SNR), we developed a respiratory-correlated segment reconstruction method (RS) and adapted it to the Feldkamp-Davis-Kress algorithm (FDK) with a 256 multidetector row CT (256MDCT). The 256MDCT scans approximately 100 mm in the craniocaudal direction with a 0.5 mm slice thickness in one rotation. Data acquisition for the RS-FDK relies on the assistance of a respiratory sensing system operating in cine scan mode (continuous axial scan with the table stationary). We evaluated the RS-FDK for volume accuracy and image noise in a phantom study with the 256MDCT and compared results with those for a full scan (FS-FDK), which is usually employed in conventional 4DCT and in half scan (HS-FDK). Results showed that the RS-FDK gave a more accurate volume than the others and had the same SNR as the FS-FDK. In a subsequent animal study, we demonstrated a practical sorting process for projection data which was unaffected by variations in respiratory period, and found that the RSFDK gave the clearest visualization among the three algorithms of the margins of the liver and pulmonary vessels. In summary, the RS-FDK algorithm provides multi-phase images with higher temporal resolution and better SNR. This method should prove useful when combined with new radiotherapeutic and diagnostic techniques.

Continuing rapid progress in computer hardware and software has led to better radiation therapy planning and dramatic improvements in delivery. New types of conformal planning and delivery technology, of which intensity-modulated radiation therapy (IMRT) is a prominent example [1, 2], have the potential to achieve a much higher degree of target conformity and normal tissue sparing than existing treatment techniques. Since higher target conformity generally requires more accurate definition of the target, the new radiation therapy requires a reduction in the target volume error arising from the respiratory motion of organs such as the lung and liver. Moreover, the human respiratory cycle is not strictly regular, but varies in amplitude and period from one cycle to the next [3–5], and greater time is spent in exhalation than in inhalation [6, 7]. These complexities hamper the accuracy of radiotherapy, for example in the determination of peak-to-peak amplitude of external chest motion. Voluntary or imposed breath-hold techniques have been proposed to reduce or eliminate these effects of breathing during both imaging and radiotherapy treatment [8–11], but these prolong treatment and in any case cannot be tolerated by many patients. Respiratory motion during CT acquisition may produce artefacts that resemble disease symptoms [12], and these remain a significant source of The British Journal of Radiology, September 2006


error in radiotherapy treatment planning for the thorax and upper abdomen [12–14]. Many investigators have introduced automatic respiratory-gated CT [15, 16] or four-dimensional CT (4DCT) [17, 18] acquisition techniques which provide wide craniocaudal (CC) coverage beyond the CT detector width without image gaps during free breathing, and have adapted them to radiation therapy [19–21]. In the respiratory-gated CT technique, CT images are taken by an axial scan gated to respiratory signals from a respiratory motion detection system. In contrast, the 4DCT technique obtains CT images by cine scan and sorts the CT images obtained in the same respiratory phase. The temporal resolution, therefore, is determined by the gantry rotation time. Due to their relatively low temporal resolution, however, these techniques do not remove motion artefacts completely. Segment reconstruction correlated to a physiological signal was first developed using electrocardiograms (ECG), and ECG-correlated reconstruction has been widely used to delineate coronary arteries with intravenous injection of contrast agent [22–26]. To our knowledge, however, few papers have appeared on respiratorycorrelated segment reconstruction. Koenig et al [27] introduced data acquisition processing on several half-turns in order to reduce the dose delivered per 745

S Mori, M Endo, R Kohno and S Minohara

rotation with the same signal-to-noise ratio (SNR). Sonke et al [28] reported the use of respiratory-correlated cone beam CT integrated with a linear accelerator, but their rotation speed of approximately 4 min is considerably longer than that with our 256 multidetector row CT (256MDCT), rendering their experience not directly applicable to the present study. To improve temporal resolution and image quality, and thereby minimize planning target volume (PTV) margin due to respiratory motion, we developed a respiratorycorrelated segment reconstruction method (RS) and adapted it to the Feldkamp-Davis-Kress algorithm (FDK) [29] with a 256MDCT [30]. The RS-FDK algorithm provides multiphase images with higher temporal resolution and a better SNR than the conventional respiratorygated CT and 4DCT techniques, and should therefore prove useful when combined with new radiotherapy techniques such as four-dimensional (4D) radiation therapy [31–34]. Here, with a view to clinical utilization, we evaluated the use of RS-FDK with the 256MDCT in phantom and animal studies.

Materials and methods Prototype 256 multidetector row CT-scanner (256MDCT) The 256MDCT was developed at the National Institute of Radiological Sciences (NIRS) [30, 35]. A wide-area two-dimensional (2D) detector was designed on the basis of existing CT technology [36] and mounted on the gantry frame of an advanced MDCT. The number of elements was 912 (transverse)6256 (craniocaudal), each of approximately 0.5 mm60.5 mm at the centre of rotation. Rotation time was 1.0 s. Owing to disk storage limitations, continuous scan time was limited to 15 s for acquisition at 256 mm60.5 mm, 30 s at 128 mm61.0 mm, and 60 s at 64 mm62.0 mm. The scanner could scan approximately 100 mm in the craniocaudal direction in one rotation. Data sampling rate was 900 views/s, and dynamic range of the A–D converter was 16 bits. The detector element consisted of a scintillator and photodiode, the former being the same as that used in the MDCT (Toshiba Aquilion; Toshiba Medical Systems, Japan). A FDK algorithm was used for reconstruction. Reconstruction of volume data of 51265126128 voxels with a high-speed image processor in a field-programmable gate array (FPGA)-based architecture took less than 1 s.

External respiratory signal tracking system and signal processing Generally, two kinds of external respiratory signal tracking system are used. These have tagging points with artificial markers that reflect or emit light, termed ‘‘passive’’ and ‘‘active’’. The latter type is routinely used for gated irradiation and CT acquisition during heavy ion radiotherapy at our institution [20]. A number of problems with these markers have been reported, such as a residual error between the marker position and surface point of interest, patient setup, and a reduction in resolution in 746

orthogonal directions [37–39]. Here, however, the passive marker was used to obtain respiratory phase alone, and not absolute distance of motion. The respiratory sensing system consisted of a workstation (Dell, Roundrock, TX) equipped with a real-time digital video analyser, in-house gating software and user interface within the PV-WAVE programming package (Visual Numerics, San Ramon, CA), and a chargedcoupled-device (CCD) camera (XC-EI50; Sony Corp., Tokyo, Japan) with an attached infrared illuminator (OTR, Tokyo, Japan) (Figure 1a). SNR of the camera signal was 60 dB and vibration tolerance was 10 G (20–200 Hz). Since the reflected marker was captured around the centre of view, distortion was negligible. The respiratory sensing system was affixed to the patient table in this study. To perform an RS scan using conventional CT, movement of the couch to the adjacent position is necessary to obtain the next respiratory cycle. Because this process is repeated until the entire scan range is completed, the respiratory sensing system may induce significant mechanical disturbances that would be amplified by zoom optics. In contrast, couch movement is not necessary for RS-FDK using the 256MDCT, and the vibration induced by rotation of the gantry does not significantly disturb the couch. Mechanical disturbance with this equipment is therefore negligible. The cycle of respiratory phase was monitored with a CCD camera focused on a reflective marker seal on the patient’s chest or abdominal area. Video signals from the camera were transferred to the workstation and software routines were run to determine respiratory function from the detected marker position. Tracking of the marker was facilitated by observation of the seal at very high contrast. A spline curve was fitted to the respiratory function to reduce noise in the detected signal and to correlate the sampling rate of 30 frame/s for respiratory function with the 900 frames/s for the projection data [40].

Reconstruction algorithms Three types of reconstruction algorithm based on the FDK were used, namely full scan (FS-FDK), half scan (HS-FDK) and respiratory-correlated segment FDK (RSFDK). The usual reconstruction algorithm, FS-FDK, uses a uniform weight over the 2p of the projections, and is usually employed in conventional respiratory-gated CT. To increase temporal resolution, we employed the HSFDK, which uses only p plus the fan angle of the projections with a Parker weighting function applied prior to the filtered back projection operation [41]. Results showed an effective scan time of 500 ms (central ray) when a 1.0 s rotation mode was used. Details of the HS-FDK have been described elsewhere [42]. With regard to the RS-FDK, its essential concept is to sort projection data in the same respiratory phase, rather than reconstructing CT images as in the conventional respiratory-gated CT technique. Data acquisition for anatomical sites subject to breathing motion such as the lung and abdomen relies on the assistance of the respiratory sensing system to reduce the impact of respiratory motion. The 256MDCT uses a cine scan mode (continuous axial scanning with the table stationary) to acquire all respiratory phase projection data (Figure 1a). The respiratory phase is determined from the respiratory The British Journal of Radiology, September 2006

Improved motion compensation in 3D-CT

Figure 1. Sorting process in the respiratory-correlated segment reconstruction method (RS). (a) A respiratory signal is obtained from the infrared video camera. (b) Projection data with projection angle are obtained from the 256MDCT. (c) Relationship between projection angle and moving object position. (d) Projection data for the same respiratory phase are sorted to obtain four data sections (sections A–D). (e) The four projectionquadrant sections – each section corresponds to p/2 of projections. (f) RS-FDK images are obtained after generating cone-beam back projection (FDK).

signal collected by the respiratory sensing system (see below) during the cine scan (Figure 1b,c). Projection angle shift per respiratory period, PS, between rotation time Trot and respiratory period Tres is defined as: PS~

1 mod (Tres ,Trot ) Trot

ð1Þ

where PS is normalized by Trot, and is given by a fraction of rotation. The number of sections, Ns, is then obtained as: 8 1 > > {1 > > < PS Ns ~ 2 > > i >h > : ð1{PSÞ{1

PS~0 0vPSv1=2 PS~1=2

ð2Þ

1=2vPSv1

[ ] denotes floor function, which yields the greatest integer lower or equal. Projection data for the same respiratory phase are sorted to obtain RS-FDK projection The British Journal of Radiology, September 2006

data PRS(b) as follows: 8 2p > > p(b) 0ƒbƒ > > > Ns > > > ::::::: > < PRS (b)~ p(bz2np) 2np ƒbƒ 2(nz1)p > Ns Ns > > > ::::::: > > > > > : p(bz2(Ns {1)p) 2(Ns {1)p ƒbƒ2p N

ð3Þ

s

where p(b) is the projection datum obtained from the cine scan at projection angle b, and n is an integer between 0 and Ns–1. The 2p projection data set thereby acquired is used for FDK reconstruction. For RS-FDK, the total scan time TS and temporal resolution TR are given as follows: TR~Trot =Ns

ð4Þ

TS~Tres :Ns

ð5Þ 747


We noted that it was necessary to desynchronize the respiratory period from the gantry rotation time, because if the respiratory period is a harmonic (PS50) or subharmonic (PS50.5) of this time, the restricted temporal resolution Trot or Trot/2 is obtained, respectively. In this study, the data set was divided into four projection-quadrant sections (Ns54), with each section therefore corresponding to p/2 of the projections (sections A–D) (Figure 1d), covering 2p (Figure 1e). Since temporal resolution was increased in proportion to the number of sections (Equation 5), temporal resolution of the RS-FDK was four and two times as high as those of the FS-FDK and HS-FDK, respectively. However, total scan time was also increased in proportion to the number of sections (Equation 5). To avoid high patient doses, total scan time was conservatively limited to 17 s, less than that proposed by the Institutional Review Board (IRB) of the National Institute of Radiological Sciences (NIRS). Multiphase images were obtained by shifting the start projection angle for FS-FDK and HS-FDK. Both algorithms provide volumetric multiphase images over the total acquisition period. In contrast, RS-FDK provided multiphase images by sorting projection data for each respiratory phase, and thereby produced images for only one respiratory cycle. Image noise ratios for FS-FDK, HS-FDK and RS-FDK are given in Appendix 1.

Phantom study Volume accuracy results for RS-FDK were compared with those for FS-FDK and HS-FDK following evaluation with a moving phantom designed for the 256MDCT (Figure 2). The phantom contained a 30 mm diameter acrylic ball as the target volume placed on a moving table connected to a mechanically driven motor with speed adjustment. To simulate human respiratory motion, sinusoidal movement was employed in an oblique direction by setting the moving phantom at 45 ˚ to the CC direction. Motion distance and time period of the moving phantom were 40 mm and 3.74 s, respectively. The phantom position was denoted using a motion function f(t) as follows: f (t)~20:sin(2p:t=3:74½s)½mm

ð6Þ

pffiffiffi pffiffiffi ðLR(t),AP(t),CC(t)Þ~ f (t)= 2,0,f (t)= 2

ð7Þ

where t denotes time [s], and (LR(t), AP(t), CC(t)) are relative coordinates of the ball’s position where LR, AP and CC denote the left–right, anterior–posterior and craniocaudal directions, respectively. A slow-motion phase was defined as the ball positioned at f(t1)5¡20 [mm] and having a displacement in 1 s of 6.8 mm. A fast-motion phase was defined as the ball positioned at f(t2)50 [mm] and having a displacement in 1 s of 30.0 mm. The ball volume was obtained by applying a threshold to images, where the actual volume of the ball was 14 137 mm3 (54/36p6153). Volume error in the moving phantom was assessed as the percent ratio (%) of volume from images of the moving phantom to the actual volume. Scan conditions were 120 kV, 150 mA using a 0.5 mm 256 row detector and 15 s acquisition time. Reconstruction parameters were a voxel size of 0.35 mm6 0.35 mm60.35 mm, matrix size of 51265126256 and 0.35 mm reconstruction increment. The convolution kernel was a standard body kernel (FC10).

Animal study Image quality of RS-FDK was evaluated in an animal study in comparison with those for FS-FDK and HSFDK. Four domestic pigs were used to simulate a human model. The animals were aged 21–23 weeks, weighed 20– 25 kg and had a diameter and circumference of 130 mm and 590 mm, respectively. Although diameter and circumference were smaller than those of adult humans, pigs were selected owing to their ease of handling, and well-developed interlobular septa and anatomic structures that are similar to those of the human lung [43]. All animal procedures were approved by the IRB of the NIRS. The pigs were given an intramuscular injection of a mixture of 10 mg kg21 of ketamine hydrochloride (Sankyo Yell, Tokyo, Japan) and 7 mg kg21 of xylazine 2% (Bayer, Tokyo, Japan) and sedated and ventilated with a respiratory pump with isoflurane 2–2.5%. The breathing cycle was adjusted to 4.2–4.3 s. A video camera was focused on a reflective marker seal set on the interseptum to track respiratory motion and obtain the respiratory signal. Scan conditions were 120 kV, 200 mA using a 1.0 mm 128 row detector and 17 s acquisition time, and cine scan mode. Reconstruction parameters were a voxel size of

Figure 2. Photographs of the moving phantom. (a) The phantom was set 45 ˚ to the caudocranial direction and connected to a mechanically driven motor with speed adjustment. (b) The phantom contained a 30 mm diameter acrylic ball as the target volume placed on a moving table. 748



Figure 3. Motion functions for oblique motion (a). CT images of the 30 mm diameter acrylic ball in oblique motion in (b) the slow- and (c) fast-motion phases. Images in each of the vertical frames from top to bottom were reconstructed with FS-FDK, HSFDK and RS-FDK. The left–right, anterior–posterior and caudocranial directions are denoted as LR, AP and CC, respectively. Window level is –373 HU and window width is 1190 HU for all images.

0.47 mm60.47 mm60.47 mm, matrix size of 5126 5126216 and 0.47 mm reconstruction increment. The convolution kernel was FC10. Effective dose was estimated as 38.6 mSv (52.27 mSv s21617 s).

Results Phantom study Motion functions (a) and reconstructed images for (b) slow- and (c) fast-motion phases are shown in Figure 3. With regard to the slow-motion phase, FS-FDK images were severely degraded and distorted while those with HS-FDK did not seem spherical. In contrast, the RS-FDK images visualized the ball as spherical and were of better quality than the HS-FDK images due to their better temporal resolution. Furthermore, image quality with RS-FDK was the same for three different cross-sections. The British Journal of Radiology, September 2006

For the fast-motion phase, although the reconstructed images showed greater degradation than those of the slow-motion phase for all algorithms, image quality with RS-FDK was nevertheless better than with the others. In particular, the RS-FDK images show the ball as round in both coronal and sagittal sections. The magnitude of artefacts differed between the transverse and longitudinal sections. Volume using the threshold of 50% of CT number in the static image was 14 123 mm3, which nearly equals the actual value of 14 137 mm3. The volume ratios decreased as threshold CT number increased. The threshold was fixed at 50% of the difference between the CT number of the ball and the background CT number, because this gives the approximate actual volume of the ball. Figure 4 shows the volume percentage of the ball obtained from images of oblique motion. The volume percentage decreased from unity for all cases, with that by RS-FDK larger than with the other two, and that by 749


inhalation phases. The length of respiratory motion was approximately 17 mm in the CC direction. Coronal and sagittal images for each respiratory phase (pig 1) are shown in Figures 8 and 9, respectively. These images captured each respiratory phase exactly. The margins of the liver and the pulmonary vessels were sharper in the RS-FDK than in the FS-FDK and HS-FDK images. FS-FDK gave the worst image quality due to its low temporal resolution. For mid-exhalation and midinhalation, geometrical distortion of the pulmonary vessels and interlobular septa became visible in FSFDK and HS-FDK images, and the latter exhibited a streak artefact on the ribs in spite of their good temporal resolution. These results indicate that the RS-FDK algorithm provides the best image quality for all respiratory phases due to its good temporal resolution.

Discussion Figure 4. Volume percentage obtained from the axial image where the ball motion is in the oblique direction.

FS-FDK being the smallest. For the fast phase, the volume ratio by RS-FDK was 5% larger than that for HS-FDK and more than 40% larger than that by FS-FDK.

Animal study The phase of respiratory signals for all animals as sensed by the respiratory sensing system is shown in Figures 5 and 6, and respiratory periods for all animals are summarized in Table 1. Animal respiration was controlled with almost complete regularity by the ventilator, with only a few irregular periodic motions due to voluntary breathing. Breathing motion at peak inhalation was slightly faster than that at peak exhalation. This is similar to the behaviour seen for human respiratory signals, which are not sinusoidal; dwell time at peak exhalation is often longer than is shown by a sinusoidal signal. Figure 7 shows projection data of the pig in the AP direction at peak exhalation, mid inhalation and peak

In the present study, we developed the RS-FDK algorithm to allow the precise capture of anatomical structures under free breathing conditions, and compared results with those for two other reconstruction algorithms, FS-FDK and HS-FDK, in experiments using a moving phantom and in animal studies. Results showed that RS-FDK gives better temporal resolution than HSFDK and FS-FDK and the equivalent SNR as FS-FDK (Appendix 1). Although conventional respiratory-gated CT and 4DCT techniques provide wide CC coverage beyond the detector width for the same respiratory phase, the inclusion of motion artefact is not completely mitigated due to the relatively low temporal resolution. This disadvantage means that the actual position of a moving object in radiotherapy treatment planning CT is uncertain, as is the volume of a moving object relative to its actual volume. These uncertainties result in planning errors. Furthermore, respiratory motion during irradiation can cause the radiation beam to miss part of the target itself. To avoid these errors, PTV is defined with a margin of several centimetres added to the clinical target volume (CTV); but this in turn carries the risk that an excessively wide margin will unnecessarily irradiate normal tissue. It is therefore necessary to minimize the

Figure 5. Respiratory signal obtained from the external respiratory signal tracking system (pig 1). The solid line shows raw respiratory data and the broken one shows a spline fitting to the raw data.

750



Figure 6. Respiratory signal for four pigs; (a) pig 1, (b) pig 2, (c) pig 3 and (d) pig 4. Four respiratory cycles are overlapped to show reproducibility of each cycle.

PTV margin. As seen in the phantom study, the reconstructed images did not visualize the actual shape due to distortion or motion artefacts. These errors may result in the delivery of an insufficient dose to the target volume or an excess dose to normal tissue. HS and RS allow the precise imaging of a periodically moving object. The shorter acquisition time with HS, however, may lead to a lower SNR than that provided by a conventional acquisition time [44], necessitating careful Table 1. Respiratory cycle for four pigs Animal

1 2 3 4

Respiratory cycle (s)

1 4.24 4.20 4.21 4.13

Average SD

2 4.45 4.24 4.32 4.23

3 4.35 4.40 4.46 4.33

4 4.30 4.20 4.38 4.28

4.33 4.26 4.34 4.24


0.09 0.09 0.10 0.09

analysis of the trade-off between temporal resolution and image quality (e.g. artefact and image noise). In contrast, RS provides multiphase images with better temporal resolution than HS and the same SNR as FS, and may supply useful information for 4D radiation therapy planning. With the multiphase images we may make a dynamic treatment plan that enables irradiation with continually adjusted beam shape to match the respiratory phase. If such a dynamic treatment could be realised, it brings greater accuracy in radiation therapy than respiratory-gated radiation therapy, which irradiates at the most stable point, such as peak exhalation [45]. The limitation of RS is its higher patient dose. This occurs because the CT scans the same position continuously, and because the effective dose increases in proportion to acquisition time. Acquisition time should be minimized even for patients receiving radiation therapy. Because the respiratory period at small PS values such as 0.1 requires a longer scan time, rotation time should be adjusted so that total scan time is 751


Figure 7. Images of projection data in the anterior–posterior direction. (a) Peak exhalation, (b) mid inhalation and (c) peak inhalation. Window level is –373 HU and window width is 1190 HU for all images.

Figure 8. Coronal images obtained at four respiratory phases (peak inhalation, mid exhalation, peak exhalation and mid inhalation). Reconstruction increment is 0.4 mm and slice thickness is 0.4 mm. (a) FSFDK, (b) HS-FDK and (c) RS-FDK. Window level is –373 HU and window width is 1190 HU for all images.

Figure 9. Sagittal images obtained at four respiratory phases (peak inhalation, mid exhalation, peak exhalation and mid inhalation). Reconstruction increment is 0.4 mm and slice thickness is 0.4 mm. (a) FSFDK, (b) HS-FDK and (c) RS-FDK. Window level is –373 HU and window width is 1190 HU for all images.

decreased. If this is difficult, an alternative is to scan in one or two respiratory cycles, and thereafter to scan in appropriate respiratory phases only. Devices providing such functionality are essential to avoiding excessive patient dose. In the present study, scan time was minimized by adjusting respiratory period to 3.74 s for the phantom study and to 4.2–4.3 s for the animal study. Many authors have reported the limitations of the abdomen as an external surrogate for the respiratory phase, namely that longer CT scanning time under free breathing and the placement of the marker result in inconsistency between the position of the marker and 752

internal anatomy motion [16, 18, 46]. Lujan et al [6] have reported that the motion of the diaphragm due to respiration is predominantly in the CC direction and is periodic but asymmetric, with more time spent at the end of expiration than at the end of inspiration. This may lead to erroneous prediction of the dose delivered to the patient, and when examination time is prolonged, may result in the degradation of RS-FDK image quality. However, examination time in the present study was only 17 s, and the pigs were sedated and ventilated with a respiratory pump with an almost constant respiratory cycle. The relationship between the respiratory signal The British Journal of Radiology, September 2006


and target motion is therefore better correlated using the amplitude and cycle of the respiratory cycle. Since the breathing characteristics of patients are not always as reproducible as those of sedated pigs, the respiratory control discussed below is necessary when the RS-FDK is used in free-breathing patients. Moreover, it should be remembered that a tumour located in a lower lobe of the lung has a larger range of tumour motion than one in an upper lobe due to proximity to the diaphragm [47]. This may also impact RS-FDK in clinical use, which requires correction of the correlation between respiration and target motion, such as by control of the patient’s breathing by operator guidance [8–10, 16, 48, 49] or by means of an occlusion valve [11, 50, 51]. It will be necessary to control for this time dependency of respiratory asymmetry through suitable correction methods. With regard to the projection data sorting process, adaptation of spline curve fitting may induce overlapping or a lack of projection data at adjacent quadrant sections (A–D in Figure 1). However, the spline curve of respiratory function showed good regulation (Table 1). In this study, respiratory period for all animals was controlled with a ventilator to between 4.2 s to 4.3 s. Generally, if the period shifts between respirations, errors in the sorting process may lead to incorrect classification of the projection data into the wrong projection data section, and thereby result in image artefacts in the resultant RS-FDK. The same problem is seen in cardiac imaging with ECG gating [52, 53]. In this case, compensation of the projection data must be done by extending or shortening the range of the quadrant in both quadrant sections. Here, however, thanks to the use of the ventilator, the respiratory period was controlled and these errors were minimized, resulting in acceptable image quality. We have carried out clinical trials at NIRS using carbon-ion beams [54]. Gated irradiation is a useful method and several findings have been reported [20, 21]. Depth dose distributions of charged particles exhibit a strong Bragg peak at the very end of the maximum distance that a charged particle travels in a tissue (range), beyond which the dose very rapidly falls to zero [55]. Charged particle beam radiotherapy may therefore be valuable when the organs at risk are closely proximal to the target volume. However, the presence of a rangeshortening tissue inhomogeneity such as the soft tissue moving into the lung would not only reduce the target dose from the value in the stretched Bragg peak to essentially zero, but the dose would also put nearby organs at risk. Range is therefore an important factor in charged particle radiation therapy, even more so than in photon beam therapy. The use of RS-FDK in charged particle therapy as well as 4D radiation therapy should allow a significant increase in dose distribution accuracy and provide sufficient information to allow the minimization of the PTV margin in radiation therapy planning. We are now investigating these points.

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Appendix 1 The image noise ratios for FS-FDK, HS-FDK and RSFDK were calculated as follows. If sP is the noise magnitude (the standard deviation) of the measured projection data, noise magnitude of FS-FDK sFS, HS-FDK sHS and RS-FDK sRS can be calculated by noise propagation analysis by integrating the squared weights over the range of 2p: sffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi ffi ð 2p ð cm 1 2 2 ( ) sp dbdc sFS ~ {cm 2 0 sffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi ð ð 2p

cm

0

{cm

sHS ~

v2HS (c,b)sp 2 dbdc~1:36sFS

sffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi ffi ð 2p ð cm 1 2 2 sRS ~ ( ) sp dbdc~sFS 0 {cm 2 The British Journal of Radiology, September 2006


8 p b > > ) b [ ½0:2cm {2p sin2 ( > > > 4 cm {c < b [ ½2cm {2p,p{2c vHS (c,b)~ 1 > > > p pz2c {b > 2 m > ) b [ ½p{2c,pz2cm : sin ( 4 cm zc where vHS(c, b) is the Parker weighting, cm is the fan angle, b denotes the range of the projection angle and c denotes the ray-sum angle within a projection. RS also uses a uniform weight over the 2p of projections, as for FS. Figure A1 shows 3D images of weighting functions of FS-FDK, HS-FDK and RS-FDK. From the above analysis, the magnitude of the image noise for HS-FDK is 1.36


times larger than that for FS-FDK, whereas RS-FDK gives the same magnitude of image noise as FS-FDK.

Figure A1. 3D images of the weighting functions of (a) FSFDK and RS-FDK, and (b) HS-FDK. b denotes the projection angle and c denotes the ray-sum angle within a projection.

755


A mathematical model for patient skin dose assessment in cardiac catheterization procedures R E MORRELL,

PhD, MIPEM

and A T ROGERS,

MSc, MIPEM

Medical Physics and Clinical Engineering, City Hospital Campus, Nottingham University Hospitals NHS Trust, Hucknall Road, Nottingham NG5 1PB, UK ABSTRACT. A mathematical model has been developed for the assessment of patient skin doses from cardiac catheterization procedures. This uses exposure and projection data stored in the DICOM image files. Since these contain only information about the acquisition runs, a correction is needed to estimate and include the contribution from fluoroscopy. Maximum skin doses calculated by the model were found to correlate well with those measured on Kodak EDR2 film. Three methods for including the contribution from fluoroscopy were investigated, and all successfully identified patients receiving skin doses in excess of 1 Gy. It is hoped to automate this tool for routine assessment of skin doses in our cardiac catheterization laboratories.

Cardiac catheterization procedures can result in high radiation doses to the patient’s skin, sufficient to cause deterministic effects [1–6]. In order to comply with the recommendations of the International Commission on Radiological Protection [7], a robust method for routine assessment of patient skin dose is needed. Dose–area product (DAP) alone is not an adequate indicator of skin dose, since in many cases there is a poor correlation between the two quantities [8–12]. Slow radiographic film can be used to measure the dose distribution across the patient’s skin. However, its useful range is limited by its saturation point. Kodak EDR2 film (Eastman Kodak Company, Rochester, NY) has previously been used for skin dosimetry [12–14], and is the least sensitive of the slow films designed for portal imaging and quality control in radiotherapy. However, it saturates at around 1 Gy to 1.5 Gy, depending on the processing conditions applied [13–15]. There is now a growing range of ‘‘Gafchromic’’ films (International Specialty Products, Wayne, NJ), which saturate at higher radiation doses and do not require processing, but at present these are much more expensive than those in the Kodak range. Smaller detectors such as thermoluminescent dosemeters and scintillation detectors cover only a small area, so are liable to miss the region of maximum dose. In modern cardiac catheterization laboratories, information about the exposure parameters is stored in the DICOM file for each image series. The DICOM standard [16] specifies fields for data such as the number of frames in the series, angulation of the X-ray imaging unit, detector position and field size, imaging mode, beam energy and tube current. If these fields are populated, they allow the position and magnitude of the radiation dose to the patient’s skin to be estimated, for each image series. Since fluoroscopic images are not usually stored, this detailed dose information is only available for the acquisition runs. The purpose of this study was to develop a mathematical model to calculate the skin dose distribution across the patient’s back, using the exposure and projection data 756


stored in the image files. Maximum skin doses predicted by the model were compared with film dosimetry measurements for coronary angiography (CA) and percutaneous transluminal coronary angioplasty (PTCA) procedures. Three methods for including an estimated contribution from fluoroscopy were investigated.

Method The dose model was developed in Matlab version 7 (The Mathworks, Inc., Natick, MA). It was designed to extract and use the data from DICOM image files from an Integris H5000F C-arm imaging unit (Philips Medical Systems, Best, The Netherlands).

The dose model The exposure information is first extracted from all of the image files in the patient examination folder, and written to a structure array. An example of these data is shown in Table 1. Each row relates to one file, i.e. one image series. The columns contain the series or run number (Run), number of frames in that series (Frames), imaging protocol (Protocol), peak beam kilovoltage (kVp), tube current (mA), pulse width (ms), primary and secondary angles describing the orientation of the imaging unit in degrees (Ang1 and Ang2), source to image distance in millimetres (SID) and detector field size in millimetres (II). ‘‘Ang1’’ relates to left–right rotation of the gantry, and is positive for left anterior oblique and left lateral views, when the detector is towards the patient’s left-hand side. ‘‘Ang2’’ describes craniocaudal rotation, and is positive when the detector is closer to the patient’s head. Additional data required by the model are:

N

centreheight – the height of the centre of rotation of the C-arm from the floor (107 cm). The British Journal of Radiology, September 2006

A mathematical model for patient skin dose assessment Table 1. Example of exposure data extracted from DICOM image files Run

Frames

Protocol

1 2 3 4 5 6 7 8 9 10

62 52 54 53 53 48 61 44 8 29

12.5 12.5 12.5 12.5 12.5 12.5 12.5 12.5 12.5 12.5

FPS FPS FPS FPS FPS FPS FPS FPS FPS FPS

Coronary Coronary Coronary Coronary Coronary Coronary Coronary Coronary Coronary Coronary

kVp

mA

ms

Ang1

Ang2

SID

II

81 83 78 74 76 71 83 73 74 81

835 870 720 710 758 625 860 650 720 889

7 7 7 7 7 7 7 7 7 7

38 42 –27 –27 –1 –1 40 –30 –28 –28

22 –28 –20 27 –28 24 0 0 –20 –20

987 1031 1019 1019 1034 983 976 962 1037 1037

170 170 170 170 170 170 170 170 170 170

kVp, peak beam kilovoltage (kV); mA, tube current (mA); ms, pulse width (ms); Ang1, left-right angulation of imaging unit; Ang2, craniocaudal angulation of imaging unit; SID, focal spot to detector distance (mm); II, field size at detector face (mm).

N N

N

radius – the distance from the focal spot to the centre of the C-arm (77.5 cm). couchheight – the height of the couch above the floor. This can be adjusted during the examination and is not recorded by the imaging unit. A typical value (87 cm) was estimated from the recorded SIDs and measured field sizes for a selection of patient films from a previous dose survey [12]. ESDRphantom – the dose rate (ESDR) at the entrance surface of a 20 cm polymethyl methacrylate (PMMA) phantom, normalized for kVp, mA, pulse time (ms), frame rate and focus to PMMA distance in centimetres (fsd), as shown in Equation (1). This has a value of 5.9761025 mGycm2 (kVp2 mA ms)21 per frame. ESDRphantom ~

ESDR|fsd2 ð1Þ kVp 2 |mA|ms|frame rate

The entrance dose rate to the phantom was measured using a Radcal 9010 series dosemeter, with a 60 cm3 ionization chamber (Radcal, Monrovia, CA), as shown in Figure 1. The standard clinical acquisition mode (12.5 FPS Coronary) was used. This mode does not employ any copper filtration. Since the imaging unit stores no information about couch position, the model assumes that the couch remains at a fixed height throughout the procedure, and that no panning is used. The dose distribution is calculated at 1 mm intervals in a plane at the height of the couch top, that extends from 230 cm to +30 cm in the x (cross-couch) direction and 220 cm to +20 cm in the y (parallel to couch axis) direction. Positive x is defined towards the right-hand side of the patient, and positive y towards the patient’s head, so that the dose map appears as though viewing the patient from behind. For each acquisition run, the position of the focal spot (xspot, yspot, zspot) is calculated in Cartesian coordinates, using Equations (2) to (4). xspot ~radius| sinðAng1Þ| cosðAng2Þ

ð2Þ

yspot ~{radius| sinðAng2Þ

ð3Þ

zspot ~centreheight{radius| cosðAng1Þ| cosðAng2Þ ð4Þ The British Journal of Radiology, September 2006

The distance (ffd, in centimetres) from the focal spot to each point in the plane is then calculated using Pythagoras’ law. The dose at each point on the film, in the absence of beam collimation, would be: dose~ESDRphantom |

kVp 2 |mA|ms|frames ffd2

ð5Þ

where kVp2, mA, ms and frames are taken from the DICOM file for that particular run. Beam limitation is calculated assuming square radiation fields, and ignoring any secondary collimation applied by the user. The angle (h) between the beam central axis and its collimated outer edge is calculated from the field size and SID. ‘‘SIDcoll’’ is the radius of a sphere centred at the focal spot that passes through the four corners of the collimated field, at the detector face. It is calculated from field size and SID, using Pythagoras’ law. The positions of the field corners at the detector face are calculated in Cartesian coordinates, using Equations (6) to (8).

Figure 1. Set-up for determining the entrance dose rate to phantom. PMMA, polymethyl methacrylate; fsd, as used in Equation (1). 757


ycoll ~yspot zSIDcoll | sinðAng2+hÞ

ð6Þ

xcoll ~xspot {SIDcoll | sinðAng1+hÞ| cosðAng2+hÞ ð7Þ zcoll ~zspot zSIDcoll | cosðAng1+hÞ| cosðAng2+hÞ ð8Þ They are then translated from the detector face onto the plane of the couch top, by means of scaling. xcoll {xspot xcouch ~xspot z | couchheight{zspot zcoll {zspot

ð9Þ

ycoll {yspot | couchheight{zspot zcoll {zspot

ð10Þ

ycouch ~yspot z

patient’s back [12]. The current study employed the dose model to calculate a dose map for each of the patients included in the survey. The maximum doses determined by the two methods were compared. The calculated skin dose maps were first compared visually with the films. Patients were excluded from the study if the region of maximum dose predicted by the model was outside the area of the film. For the subset of films that showed no saturation, the Pearson correlation coefficients between calculated and measured maximum doses were computed. This subset included 14 CA and 20 PTCA procedures.

Results

To apply the collimation, the values of the dose array are set to zero at all locations outside the trapezium formed by these four points. The completed dose maps from each run are summed to give a total dose array, which is then displayed as a filled contour map. The model described above considers only the acquisition run data, which is stored in the image files. In clinical practice, some examinations involve extended fluoroscopy times, and where fluoroscopy is performed primarily in one projection this can raise the maximum skin dose considerably. Three options for estimating and including the contribution from fluoroscopy were investigated. 1. Dose–area product. The DAP was calculated from the exposure parameters for each acquisition run, and summed to give acquisition DAP for the whole procedure. The final dose map was multiplied by the ratio of displayed total DAP to calculated acquisition DAP. 2. Fluoroscopy time. The entrance dose rate to the 20 cm PMMA phantom in the standard fluoroscopy mode (Low Fluoro), at the 23 cm field size, with an SID of 100 cm and fsd of 67 cm has been measured as approximately 40 mGy min21. The fluoroscopy time (in minutes) was divided by the number of runs, multiplied by this dose rate, and added to the dose array for each run, before applying beam collimation. 3. Concentration factor. An attempt was made to relate maximum skin dose to DAP for individual procedures, by means of a ‘‘concentration factor’’. This was simply the ratio of the maximum calculated skin dose, to the sum of doses at all points in the dose map. Peak skin dose was predicted for each patient using the product of DAP and concentration factor, together with a constant that forced the regression line between predicted and measured doses to have a gradient of one.

Figure 2 shows an example of the data output from the skin dose model. A visual comparison with the corresponding dosimetry film (Figure 3) shows that the region of maximum dose has been correctly identified. The fields visible on the calculated dose map can be broadly matched with some of those appearing on the film. However, fields arising from purely fluoroscopic exposures are seen only on the film, since the model has no projection data for fluoroscopy. Figure 4 shows the calculated doses for procedures resulting in film saturation, for each version of the dose model. The film saturates at about 1 Gy, so saturation indicates a skin dose of 1 Gy or more [15]. The most basic form of the dose model, which includes only acquisition data, generally predicted doses of at least 800 mGy for these procedures. However, in one case, a dose of only 332 mGy was computed. Although this patient had a maximum skin dose of at least 1 Gy and may therefore be at risk of deterministic effects, the model would not have identified her as having received a dose that may be of concern. On applying any of the correction methods to include the contribution from fluoroscopy, the calculated dose to this patient increased to more than 1 Gy. When corrected by fluoroscopy time or concentration factor, the model successfully identified all procedures that resulted in film saturation. When adjusted by the ratio of displayed

Comparison of calculated and measured doses A dose survey has previously been carried out in our cardiac catheterization laboratory, using Kodak EDR2 film to measure the skin dose distribution across the 758

Figure 2. Example output from dose model. The British Journal of Radiology, September 2006

A mathematical model for patient skin dose assessment

Figure 5. Calculated versus measured doses, using only acquisition data.

Figure 3. Dosimetry film for the patient whose calculated skin dose map is shown in Figure 2.

to calculated DAP, the model predicted one of these patients to have a maximum skin dose of 928 mGy, and all others to have maximum doses of more than 1 Gy. Figures 5–8 show calculated versus measured doses for all the procedures that did not result in film saturation, i.e. for which measured doses were less than 1 Gy. The error bars show the expected uncertainty in film dosimetry measurements [12]. In each case, a trend line has been fitted, that passes through the origin. The equation of the trend line and the square of the Pearson correlation coefficient are shown on each figure. When using only the acquisition data, the gradient of the trend line was 0.735 (Figure 5), suggesting that the model underestimates doses by 26% on average. There was a strong correlation between calculated and measured doses (R250.672). On using DAP or fluoroscopy time to estimate the contribution from fluoroscopy the gradient of the trend line was closer to 1 (Figures 6 and 7). The strength of the correlation also increased (R250.708 and R250.716, respectively). The product of concentration factor and DAP was multiplied by 181 000, to force the gradient of the trend

Figure 4. Calculated doses for procedures that resulted in film saturation, using each version of the dose model. Fluoro Correction 1 uses DAP, 2 uses fluoroscopy time, 3 uses concentration factor to estimate the contribution from fluoroscopy. The British Journal of Radiology, September 2006

Figure 6. Calculated versus measured doses, using dose– area product (DAP) to estimate the contribution from fluoroscopy.

line to 1 (Figure 8). This method gave the strongest correlation between calculated and measured doses (R250.735).

Discussion For doses of up to 1 Gy, maximum skin doses computed using the model correlate well with those measured using Kodak EDR2 film. Unlike film dosimetry, the model has no limit on the magnitude of the doses that can be evaluated. Using only the acquisition data stored in the DICOM files can lead to large errors in dosimetry for examinations

Figure 7. Calculated versus measured doses, using fluoroscopy time to estimate the contribution from fluoroscopy. 759


develop real-time skin dose monitoring software that utilizes much of this information, none is currently available for purchase. As with film dosimetry, the model considers only those radiation beams that pass through the plane of the couch, and ignores any contributions to skin dose from lateral views. A potential improvement would be to use a three-dimensional model of the patient to estimate skin doses over the whole surface of the thorax. However, this would require couch positioning data, to achieve any degree of accuracy.

Figure 8. Calculated versus measured doses, using concentration factor to estimate the contribution from fluoroscopy.

involving long fluoroscopy times and few projections. Three options for including an estimated contribution from fluoroscopy were explored. Each improved the accuracy of the model for procedures involving a large proportion of fluoroscopy, and predicted doses of at least 928 mGy for all procedures resulting in film saturation. The method using displayed DAP together with a ‘‘concentration factor’’ gave the strongest correlation between calculated and measured doses. Using this method, 95% of calculated doses were within ¡ 270 mGy of measured doses, for films showing no saturation. Each method for including the contribution from fluoroscopy requires additional information (either DAP or fluoroscopy time) that is not stored in the image files and must therefore be obtained from another source. At present, displayed DAP and fluoroscopy time for all procedures are manually recorded in a book, and in an Oracle database (Oracle Corporation, Redwood Shores, CA). In the near future, these may be automatically stored in a new catheterization laboratory information system. It is hoped that clinical application of the dose model can eventually be completely automated – to extract the relevant data from the image files and obtain the fluoroscopy time from the information system or database as each examination is completed, to calculate the skin dose distribution and alert a member of staff if the maximum dose exceeds a certain threshold. This would enable staff to follow up at-risk patients by examining their skin and warning them of potential effects before they leave the hospital. The accuracy of the model is limited by a number of unknown variables, about which assumptions have had to be made. No exposure factors are stored for fluoroscopy, as has been previously discussed, necessitating an estimation of the contribution to total dose from the fluoroscopic parts of the procedure. The actual couch position for each procedure is not known, so it is necessary to assume a certain couch height, and to ignore any horizontal movement. No information about secondary collimation or the use of the wedge filter is available, so these must be assumed not to have been used. Improved accuracy is dependent on manufacturers building in the facilities to make this information available. Whilst den Boer et al [17] and Chugh et al [18] have worked with equipment manufacturers to 760

Conclusion A dose model has been developed to calculate the skin dose distribution across the patient’s back from cardiac catheterization procedures. This utilizes the exposure and projection data stored in the DICOM image files, as well as DAP or fluoroscopy time. Maximum doses calculated by the model correlated well with those measured using Kodak EDR2 film. After applying a correction to include the estimated contribution from fluoroscopy, the model successfully identified patients receiving skin doses in excess of 1 Gy. It is hoped to automate the dose model for use as a routine dosimetry tool in the cardiac catheterization laboratory. The accuracy of the model is limited by several unknown variables that are not recorded by the imaging system. Improvement of dosimetric accuracy is dependent upon manufacturers developing methods for storing and accessing this information.

Acknowledgments We are grateful to Dr Nick Gibson for his assistance with the Matlab code, and to Prof. Alan Perkins for helpful discussions.

References 1. Kawakami T, Saito R, Miyazaki S. Chronic radiodermatitis following repeated percutaneous transluminal coronary angioplasty. Br J Dermatol 1999;141:150–3. 2. D’Incan M, Roger H, le Boudec MCF, Souteyrand P. Radiodermatitis following cardiac catheterization. Arch Dermatol 1997;133:242–3. 3. Dehen L, Vilmer C, Humiliere C, Corcos T, Pentousis D, Ollivaud L, et al. Chronic radiodermatitis following cardiac catheterisation: a report of two cases and a brief review of the literature. Heart 1999;81:308–12. 4. Shope TB. Radiation-induced skin injuries from fluoroscopy. Radiographics 1996;16:1195–9. 5. Vano˜ E, Goicolea J, Galvan C, Gonzalez L, Meiggs L, Ten JI, et al. Skin radiation injuries in patients following repeated coronary angioplasty procedures. Br J Radiol 2001;74: 1023–31. 6. Vano˜ E, Arranz L, Sastre JM, Moro C, Ledo A, Garate MT, et al. Dosimetric and radiation protection considerations based on some cases of patient skin injuries in interventional cardiology. Br J Radiol 1998;71:510–6. 7. International Commission on Radiological Protection. Report 85: Avoidance of radiation injuries from medical interventional procedures. Ann ICRP 2000;30:45–7. 8. van de Putte S, Verhaegen F, Taeymans Y, Thierens H. Correlation of patient skin doses in cardiac interventional


A mathematical model for patient skin dose assessment

9.

10.

11.

12.

13.

radiology with dose-area product. Br J Radiol 2000;73: 504–13. Vano˜ E, Gonzalez L, Ten JI, Fernandez JM, Guibelalde E, Macaya C. Skin dose and dose-area product values for interventional cardiology procedures. Br J Radiol 2001;74:48–55. Waite JC, Fitzgerald M. An assessment of methods for monitoring entrance surface dose in fluoroscopically guided interventional procedures. Radiat Prot Dosim 2001;94:89–92. Delichas MG, Psarrakos K, Giannoglou G, MolyvdaAthanasopoulou E, Hatziioannou K, Papanastassiou E. Skin doses to patients undergoing coronary angiography in a Greek hospital. Radiat Prot Dosim 2005;113: 449–52. Morrell RE, Rogers AT. Kodak EDR2 film for patient skin dose assessment in cardiac catheterization procedures. Br J Radiol 2006;79:603–7. Guibelalde E, Vano˜ E, Gonzalez L, Prieto C, Fernandez JM, Ten JI. Practical aspects for the evaluation of skin doses in


14.

15.

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17.

18.

interventional cardiology using a new slow film. Br J Radiol 2003;76:332–6. Vano˜ E, Prieto C, Fernandez JM, Gonzalez L, Sabate M, Galvan C. Skin dose and dose-area product values in patients undergoing intracoronary brachytherapy. Br J Radiol 2003;76:32–8. Morrell RE, Rogers A. Calibration of Kodak EDR2 film for patient skin dose assessment in cardiac catheterization procedures. Phys Med Biol 2004;49:5559–70. National Electrical Manufacturers Association. Digital Imaging and Communications in Medicine (DICOM). Part 6: Data Dictionary (PS 3.6-2006). Rosslyn, Virginia: National Electrical Manufacturers Association, 2006. den Boer A, de Feijter PJ, Serruys PW, Roelandt JRTC. Realtime quantification and display of skin radiation during coronary angiography. Circulation 2001;104:1779–84. Chugh K, Dinu P, Bednarek DR, Wobschall D, Rudin S, Hoffmann K, et al. A computer-graphic display for realtime operator feedback during interventional x-ray procedures. Proc Soc Photo-Opt Instrum Eng 2004;5367:464–73.

761


Nurse-led central venous catheter service: Christie experience K GOPAL,

MRCP, FRCR,

L FITZSIMMONS,

RGN

and J A L LAWRANCE,

MRCP, FRCR

Christie Hospital, Wilmslow Road, Manchester M20 4BX, UK ABSTRACT. The aim of this study was to evaluate the success and complication rate of Nurse-led subclavian central venous catheter (CVC) insertion using the landmark technique. A prospective study was performed on all subclavian CVC insertion between 13/01/03 to 01/07/03. Data recorded included indication for catheter insertion, type of catheter, complications during the procedure and patient satisfaction. A total of 348 subclavian cuffed tunnelled catheters were inserted over the study period. The age group ranged from 31 years to 84 years with a mean age of 53 years. This included 192 women and 156 men. The majority of CVC insertions were through the right subclavian (79%) and were single lumen (76%). In total, complications were encountered in 48 patients (14%). These included misplaced tip in 29 (8%), arterial puncture in 16 (4%), pneumothorax in 3 (1%) and the procedure failed in 3 (1%). Of these multiple complications were seen in 3 (1%). No interventions were required for the pneumothoraces or for the arterial punctures. In conclusion, nurse-led subclavian CVC placements using the landmark technique are both safe and effective.

Venous access is a critical issue in the care and management of patients with a wide variety of malignancies. The insertion of tunnelled central venous catheters (CVC) for patients requiring long-term venous access is now common. In the NHS, approximately 200 000 CVC are inserted in adult patients per year [1]. This is partly due to increased usage in cancer patients. CVC are inserted by surgeons, anaesthetists, interventional radiologists, medical oncologists and nurses. The British Committee of Standards in Haematology (BCSH) guidelines [2] state that insertion should be performed by experienced operators, regardless of speciality. In our centre, these lines are inserted by a team of experienced nurses. Nurses undergo a supervised training programme for 5 weeks (discussed below) in which they not only observe but perform at least 50 CVC insertions. They are then formally assessed to determine whether they may perform the line insertions independently. Various studies have been published supporting the extension of nurse’s roles in this area. One of the main advantages of nurse-led CVC insertion is that the procedure can be done at the bedside. Others include reduced waiting time for patients, stable workforce and successful outcome due to increased frequency of practice [3, 4]. Although central venous access devices are clearly advantageous with respect to delivery of therapy, their placement and maintenance is not without potential complications. McBride and colleagues [5] demonstrated that there is a steep operator learning curve and the complication rate improves notably with experience. In recent years, The National Institute of Clinical Excellence (NICE) has recommended the use of ultrasound guided placement of CVC [6]. However, this has been received with a mixed response form various specialities. At our institution, a large oncology centre in the UK, central venous access is performed by the Address correspondence to: Dr J A L Lawrance.

762

Received 11 July 2005 Revised 21 December 2005 Accepted 8 February 2006 DOI: 10.1259/bjr/79794134 ’ 2006 The British Institute of Radiology

landmark technique led by a team of nurse specialists. We undertook this study to evaluate the success and complication rate of nurse-led CVC insertion via the subclavian route.

Materials and methods We perform approximately 950 nurse-led CVCs per year. These include both CVC and peripherally inserted central catheters (PICC). A prospective study was performed on all CVC insertion between 13/01/03 and 01/07/03. The following information was recorded, including: diagnosis and indication for catheter insertion, type and position of the catheter, complications during the procedure and patient satisfaction.

Insertion technique Pre-procedure investigations include full blood count and coagulation profile. Local anaesthesia (5–10 ml 2% lidocaine) is used to anaesthetise the skin. Prophylactic intravenous antibiotics are not routinely administered [7]. Catheters are inserted blindly using the landmark technique. The right subclavian vein is used unless contraindicated, e.g. due to thrombosis, infection, left pulmonary compromise or right sided mastectomy and radiotherapy. The catheters are secured with ethilon sutures at the insertion site. The intended catheter tip position is the distal superior vena cava (SVC), SVC/ right atrial junction or right atrium [8]. Following the insertion, a chest radiograph is performed to check catheter position and identify potential complications such as pneumothorax. As most of the patients received continuous infusion of chemotherapy as out-patients, they are taught to look carefully for signs of catheter infection, blockage or accidental removal. Written The British Journal of Radiology, September 2006

Nurse lead venous catheter service

instructions are provided for reference. Patients were then asked to fill out an audit questionnaire regarding the procedure, which included patient satisfaction, waiting time, etc.

Results 348 CVC were inserted over this 6 month period. Similar numbers of PICC lines were also inserted, but these were excluded from the study group. Indications for catheter insertion include carcinomas such as breast, gastrointestinal, leukaemia, lymphoma, ovarian and multiple myeloma (Table 1). The age group ranged from 31 years to 84 years, with a mean age of 53 years. This included 192 women and 156 men. The majority of the patients were of normal build except for 16% who were either obese or emaciated, which was calculated based on their body mass index (Table 2). The main indication for CVC insertion in patients was for chemotherapy. Others include bone marrow transplant, total parental nutrition, difficult venous access and combination of the above (Table 3). The right subclavian route was the preferred route of insertion unless contraindicated (Table 4). All insertions Table 1. Subset of disease groups Subset

Frequency (%)

Gastrointestinal Breast Leukaemia Lymphoma Ovarian Multiple myeloma Other

190 (54%) 38 (11%) 24 (7%) 24 (7%) 12 (3%) 8 (3%) 52(15%)

were done using the blind technique under local anaesthesia. Sedation was not routinely administered and in our study only 78 patients (22.4%) received sedation. Two patients had to have the procedure under general anaesthesia. Single, double or triple lumen catheters were used. Although multi-lumen catheters facilitate concurrent administration of different medications and fluids, we preferred the usage of single lumen catheters to reduce the rate of infection (Table 5). Nurseled CVC insertions are usually performed at the bedside in the Day Case Unit (Table 6). The procedures are usually performed as day cases. CVC catheters were inserted with a single or double needle pass in the majority of patients (Table 7). Insertion related complications were noted in a total of 48 patients (14%). There were three pneumothoraces (1%), and 16 arterial punctures (4%). None of the former required chest drains and none of the latter caused significant haemodynamic problems, such as hypotension or tachycardia. In three patients there was failure of catheter insertion and an alternative approach, i.e. internal jugular vein was used. Chest radiographs confirmed tip malposition in 29 patients (8%). In cases where the tip is malpositioned, catheters are repositioned under fluoroscopy guidance by the nurse specialists themselves. This can usually be achieved by a technique developed by one of the authors (JALL) consisting of a combination of a forced saline injection and hyperventilation, the combination of which will usually reposition a contralateral line tip to the SVC. In cases where this fails, the tip of the CVC line is withdrawn and manipulated to lie within the distal SVC/right atrium fluoroscopically. In three patients (1%) multiple complications were seen (Tables 8 and 9). The latter patients were hospitalized, but all complications resolved conservatively. 327 patients (94%) tolerated the procedure very well. Minor problems, e.g. discomfort Table 5. Type of line inserted

Table 2. Patients’ body habitus

Lumen

Body habitus

Frequency (%)

Normal Obese Emaciated Not known

287 (82.5%) 37 (10.6%) 21 (6%) 3

Table 3. Indication for central venous catheter (CVC) insertion Indication

Frequency (%)

Chemotherapy Bone marrow transplant Difficult venous access Total parental nutrition Other

312 18 5 7 6

(90%) (5%) (1.5%) (2%) (1.5%)

Table 4. Site of insertion Site

Frequency (%)

Right subclavian Left subclavian Not known

274 (79%) 71 (20%) 3 (1%)


Single Double Triple Not known

264 (76%) 71 (20%) 7 (2%) 6 (2%)

Table 6. Place of insertion Location

Frequency (%)

Bed side Radiology department Not known

342 (98%) 2 (0.5%) 6 (1.5%)

Table 7. Number of attempts to cannulate the subclavian vein No. of passes

Frequency (%)

1 2 3 4 Several Not known

168 (48%) 67 (19%) 27 (8%) 6 (2%) 45 (13%) 35 (10%)

763

K Gopal, L Fitzsimmons and J A L Lawrance Table 8. Summary of insertion related complications Summary

Frequency (%)

No. of patients No complications Complications

348 (100%) 299 (86%) 48 (14%)

Table 9. Subdivision of complications Complication

Frequency (%)

Misplaced tip Arterial puncture Pneumothorax Failed procedure Multiple complications

29 (8%) 16 (4%) 3 (1%) 3 (1%) 3 (1%)

and pains were reported in 6%. Long term follow-up of these catheters, in relation to thrombosis, was not assessed in this study as this has been done in a previous publication from our institution [8].

Discussion Initially implemented by Broviac et al [9] in 1973 and subsequently modified by Hickman et al [10] in 1979, indwelling central venous access devices have revolutionized the care and treatment of the oncology patient. Various studies have been published supporting the extension of the nurse’s role in this area. Indeed, BCSH guidelines state that insertion should be performed by experienced operators regardless of speciality [2]. In our institution, all nurses undergo intensive training for 5 weeks. This is divided into Phase 1 and Phase 2, which last for 1 week and 4 weeks, respectively. During Phase 1, the nurses not only observe the insertion of CVC but also become familiar with all types of CVC access devices, their uses and the advantages and disadvantages associated with the various types of catheters. Later in Phase 2 they undertake at least 50 CVC insertions under direct supervision and with the assistance of a mentor. Concluding the training period, a formal documented assessment of competency is undertaken by an independent assessor and, once satisfactory, nurses are then allowed to perform independently. Nurse placements of CVC at the patient’s bedside mean that insertions can take place at the optimal time in a patient’s management as waiting lists for theatres or Xray suites are no longer barriers to insertion, which helps to reduce the patients waiting time for CVC insertion. The nurse specialists are a stable workforce and the frequency of practice correlates with successful outcome [3, 4, 11]. In addition, the numbers of patients encountered are too great to be accommodated in the X-ray department with the limited resources available. Nurses are also being trained to use fluoroscopy, which they use to reposition CVC lines if blind insertions lead to misplaced catheter tips. In our centre, CVC are inserted by a team of experienced nurses under blind percutaneous placement. Image guidance is not used for venous puncture. One of the authors (JALL) has trained the team in the use of fluoroscopy and the team has free access to the interventional suite when uncertain about CVC 764

placement or in patients in whom they suspect potential problems might arise. In rare cases, patients are referred for radiologically inserted internal jugular lines, using ultrasound and fluoroscopic guidance. The right subclavian vein is punctured and the operator advances the catheter blindly to the lower part of the SVC/right atrium. Fluoroscopy is usually only used to reposition misplaced catheter tips. We have used the subclavian route historically and our CVC service was started before NICE guidelines and before the widespread use of ultrasound guidance for central line insertion. Due to the high success and acceptable complication rates over the years, this service has been continued. Given the high number of line insertions we perform and our limited resources, we have continued with this service. Some advantages of bedside CVC insertion are the fact that they can be inserted on the ward and patients may be less anxious of the ward environment compared with operating theatres or X-ray suites, as well as having cost implications. The main disadvantage is risk of catheter tip misplacement. As the use of tunnelled CVC is increasing, it is important to recognize that their use is not without complications. The experience of the operator is an important factor in keeping complications to a minimum. In our study, the total incidence of insertion related complications was 14%, which compares favourably with other reports [12, 13]. The majority of these complications was related to catheter tip malposition (8%) with data from other studies ranging between 12% and 32% [13, 14]. Early complications such as pneumothorax (1%) and arterial puncture (4%) are also similar to previous reports in the literature [15, 16]. It is said that the risk increases with blind punctures where the operator is solely reliant on anatomical landmarks for the venipuncture. However, in our case the risks were well within the limitations and all the pneumothoraces encountered were small and none required chest drain insertion. In our study, 1% of procedures failed and had to be abandoned [17] with multiple complications seen in a similar number [18], which are similar in number to those reported earlier. Sedation is not routinely administered for this procedure and when used we perform this within the remit of The Safe Sedation Policy prescribed by the trust. The hospital policy on sedation is widely published in our department and intranet. As part of the training, the nurses undergo a formal training on sedation and, in addition, all nurse practitioners have a Post Graduate Diploma assessment on their ability to use drugs and side effects. All the procedures are performed under indirect supervision of a doctor in the outpatient or radiology department. To reduce the risk of infection, we prefer the insertion of single lumen catheters (76%) as it is apparent that the greater the number of catheter lumens, the greater the potential for sepsis [19]. The total life span of the catheters depends on need. For example, in haematology patients they can remain in situ for 12 months from diagnosis to post-transplantation and in solid tumour patients receiving chemotherapy, approximately 4 months. With respect to immediate complications involved in the placement of CVC, our results are comparable with those described previously by several investigators. Although the cost for nurse-led CVC was not within the remit of the study, we believe that the cost The British Journal of Radiology, September 2006

Nurse lead venous catheter service

of the procedure as performed by nurses, including the complications, should be less than radiologically inserted lines. Given the European Working Time Directive and the move to roll redesign, this is an excellent example of the successful deployment of nurses to perform what was previously the doctors’ role. In 2002, NICE recommended the use of ultrasound for the insertion of CVC into the internal jugular vein in adults and children [6]. This has been received with a mixed response across various specialities. NICE had largely concentrated on the complication of inadvertent arterial puncture and had made the above recommendation. Ultrasound does reduce the risk, but it is usually small [20]. Two recent studies make clear that ultrasound guidance offered incomplete protection against arterial injury [21] and did not improve the result of right IJV cannulation, compared with a meticulous landmark based technique [22]. NICE had considered only one randomized control study that analysed the effect of ultrasound and landmark technique in the placement of subclavian CVC catheters. The operators in both groups were relatively inexperienced with failure rates as high as 55% in the landmark technique, which is far higher that reported in trials (9–19%) [6]. Muhm concluded that ultrasound guidance improved the number of attempts per cannulation and successful first attempts for catheterization of the internal jugular vein, but not the subclavian approach [23]. In addition, the Department of Health has suggested that the subclavian route was associated with less infection [24]. Furthermore, none of the trials quoted by NICE involved nurses [6]. NICE admits that the landmark method is safe in experienced hands and operators should maintain their ability to use the landmark method. The Royal College of Anaesthetists agrees and advises that utilization of the landmark method is still an acceptable alternative whether ultrasound is available or not [25].

Conclusion The use of CVC is now well established in the treatment of cancer patients. With its increasing demand, manpower and facilities are its main limitation. Our study has indicated that complication rates are well within acceptable limits using a landmark technique for the subclavian route. Operator experience, irrespective of the speciality, is the key to limiting the complication rate and regular audits are necessary to prove their efficiency. Nurse-led CVC placements using the landmark technique are both safe and effective.

References 1. Elliot TSJ, Faroqui MH, Armstrong RF, Hanson GC. Guidelines for good practise in central venous catheterisation. J Hosp Infect 1994;28:163–76. 2. BCSH Working party. BCSH guidelines on the insertion and management of central venous lines. Br J Haematol 1997;98:1041–7. 3. Fitzsimmons CL, Gilleece MH, Ranson MR, Wardley A, Morris C, Scarffe H. Central venous cathter placement: extending the role of nurses. J R Coll Physicians Lond 1997;31:533–4.


4. Hamilton H, O’Byrne M, Nicholai L. Central lines inserted by clinical nurse specialists. Nurs Times 1995;91:38–9. 5. McBride KD, Fisher R, Warnock N, Winfield DA, Reed MW, Gaines PA. A comparative analysis of radiological and surgical placement of central venous catheters. Cardiovasc Intervent Radiol 1997;20:17–22. 6. Guidance of the use of ultrasound locating devices for placing central venous catheters. National Institute for Clinical Excellence Technology appraisal guidance No 49, September 2002. National Institute for Clinical Excellence. 7. Ranson MR, Oppenheim BA, Jackson A, Kamthan AG, Scarffe JH. Double-blind placebo controlled study of vancomycin prophylaxis for central venous catheter insertion in cancer patients. J Hosp Infect 1990;15:95–102. 8. Cadman A, Lawrance JA, Fitzsimmons L, et al. To clot or not to clot? That is the question in central venous catheters. Clin Radiol 2004;59:349–55. 9. Broviac JW, Cole JJ, Scribner BH. A silicone rubber atrial catheter for prolonged parental alimentation. Surg Gynecol Obstet 1973;136:602–6. 10. Hickman RO, Buckner CD, Clift RA, et al. A modified right atrial catheter for access to the venous system in marrow transplant recipients. Surg Gynecol Obstet 1979;148:871–5. 11. Hamilton HC. Selecting the correct intravenous device: nursing assessment. Br J Nurs 2000;9:968–78. 12. Ray S, Stacey R, Imrie M, Filshie J. A review of 560 Hickman catheter insertions. Anaesthesia 1996;51:981–5. 13. DeJong PCM, Meyenfeldt MR, Rouflart M, Wesdorp RIC, Soeters PB. Complications of central venous catheterization of the subclavian vein: the influence of a parenteral nutrition team. Acta Anaesth Scand 1985;Suppl. 81:48–52.. 14. Dewey JC, Robert WH. Aberrant locations and complications in initial placement of subclavian vein catheters. Arch Surg 1984;119:293–5. 15. Gray RR. Radiological placement of indwelling central venous lines for dialysis, TPN and chemotherapy. J Interv Radiol 1991;6:133–44. 16. McBride KB, Warnock N, Fisher R, Reed NW, Winfield DA, Gaines PA. A comparative review of surgically and radiologically placed Hickman catheters. Cardiovasc Interv Radiol 1994;17:S122. 17. Rosen M, Latto P, Ng S. Percutaneous central venous catheterisation. London: W.B. Saunders, 1992. 18. Mansfield PF, Hohn DC, Fornage BD, Gregurich MA, Ota DM. Complications and failures of subclavian-vein catheterization. N Engl J Med 1994;331:1735–8. 19. Eastridge BJ, Lefor AT. Complications of indwelling venous access devices in cancer patients. J Clin Oncol 1995;13:233–8. 20. Chalmers N. NICE should reconsider its guidance on the use of ultrasound for placing central venous catheters, (18 December 2002) Rapid response for Muhm. BMJ 2002;325:1373–4. 21. Augoustides JG, Diaz D, Weiner J, et al. Current practise of internal jugular venous cannulation in a University anaesthesia department. J Cardiothor Vasc Anesth 2002;16:567–71. 22. Hayashi H, Amano M. Does ultrasound imaging before puncture facilitate internal jugular vein cannulation? J Cardiothor Vasc Anesth 2002;16:572–5. 23. Muhm M. Ultrasound guided central venous access. BMJ 2002;325:1373–4. 24. Pratt RJ, Pellowe CM, Loveday HP, et al. The epic project: developing national evidence-based guidelines for preventing healthcare associated infections. Phase1: Guidelines for preventing hospital acquired infections. J Hosp Infect 2001;47(Suppl):S1–82. 25. Ball DR. Evidence based central venous access. Rapid response for Muhm. BMJ 2002;325:1373–4.

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Breast movement during normal and deep breathing, respiratory training and set up errors: implications for external beam partial breast irradiation 1

S CHOPRA,

MD, DNB,

1

K A DINSHAW,

DMRT, FRCR,

1

R KAMBLE,

MSc

and 2R SARIN,

MD, FRCR

1

Radiation Oncology, Tata Memorial Hospital, Mumbai, Maharashtra and 2ACTREC, Kharghar, NaviMumbai, Maharashtra, India

ABSTRACT. This study was designed to evaluate interfraction and intrafraction breast movement and to study the effect of respiratory training on respiratory indices. Five patients were immobilized in supine position in a vacuum bag and three-dimensional set up errors, respiratory movement of the breast during normal and deep breathing, tidal volume and breath hold time were recorded. All patients underwent respiratory training and all the respiratory indices were re-evaluated at the end of training. Cumulative maximum movement error (CMME) was calculated by adding directional maximum set up error and maximum post training movement during normal breathing. The mean set up deviation was 1.3 mm (SD ¡ 0.5 mm), 1.3 mm (SD ¡ 0.3 mm) and 4.4 mm (SD ¡ 2.6 mm) in the mediolateral, superoinferior and anteroposterior dimensions. Pre-training mean of the maximum marker movement during normal breathing was 1.07 mm, 1.94 mm and 1.86 mm in the mediolateral, superoinferior and anteroposterior dimensions. During deep breathing these values were 2 mm, 5.5 mm and 4.8 mm. While respiratory training had negligible effect on breast movement during normal breathing, it resulted in a modest reduction during deep breathing (p50.2). The mean CMME recorded for these patients was 3.4 mm, 4.5 mm and 7.1 mm in the mediolateral, superoinferior and anteroposterior dimension. Respiratory training also resulted in an increase in breath hold time from a mean of 31 s to 44 s (p50.04) and tidal volume from a mean of 560 cm3 to 1160 cm3 (p50.04). With patients immobilized in the vacuum bag the CMMEs are relatively less. Individualized directional margins may aid in reduction of planning target volume (PTV).

Accelerated partial breast irradiation (APBI) is presently the focus of research in radiotherapy for early breast cancer (EBC) [1]. In addition to interstitial brachytherapy, three-dimensional conformal radiation therapy (3DCRT) and intensity-modulated radiation therapy (IMRT) are currently being evaluated for APBI in this group of patients [1–5]. For successful implementation of these high precision techniques both intrafraction and interfraction movements of the breast have to be considered. Efforts have been made to decrease respiratory movement by the use of active breathing control (ABC), respiration correlated cone beam CT (RC-CBCT), real time tumour tracking and helical tomotherapy while implementing 3DCRT for APBI [6–10]. While efforts are being made to incorporate image guidance for precise localization of planning target volume (PTV) for APBI, there is a paucity of data regarding cumulative positional uncertainties due to intrafraction respiratory movement and interfraction-repositioning errors with different positioning devices in patients with breast cancer. This study was designed with the aim: Address correspondence to: Prof. Rajiv Sarin, Director, ACTREC, Tata Memorial Centre, Kharghar, NaviMumbai, Maharashtra, India; E-mail: [email protected].

766

Received 11 January 2006 Revised 25 March 2006 Accepted 27 March 2006 DOI: 10.1259/bjr/98024704 ’ 2006 The British Institute of Radiology

(a) To ascertain intrafraction respiratory movement of the breast. (b) To evaluate interfraction movement of breast clinical target volume (CTV) by studying the re-positioning set up errors with individualized vacuum cast immobilization. (c) To evaluate the possible efficacy of respiratory training in reducing breast movement for the purpose of external beam APBI with 3DCRT and IMRT. (d) To devise guidelines for CTV to PTV margins based on cumulative maximum movement error (CMME) combining intrafraction or interfraction maximum respiratory motion during normal breathing and maximum re-positioning set up errors.

Methods and materials Five patients with EBC who had undergone breast conservative surgery at our institution and were potentially eligible for APBI as per standard eligibility criteria for APBI [1] were included in this study. Patients with restricted arm movements after surgery and preexisting respiratory problems were excluded. Written informed consent was obtained from all the patients. While these women participated in this investigational The British Journal of Radiology, September 2006

Breast movement and implications for partial breast irradiation

study for respiratory movements and set up errors during positioning for APBI, they were treated with the standard radiotherapy protocol of our department. The standard treatment included 5 weeks of daily whole breast radiotherapy with bitangential portals with 6 MV photons on an inclined breast board followed by an en face electron boost. For the purpose of this study, the women underwent a separate process of evaluation of set up errors for which they were immobilized in an individualized body cast (vacuum bag). Baseline movements during normal and deep breathing, breath hold time and tidal volume were evaluated for all the women. Breath hold time and tidal volume were measured on a daily basis during the course of respiratory training. After a short course of 8– 10 days of respiratory training, breast movement during normal and deep breathing was re-evaluated.

Immobilization and surface markers Patients were immobilized in individualized body casts made of a polyurethane bag filled with tiny polystyrene pellets, which sets according to the body shape upon application of vacuum. Patients were positioned in the body cast in supine position with arms above the head without using the breast board. The lumpectomy cavity was localized and outlined on the breast surface using the information from pre-operative clinical description, mammography films, intraoperative findings and lumpectomy scar. With an additional margin of 1.5 cm in the x and y dimensions the CTV was delineated on the patient’s surface. Lead markers (2 mm diameter) were placed at the centre of each of the four borders of the outlined CTV and at the centre of the CTV (Figure 1). For the purpose of the study it was assumed that the movement of these five markers on the anterior surface of the CTV would represent the movement of the entire CTV.

Interfraction movement: re-positioning set up errors After carefully positioning a patient in the customized body cast on the simulator (Ximavision 2.1H; Varian Medical Systems, Palo Alto, CA), a field size of 10 cm 6 10 cm was centred on the CTV at a target to skin distance of 95 cm. Using an isocentric technique, orthogonal anterior and lateral images were acquired. Patients were instructed to hold their breath after deep inspiration during image acquisition. This served as a baseline image against which set up reproducibility was evaluated. After the baseline image acquisition patients were then taken off the couch and re-positioned four times at different time intervals on the same day and orthogonal images were repeated. For each patient, four pairs of orthogonal films were compared against the baseline orthogonal pair of films. On each of these films, the perpendicular distance between the centre of each spherical marker and the nearest delineator wire or the central crosswire of the simulator was measured using an inbuilt measuring scale with an accuracy of 0.1 mm. The differences between the measurements in The British Journal of Radiology, September 2006

Figure 1. Localization of clinical target volume (CTV) on the breast surface with lead markers.

the baseline and subsequent films were recorded for each of these markers in the mediolateral, superoinferior and anteroposterior dimensions. Internal surgical titanium clips demarcating the tumour bed were present in one patient. However, the visualization of these titanium clips was poor in orthogonal films and not suitable for evaluation of tumour bed movement. Systematic error (S), defined as the variation between the planned position and average position on repositioning, was calculated as the average value of the mean deviation of each patient. Random errors (s) defined as fraction to fraction variations around the mean deviation were calculated as the average of the standard deviation around the mean [11].

Intrafraction movement: respiratory movement of breast Baseline evaluation of respiratory movements was carried out in normal as well as deep breathing, using the cine-acquisition mode of XimavisionH. Orthogonal cine-fluoroscopic images were acquired at the rate of two images per second with the gantry at 0 ˚ and 90 ˚ (or 270 ˚). At the end of image acquisition, cine recordings were available for each patient and the orthogonal image sets could be used for evaluation of movement in mediolateral, superoinferior and anteroposterior dimensions. A total of 80 images (40 orthogonal images in normal breathing and 40 orthogonal images in deep breathing) were available per patient. For the purpose of measurement or tracking of movements of breast markers during respiration, the projection of the field delineator and the central cross wires were considered as a stationary structure against which the position of individual markers was measured in serial cine images 767

S Chopra, K A Dinshaw, R Kamble and R Sarin Table 1. Set up variation and cumulative set up errors in mediolateral, superoinferior and anteroposterior dimensions with the use of individualized body cast Sr. No:

Patient Patient Patient Patient Patient Mean

Mediolateral (mm) Mean¡SD (Range)

1 2 3 4 5

0.5¡0.6 2.0¡1.7 1.3¡1.4 1.5¡0.9 1.3¡0.8 1.3¡0.5

(–0.3 (–2.7 (–0.6 (0 to (–0.5

Superoinferior (mm) Mean¡SD (Range)

to +0.9) to +0.2) to –3.9) –2.1) to +2.3)

1.5¡1.3 1.4¡0.5 1.0¡0.8 1.9¡1.7 1.0¡0.7 1.3¡0.3

acquired during the respiratory cycle. On each of these images, the perpendicular distance between the marker and the nearest delineator wire or central crosswire of the simulator was measured as for set-up errors. Differences between the measurements in the baseline and subsequent images were recorded for each of these markers for ascertaining the breast movement in the mediolateral, superoinferior and anteroposterior direction during a few respiratory cycles.

Respiratory training After baseline assessment of respiratory movement, all patients were given a short course of respiratory training for a period of 8–10 days. Patients were trained by an occupational therapist (RK). Deep breathing exercises, including inspiratory and expiratory manoeuvres and forced abdominal expiration technique, were taught to all the patients. A spirometer was used on a daily basis for training. Patients carried out all these exercises for 15–20 min once daily under the supervision of the occupational therapist and repeated them at least twice daily without the instructor. Breath holding time and tidal volume were recorded before starting respiratory training. While patients received respiratory training daily breath hold time and tidal volume were noted. Detailed cine-fluoroscopic evaluation of breast movement as described above was done before starting the respiratory training and repeated once after completion of respiratory training. Any reduction in the movement of the skin markers during respiration, increase in the

(+0.3 to +3.3) (–1.9 to +1.7) (–0.3 to +1.9) (–0.4 to +4.2) (–0.2 to –1.9)

Anteroposterior (mm) Mean¡SD (Range)

4.6¡1.2 1.8¡1.7 6.0¡0.8 7.8¡1.9 1.8¡0.5 4.4¡2.6

(–2.7 to –5.2) (–0.5 to +5.0) (–5 to –7.2) (0 to –9.2) (–0.8 to –2.7)

tidal volume or breath holding time was used as an end point to evaluate the efficacy of short course respiratory training. The Wilcoxon Sign rank test was used to evaluate the statistical significance of observed differences.

Observations and results Re-positioning errors using the customized vacuum bag were evaluated in the mediolateral, superoinferior and anteroposterior dimensions for all patients. Cranial, anterior and right sided deviations were recorded in the positive direction whereas caudal, posterior and left sided deviations were recorded in the negative direction. The details of the set up errors on re-positioning are shown in Table 1. The mean deviations for all markers from their baseline positions were 1.3 mm (SD ¡ 0.5 mm) in the mediolateral; 1.3 mm (SD ¡ 0.3 mm) in the superoinferior and 4.4 mm (SD ¡ 2.6 mm) in the anteroposterior dimension. Rotational errors were not calculated. Breast movement, as judged by five surface markers, was minimal during normal breathing with pre-training mean marker movement of 1.07 mm (0.6–1.4 mm); 1.94 mm (0.7–4.5 mm) and 1.86 mm (0.6–1.9 mm) in the mediolateral, superoinferior and anteroposterior directions, respectively, during normal breathing. The breast movements during normal breathing were not affected by respiratory training (Table 2). As expected, the marker movement was more pronounced during deep breathing with mean values of

Table 2. Maximum movement during normal breathing in anteroposterior, mediolateral and superoinferior dimension before and after respiratory training Mediolateral (mm)

Patient 1 Patient 2 Patient 3 Patient 4 Patient 5 Mean of max. movement in five patients Wilcoxon Signed Rank sum test (p-value)

Superoinferior (mm)

Anteroposterior (mm)

Pre-training

Post-training

Pre-training

Post-training

Pre-training

Post-training

1.4 NA 1.6 0.6 0.7 1.07 (0.6–1.4)

1.4 1.6 0.4 0.6 1.1 1.02 (0.4–1.4)

4.5 0.8 2.3 1.4 0.7 1.94 (0.7–4.5)

4.5 1.6 1.2 1.4 0.6 1.86 (0.6–4.5)

1.9 0.8 4.9 1.1 0.6 1.86 (0.6–4.9)

1.7 0.9 1.6 0.7 1.2 1.09 (0.7–1.7)

0.6

0.5

0.3

NA, not available.

768


Breast movement and implications for partial breast irradiation Table 3. Maximum movement of breast in anteroposterior, mediolateral and superoinferior dimension during deep breathing before and after respiratory training Mediolateral (mm)

Patient 1 Patient 2 Patient 3 Patient 4 Patient 5 Mean of max. Movement in five patients Wilcoxon Signed Rank sum test (p-value)

Superoinferior (mm)

Anteroposterior (mm)

Pre-training

Post-training

Pre-training

Post-training

Pre-training

Post-training

1.7 NA 3.4 2.0 1.0 2.0 (1.0–3.4)

2.0 2.2 2.8 0.9 1.0 1.7 (0.9–2.8)

8.2 5.7 5.0 7.4 1.4 5.5 (1.4–8.2)

6.6 4.1 5.4 4.2 1.2 4.3 (01.2–6.6)

9 3.8 5.5 1.7 4.1 4.8 (1.7–9)

6 1.3 3.8 2.5 5.0 3.7 (1.3–6)

0.2

0.1

0.2

NA, not available.

Discussion

2 mm (1–3.4 mm), 5.5 mm (1.4–8.2 mm) and 4.8 mm (1.7–9 mm) in the mediolateral, superoinferior and anteroposterior directions, respectively (Table 3). Respiratory training produced a non-significant reduction of movement of marker in mediolateral, superoinferior and anteroposterior directions during deep breathing. After training the breath hold time increased from a pre-training mean of 31 s to a mean of 44 s (p50.04). The tidal volume increased in all the patients from an average of 560 cm3 in the pre-training phase to 1160 cm3 in the post-training phase (p50.04) (Table 4, Figures 2 and 3).

While several studies have previously evaluated set up errors in women undergoing daily breast radiotherapy [12–17], this is probably the first study in which both the interfraction (re-positioning set up errors) and intrafraction (respiratory) components of variability in CTV localization have been ascertained individually for each study patient. Our findings confirm that re-positioning set up errors and respiratory movements both contribute to the total uncertainty. The maximum interfraction and intrafraction movement readings have been combined to derive the cumulative maximum movement error for each patient. In our opinion, this represents a simple and perhaps more accurate and safe method of deriving the CTV to PTV margin for individual patients. In this study instead of relying on the surgical clips, which are usually placed on the posterior surface of the lumpectomy cavity, we have evaluated respiratory movements by skin surface markers representing the anterior surface of the CTV. We feel that it is reasonable to assume that the movement of the breast surface represents movement of the CTV, except in patients with large pendulous breasts where surface markers may not be so reliable for representing the underlying CTV. As we observed in one patient, titanium surgical clips are not visualized well in all projections that are required for assessing set up errors and respiratory movements, especially in the true lateral projection. Moreover, it has been reported that the surgical clips may be

CTV to PTV margins based on cumulative maximum movement error For determining the CTV to PTV margin for each study patient, we derived the CMME by combining the maximum intrafraction movement as judged by cinefluoroscopic assessment of breast respiratory motion during normal breathing and maximum interfraction repositioning set up errors obtained in these test conditions (Table 5). The CMME did not exceed 10 mm for any of these patients in any of the axes. The mean of the CMME recorded for these 5 patients was 3.4 mm (range 2.3– 4.3 mm), 4.5 mm (range 2.5–7.8 mm) and 7.1 mm (range 3.9–9.9 mm) in the mediolateral, superoinferior and anteroposterior dimensions, respectively.

Table 4. Pre- and post-respiratory training breath hold time and tidal volume Patient no.

Patient 1 Patient 2 Patient 3 Patient 4 Patient 5 Mean Median Wilcoxon Signed Rank Test (p-value)

Breath hold time in seconds

Tidal volume in ml

Pre-training

Post-training

Pre-training

Post-training

25 25 45 23 39 31.4 25 0.04

40 40 46 46 50 44.5 46

400 800 800 400 400 560 400 0.04

800 1600 1200 1200 1000 1160 1200


769

S Chopra, K A Dinshaw, R Kamble and R Sarin

Figure 2. Box whisker plot showing variation of breath hold time during respiratory training over 1 week (BH 1–75 breath hold time day 1–7). y-axis: breath hold time in seconds; x-axis: breath hold time over 1 week of respiratory training; N5 number of patients.

displaced by up to 3 mm (range 0–11 mm) [18] especially if they are not anchored to the muscle. Other authors have described using bony or soft tissue reference points for measuring set up errors or respiratory movements [15–17]. In our opinion, none of the soft tissue or bony reference points in this region are so well defined that they can be reliably and accurately localized in all projections in serial imaging.

Interfraction movement: re-positioning set up errors Previous studies have measured re-positioning set up errors using bony or soft tissue reference points [12–17] during a 5–6 week course of daily radiotherapy. In contrast, we have ascertained the set up errors by repositioning the patients five times during the same day and by using skin markers. While our approach of repositioning assessments over 1 day may not reflect time trends in set up variation over 5 weeks [19–21] it has certain other advantages. Most importantly, it removes the confounding variability in the placement of markers or problems in precisely localizing the same reference point in the thoracic bony cage or soft tissues in serial orthogonal films. For abbreviated treatment like APBI which is delivered over 5 days, time trends if any, are likely to be much less than that observed over a treatment course of 35 days [21]. Assessment over 1 day allows the use of individualized decision regarding CTV to PTV margins, the benefits of which are discussed later.

The set up errors with this individualized body cast immobilization were quite small under these test conditions, except in the anteroposterior direction (Table 1). However, we have not compared the re-positioning set up errors without body cast. Similar findings have been reported by other groups using various immobilization devices [12–15]. Nalder [12] et al evaluated set up reproducibility for breast radiotherapy using a breast board with and without a Vacfix device (PAR Scientific A/S, Denmark). They noted reduction in set up errors in the superoinferior direction using the Vacfix device. Giraud et al [22] have demonstrated that a personalized body cast leads to a significant reduction in lung apex movements along with reduction in lateral displacement of the thoracic cage. Vicini et al [4] have used vacuum bags for immobilization for APBI and 5 mm margin has been incorporated for re-positioning errors.

Intrafraction movement: respiratory movement of breast Ideally the entire treatment period should be continuously monitored and orthogonal images should be acquired simultaneously for assessing intrafraction motions. However, for reasons of logistics and in order to minimize radiation exposure to the opposite breast with lateral beams, we used cine-fluoroscopic assessment of breast motion for only 10 s. Since the beam on time for each radiation field of APBI using 3DCRT is usually in the range of 20–30 s, this 10 s assessment may

Figure 3. Box whisker plot showing the variation in tidal volume during respiratory training over 1 week. (Tidal 1–7 5 tidal volume from day 1–7). y-axis: tidal volume in millilitres; x-axis: tidal volume over 1 week of respiratory training; N5 number of patients. 770


Breast movement and implications for partial breast irradiation Table 5. Measured cumulative maximum movement error (CMME in mm) combining maximum respiratory motion in normal breathing after respiratory training and maximum set up errors Patient

Mediolateral CMME (mm)

Patient 1 Patient 2 Patient 3 Patient 4 Patient 5 Mean CMME

2.3 4.3 4.3 2.7 3.4 3.4

(1.4 (1.6 (0.4 (0.6 (1.1

+ + + + +

0.9) 2.7) 3.9) 2.1) 2.3)

Superoinferior CMME (mm)

7.8 3.5 3.1 5.6 2.5 4.5

(4.5 (1.6 (1.2 (1.4 (1.6

+ + + + +

3.3) 1.9) 1.9) 4.2) 1.9)

Anteroposterior CMME (mm)

7.2 (1.7 + 5.2) 5.9 (0.9 + 5.0) **8.8 (1.6 + 7.2) **9.9 (0.7 + 9.2) 3.9 (1.2 + 2.7) 7.1

**Use of pre-respiratory training measurements of respiratory movement would have made the combined maximum movement over 10 mm (12.1 mm and 10.3 mm).

therefore be a reasonable representation of intrafraction movement. The findings of our study confirm that respiratory movements during normal breathing are negligible (,2 mm) and may be ignored for conventional whole breast treatment, whereas deep breathing could cause significant intrafraction error. However, respiratory motion even during normal breathing merits special consideration when planning external beam APBI using 3DCRT or IMRT and even for tangential whole breast treatment if using IMRT. With IMRT, not only the CTV is moving with respiration but the lack of synchrony between respiratory motion and the dynamic sliding multileaf collimator could result in perturbation in radiation dose delivery [23]. Hence for IMRT, respiratory movements are a matter of concern even for whole breast irradiation. Other authors have evaluated movements of the breast in various dimensions with varying respiratory patterns ranging from no breathing to deep breathing and reported 2–16 mm, 2–6 mm and 2–16 mm movement in the mediolateral, superoinferior and anteroposterior dimensions, respectively, and documented that the best reproducibility was achieved when patients were asked to hold their breath [23, 24]. In our study we recorded baseline average movement of less than 2 mm in mediolateral, superoinferior and anteroposterior dimensions during normal breathing. At the William Beaumont Hospital, ABC was used in the initial few patients undergoing external beam APBI [3, 6]. However, ABC is not used currently for APBI and a generic margin of 10 mm is being incorporated (5 mm for set up errors and 5 mm for breathing movements) for CTV to PTV generation. Formenti et al [2, 5] have treated patients with APBI in prone position with the aim of reducing breathing movements. However, up to 1 cm allowance has to be incorporated for enhanced set up errors in this position. Improved accuracy, both in re-positioning and minimizing target motion, has recently been reported with the incorporation of video based surface registration, optoelectronic surface registration and RC-CBCT [8–10]. This may further aid in reducing CTV to PTV expansion, but limited access to these devices may limit its wider applicability. Although there was only minimal impact of the short course of respiratory training in reducing breast movements, the tidal volume and breath hold time increased appreciably in all the patients. The post-training mean breath hold time was 44 s as compared with 12–16 s in untrained patients as reported by other studies [25]. This paradox of increased tidal volume with similar or The British Journal of Radiology, September 2006

reduced thoracic movement may be explained by an increase in the diaphragmatic component of breathing. However, this was not specifically measured in the present study. Increasing the diaphragmatic component of respiration and increased breath hold time may have two implications in thoracic radiotherapy. First, respiratory training imparted in this fashion has the possibility of exacerbating lung movements, especially of the lower lobe. Hence, the impact of respiratory exercises, which are sometimes imparted to patients with lung cancer in order to improve their pulmonary functions, should be studied with special reference to their impact on respiratory lung movement during shallow or normal breathing. Movements during deep breathing are irrelevant since it is unlikely that any clinical context would require delivering radiotherapy during continuous deep breathing. Second, such short course respiratory training could help deep inspiration breath hold (DIBH) assisted treatment of thoracic or abdominal tumours by increasing the breath hold time, thereby shortening the treatment delivery time and possibly improving accuracy in treatment delivery.

CTV to PTV margins In the literature, there are several recommendations for deriving the CTV to PTV margins [11, 26]. Using a mathematical notation and assuming that the distribution of set up errors is gaussian, Stroom et al have recommended that a safety margin of 2.0S + 0.7s should be incorporated where S and s are as previously defined [11]. The criteria for this recommendation are based on an assumption that 99% of the CTV would be covered by 95% of the prescribed dose. While this recommendation has been favoured by other investigators [27] we feel that there are two caveats with this recommendation. First, set up errors may not always have a normal or gaussian distribution as seen in our data as well as in other studies [28, 29]. Second, such values of systematic and random errors and the CTV to PTV margins based on these formulae are derived for a particular study population and then recommended to be used for generating the CTV to PTV margin for individual patients. The present study as well as other investigators [30] have shown that there could be significant patient to patient variation both in the magnitude and vector of the mean or maximum set up error. For highly conformal radiation plans that are delivered in very abbreviated schedules, such as APBI over 5 days or highly hypofractionated schedules such as extracranial body frame stereotaxy 771

S Chopra, K A Dinshaw, R Kamble and R Sarin

[31, 32] use of generic values of CTV to PTV margins may result in under coverage or an inappropriately large margin for a substantial part of the treatment. For such highly abbreviated or hypofractionated schedules, the maximum re-positioning set up error measured over repeated re-positioning on a single day along with the measured respiratory movements can be incorporated in the initial individualized CTV to PTV generation algorithm. This would ensure the use of minimum yet safe CTV to PTV margins and in patients where the margins thus obtained are unacceptably large, the process of immobilization may be repeated or changed or the patient may be given special instruction. Finally, the appropriateness of these margins could be further validated with online portal imaging and if necessary online or offline correction [7].

Conclusions Interfraction re-positioning set up errors and intrafraction respiratory movement of the breast contribute to total uncertainty in irradiating the intended CTV. Moreover, there are significant interindividual differences in the maximum values of interfraction and intrafraction movement errors. With the use of individualized vacuum bag fixation and individually ascertained values for set up errors and respiratory movements on a single day it is possible to use tight CTV to PTV margins for external beam APBI delivered during normal breathing. It is unlikely that more advanced techniques such as ABC, real time tumour tracking and image guided radiotherapy would result in ‘‘clinically significant’’ reduction in the CTV to PTV margin for APBI over and above what is achieved by the technique described here. Although not of much use for APBI, by increasing the breath hold time, a short course of respiratory training may be useful for patients undergoing DIBH for intrathoracic or abdominal tumours.

6.

7.

8.

9.

10.

11.

12.

13.

14.

15.

16. 17.

References 1. Sarin R. Partial breast treatment for early breast cancer: emergence of a new paradigm. Nat Clin Pract Oncol 2005;2:40–7. 2. Formenti SC, Rosenstein B, Skinner KA, Jozsef G. TI stage breast cancer: adjuvant hypofractionated conformal radiation therapy to tumor bed in selected postmenopausal breast cancer patients-pilot feasibility study. Radiology 2002;222:171–8. 3. Baglan KL, Sharpe MB, Jaffray D, Frazier RC, Fayad J, Kestin LL, et al. Accelerated partial breast irradiation using 3D conformal radiation therapy (3DCRT). Int J Radiat Oncol Biol Phys 2003;55:302–11. 4. Vicini FA, Remouchamps V, Wallace M, Sharpe M, Fayad J, Tyburski L, et al. Ongoing clinical experience utilizing 3D conformal external beam radiotherapy to deliver partial breast irradiation in patients with early stage breast cancer treated with breast conserving therapy. Int J Radiat Oncol Biol Phys 2003;57:1247–53. 5. Formenti SC, Truong MT, Goldberg JD, Mukhi V, Rosenstein B, Roses D, et al. Prone accelerated partial breast irradiation (p-apbi) after breast conserving surgery: preliminary clinical results and dose volume histogram

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18.

19.

20.

21.

22.

(DVH) analysis. Int J Radiat Oncol Biol Phys 2004;60:493–504. Wong JW, Sharpe MB, Jaffray DA, Kini VR, Robertson JM, Stromberg JS, et al. The use of active breathing control (ABC) to reduce margin for breathing motion. Int J Radiat Oncol Biol Phys 1999;44:911–9. Kim L, Vicini F, Yan D, Vargas C, Martinez A, Wong J. Reduction of PTV margin for accelerated partial breast irradiation using on-line detection of surgical clips. Int J Radiat Oncol Biol Phys 2004;60:S336. Shirato H, Shimizu S, Kitamura K, Nishioka T, Kagei K, Hashimoto S, et al. Four dimensional treatment planning and fluoroscopic real time tumor tracking: radiotherapy for moving tumor. Int J Radiat Oncol Biol Phys 2000;48: 435–42. Riboldi M, Baroni G, Orecchia R, Pedotti A. Enhanced surface registration techniques for patient positioning control in breast cancer radiotherapy. Technol Cancer Res Treat 2004;3:51–8. Hui SK, Das RK, Kapatoes J, Oliviera G, Becker S, Odau H, et al. Helical tomotherapy as a means of delivering accelerated partial breast irradiation. Technol Cancer Res Treat 2004;3:639–46. Stroom JC, de Boer HC, Huizenga H, Visser AG. Inclusion of geometrical uncertainties in radiotherapy treatment planning by means of coverage possibility. Int J Radiat Oncol Biol Phys 1999;43:905–19. Nalder CA, Bidmead AM, Mubata C D, Tait D, Beardmore C. Influence of vac-fix immobilization device on the accuracy of patient positioning during routine breast radiotherapy. Br J Radiol 2001;74:249–54. Pradier O, Schmidberger H, Weiss E, Bouscayrol H, Daban A, Hess CF. Accuracy of alignment in breast irradiation: a retrospective analysis of clinical practice. Br J Radiol 1999;72:685–90. Bohmer D, Feyer P, Harder C, Korner M, Sternemann M, Dinges S, et al. Verification of set up deviations in patients with breast cancer using portal imaging in clinical practice. Strahlenther Onkol 1998;174 S2:36–9. Lirette A, Pouliot J, Aubin M, Larochelle M. Efficacy of portal imaging in tangential breast irradiation: a prospective study. Radiother Oncol 1995;37:241–5. Herman MG. Clinical use of electronic portal imaging. Semin Radiat Oncol 2005;15:157–67. Hurkmans CW, Remeijer P, Lebesque JV, Mijnheer BJ. Set up verification using portal imaging: review of current clinical practice. Radiother Oncol 2001;58:105–20. Weed D, Yan D, Martinez A, Vicini FA, Wilkinson TJ, Wong J. The validity of surgical clips as a radiographic surrogate for the lumpectomy cavity in image guided accelerated partial breast irradiation. Int J Radiat Oncol Biol Phys 2004;60:484–92. Ghizelan MJ, Jaffray DA, Siewerdsen JH, Van Herk M, Shetty A, Sharpe MB, et al. Prostate gland motion assessed with cine magnetic resonance imaging. Int J Radiat Oncol Biol Phys 2005;62:406–17. Langen KM, Pouliot J, Anezinos RTT, Aubin M, Gottschalk AR, Hsu IC, et al. Evaluation of ultrasound based prostate localisation for image-guided radiotherapy. Int J Radiat Oncol Biol Phys 2003;57:635–44. Trog D, Garbe S, Lutterbey G, Barwig P, Boldt I, Stolz A, et al. Volumetric changes of the breast during radiotherapy. Is replanning necessary for the electron boost? Strahlenther Onkol 2005;181:255–9. Giraud P, De Rycke Y, Dubray B, Helfre S, Voican D, Guo L, et al. Conformal radiotherapy (CRT) planning for lung cancer: analysis of intrathoracic organ motion during extreme phases of breathing. Int J Radiat Onc Biol Phys 2001;51:1081–92.


Breast movement and implications for partial breast irradiation 23. George R, Keall PJ, Kini VR, Vedam SS, Siebers JV, Wu Q, et al. Quantifying the effect of intrafraction motion during breast IMRT planning and dose delivery. Med Phys 2003;30:552–62. 24. Vedam SS, Kini VR, Keall PJ, Ramakrishnan V, Mostafavi H, Mohan R, et al. Quantifying the predictability of diaphragm motion during respiration with a non-invasive external marker. Med Phys 2003;30:505–13. 25. Hanley J, Debois MM, Mah D, Mageras GS, Raben A, Rozenweig K, et al. Deep inspiration breath-hold technique for lung tumors: the potential value of target immobilization and reduced lung density in dose escalation. Int J Radiat Oncol Biol Phys 1999;45:603–11. 26. Herk MV, Remeuer P, Rasch C, Lebesque JV. The probability of correct target dosage: dose-population histograms for deriving treatment margins in radiotherapy. Int J Radiat Oncol Biol Phys 2000;47: 1121–35. 27. Alheit H, Domfeld S, Dawel M, Alheit M, Henzel B, Steckler K, et al. Patient position reproducibility in fractionated stereotactically guided conformal radiotherapy using the


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BrainLab mask system. Strahlenther Onkol 2001;177: 264–8. Lujan AE, Ten haken RK, Larsen EW, Balter JM. Quantization of set up uncertainties in 3D dose calculations. Med Phys 1999;26:2397–402. Huizenga H, Levendag PC, De Porre, Visser AG. Accuracy in radiation field alignment in head and neck cancer: a prospective study. Radiother Oncol 1988;11:181–7. Erridge SC, Seppenwolde Y, Muller SH, Van Herk M, De Jaeger K, Belderbos JS, et al. Portal imaging to assess set up errors, tumor motion and tumor shrinkage during conformal radiotherapy of non small cell lung cancer. Radiother Oncol 2003;66:75–85. Zimmerman FB, Geinitz H, Schill S, Grosu A, Schratzenstaller U, Molls M, et al. Stereotactic hypo fractionated radiation therapy for stage I non-small cell lung cancer. Lung Cancer 2005;48:107–14. Balter JM, Brock KK, Lam KL, Tatro D, Dawson LA, Mcshan DL, et al. Evaluating the influence of set up uncertainties on treatment planning for focal liver tumors. Int J Radiat Oncol Biol Phys 2005;63:610–4.

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Book review Computed tomography, 2nd edn. By WA Kalender. pp. 306, 2005 (Publicis, Erlangen, Germany) J49.90 ISBN 3-89578-216-5 This is an excellent well-written and well-illustrated book, which deals with the scientific and technical aspects of computed tomography. CT scanning has developed rapidly in the last few years and now makes up half of the collective dose delivered to patients from diagnostic radiology (47% in Germany in 2003). There is a need for medical physicists, radiologists and radiographers to have a thorough understanding of the technique and the information it can provide in order that its full potential is realised. This book by an author who has played an important role in the development of spiral CT addresses this need. The book explains the fundamentals and application of CT with the main focus on physics and technology, and deals with problems relevant to the user of CT equipment. It is directed at a multidisciplinary readership and the concepts are explained clearly for all staff involved in use of the technique. This is the second edition of the textbook, which has been revised to include important new developments in detector technology and applications. The first chapter covers fundamentals, starting from first principles. Those working with the topic can skim quickly through the chapter before delving into the more meaty content of the later chapters. Others will find it a useful introduction. Chapter 2 contains technical concepts, including scanner design, collimation, filtration, detector systems, can modes and finishing off with a section on important aspects of modern equipment, such as the use of cone-beam geometry in multislice scanners and combined CT/PET scanners. Chapter 3 deals with spiral CT and chapter 4 image quality, including considerations of requirements for 64 slice scanners.

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Chapter 5 is devoted to radiation protection with explanations of the standard dose quantities. Questions regarding the possibility of dose reduction are discussed in a separate section. The chapter promotes the thesis that CT when applied properly should not be regarded as a high dose technique. Two and three dimensional approached to display of information are explained in chapter 6 with consideration of interactive approached for assessment of the large volume of data generated. Techniques are illustrated with a CD-ROM issued with the book. Example images can be loaded and viewed interactively, so that readers can undertake exercises in image viewing and adjustment themselves. This provides an invaluable teaching aid for those who do not have easy access to CT data themselves. Moreover, the author indicates in the preface to the new edition that readers may use the figures in their own presentations and teaching material. Chapter 7 deals with clinical applications, such as CT angiography, interventional CT and cardiac imaging, while in chapter 8 the author provides his view on future developments and includes some wishes about how perception and use might change in the future. The full mathematical treatment of image reconstruction is left until chapter 9 for those with the mathematical background and the need to understand the image reconstruction methods in more depth, so that others are not put off by encountering the more complex formulae in the earlier chapters. This is an excellent text covering every aspect of the physics and technology of CT, which is suitable for a wide range of staff involved in use and applications of CT scanners. It would make a useful addition to any Radiology or Medical Physics Department library.

C J MARTIN



CASE OF THE MONTH

A case of spinal cord compression of unknown cause D BUTTERISS,

FRCR

and C SOH,

FRCR

Department of Neuroradiology, Regional Neurosciences Centre, Newcastle General Hospital, Westgate Road, Newcastle-upon-Tyne NE4 6BE, UK

Received 21 June 2005 Revised 27 July 2005 Accepted 16 August 2005 DOI: 10.1259/bjr/29157329 ’ 2006 The British Institute of Radiology

Case report

sensation below the shoulders and perianal/saddle paraesthesia. A presumptive diagnosis of spinal cord compression in the mid-cervical spine was made and the patient was referred for spinal MRI. Selected axial and sagittal images of the cervical (Figure 1) and lumbar spine (Figure 2) are shown. What do the images show? Is there evidence of myelopathy? What is your differential diagnosis? CT of the cervical spine was subsequently performed, prior to surgery (Figure 3). Does this alter or narrow your differential diagnosis?

A 65-year-old woman presented to the accident and emergency department complaining of a 3-month history of falls, reduced power and numbness in her arms and legs, and urinary incontinence. Her status had deteriorated over the previous 4 weeks and she was bed-bound on admission. There was no significant past medical history. Examination revealed reduced power in all limbs, increased tone in the right upper limb, altered Address correspondence to: Dr C Soh.

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(c)

Figure 1. Sagittal (a) T1 weighted, (b) T2 weighted MR images through the cervical spine. (c) Representative axial T1 weighted image through the cervical spine. The British Journal of Radiology, September 2006

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D Butteriss and S Soh

(a)

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(c)

Figure 2. Sagittal (a) T1 and (b) T2 weighted MR images through the lumbar spine. (c) Representative axial T2 weighted image through the lumbar spine.

MRI demonstrates low/intermediate signal masses on T1 and T2 imaging, involving the C3–C6 vertebral bodies, with destruction of C4 and C5 extension into the paravertebral soft tissues. At C4–C5 there is extradural compression of the cervical cord, with high intramedullary cord signal on T2 weighted imaging consistent with myelopathy. CT of this level shows a relatively high attenuation mass containing fragmentary calcification that is destroying the cervical vertebral bodies and extending into the surrounding soft tissues and into the spinal canal with thecal compression. MRI of the lumbar spine shows involvement of the L3– S1 vertebral bodies, posterior elements and facet joints with small extradural masses, but no evidence of neural compression. Further imaging demonstrated further lesions in the sacroiliac joints, both hips and the pubic symphysis. Neurosurgical C3–C6 partial anterior vertebrectomies with iliac crest grafting and plate fixation revealed pale, cheesy deposits, shown to be gouty tophi on histological investigation. Urate-lowering medical therapy was instituted.

Discussion Gout is a relatively common metabolic condition characterized by deposition of monosodium urate crystals in joints and soft tissues. In Europe the incidence is approximately 0.2–0.5% and the male to female ratio is approximately 5:1. The acute presentation is with a monoarthritis, in 25% of cases affecting the first tarsometatarsal joint (podagra). Chronicity results in the formation of tophi within the soft tissues around joints and in the pinnae of the ears due to deposits of monosodium urate and associated inflammatory cells. This is rare before at 776

least 10 years of acute attacks. The peripheral joints are most commonly affected, but occasionally spinal involvement with gouty tophi occurs [1]. The presentation of spinal tophaceous gout is nonspecific, with local pain, radicular symptoms and paraparesis or tetraparesis occurring due to neural compression. There is usually, but not invariably, a history of prior acute attacks of gout, or of hyperuricaemia [2]. MRI appearances are variable. Tophi appear as homogeneous intermediate/low intensity masses on T1 weighted imaging, but have a variable T2 weighted appearance ranging from homogeneous low through heterogeneous signal to homogeneous high signal. Tophi may contain small areas of signal drop-out that have been shown to represent calcification on CT imaging. Enhancement post-gadolinium administration is also variable. Rim enhancement is the most common finding, but both complete lack of enhancement and homogeneous enhancement have both been described [3]. Appearances may mimic spinal infection, malignancy or haemorrhage [4]. Symptomatic spinal gout usually involves the epidural space, with variable involvement of the intradural component, ligamentum flavum, facet joints, posterior elements, vertebral bodies and paraspinal soft tissues [3]. Acute management in cases with neural compression usually requires surgical decompression, where the tophus appears as a chalky, cheesy or fibrous mass. Often the diagnosis is made on histological examination of the surgical sample. Local pain may respond to antiinflammatory medication or colchicine. Long-term allopurinol therapy may result in reduction in size or even disappearance of tophi. In conclusion, spinal tophaceous gout is a rare cause of spinal neural compression, but should be considered when The British Journal of Radiology, September 2006

Case of the month: Unknown cause of spinal cord compression

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Figure 3. Axial CT images of the cervical spine in (a) bone and (b) soft tissue windows.

MRI demonstrates intermediate/low signal on T1 weighted deposits with low signal foci on all sequences, especially if there is involvement of the posterior elements.

References 1. Huskisson EC, Drury PL. Rheumatology and bone disease. In: Kumar P, Clark M, editors. Clinical medicine. London, UK: Bailliere Tindall, 1994:409–11.


2. Paquette S, Lach B, Guiot B. Lumbar radiculopathy secondary to gouty tophi in the filum terminale in a patient without systemic gout: case report. Neurosurgery 2000;46:986–8. 3. Hsu C-Y, Shih T T-F, Huang K-M, Chen P-Q, Sheu J-J, Li Y-W. Tophaceous gout of the spine: MR imaging features. Clin Radiol 2002;57:919–25. 4. Barrett K, Miller ML, Wilson JT. Tophaceous gout of the spine mimicking epidural infection: case report and review of the literature. Neurosurgery 2001;48:1170–2.

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October 2006 Volume 79 Issue 946

October 2006, Volume 79, Issue 946

●

Virtual pulmonary arterioscopy in pulmonary embolic disease

Breast radiotherapy in women with pectus excavatum (funnel chest): is the lateral decubitus technique an answer? A dosimetric study ●

●

Hypoxia in head and neck cancer

Is routine chest radiography a useful test in the follow up of all adult patients with soft tissue sarcoma? ●

Multiple hyperechoic testicular lesions are a common finding on ultrasound in Cowden disease and represent lipomatosis of the testis ●

Duplex ultrasound of the superior mesenteric artery in chronic pancreatitis ●

Accuracy and precision of an external-marker tracking-system for radiotherapy treatments ●

Normalized data for the estimation of fetal radiation dose from radiotherapy of the breast ●

Dosimetric and treatment planning considerations for radiotherapy of the chest wall ●

Patient and staff radiation doses from early radiological examinations (189921902) ●

●

Ultrasound spectrum in intraductal papillary neoplasms of breast


SHORT COMMUNICATION

Virtual pulmonary arterioscopy in pulmonary embolic disease C HOSKINS,

BSc, FRCR

and M CARPENTER,

BSc

Department of Diagnostic Imaging, Mayday University Hospital, Croydon CR7 7YE, UK ABSTRACT. 16 slice multidetector CT provides virtual endoscopic views of the inside of arteries, or any other hollow structures. This is performed non-invasively using postprocessing of three-dimensional isotropic image data sets, acquired during standard CT examinations. These virtual endoscopic views are simultaneously correlated with the standard multiplanar reconstructions, with the ability to navigate a virtual camera through the hollow structure under study. Normal and abnormal volume rendered images of the pulmonary arteries are presented in correlation with the multiplanar reformats. The abnormal images show the volume rendered appearances of acute and chronic pulmonary embolic disease. It is also postulated that this technique has a problem solving role in the differential diagnosis of chronic mural emboli from extravascular structures such as adjacent lymph nodes or bronchiolar impaction. This technique may also have a role in medical education, providing clinicians and medical students with interactive three-dimensional representations of disease processes.

The establishment of CT pulmonary angiography as the primary imaging technique for assessment of pulmonary embolic disease is widely recommended in many centres, as it combines the ability to directly visualize thrombus and evaluate the mediastinum and the lung parenchyma within one examination [1, 2]. Production of near isotropic data sets with 16+ slice multidetector CT has enabled the introduction and or refinement of numerous image processing techniques, avoiding the inherent distortion and artefacts associated with non-isotropic data [3]. Volume rendering (VRT) is one such technique that produces a 3D coloured image with depth perception. Unlike the earlier method of shaded surface display (SSD), VRT evaluates all the voxel intensities from the standard isotropic data set normally acquired for a routine examination. The VRT software assigns differing colours and degrees of transparency to CT number ranges within the volume and is thus able to display overlapping structures. It is also possible to position a virtual ‘‘endoscopic’’ camera within the arterial lumen guided by the multiplanar reconstruction (MPR) displays. This technique of ‘‘fly through’’ or perspective volume rendering (pVRT) allows visualization of the inner wall and any intraluminal contents of the arteries. Development of VRT relied heavily on the advancement of computing power [4], but is now universally applied to virtual colonoscopy, virtual bronchoscopy and virtual angioscopy/interior vessel analysis [5]. Identified applications for virtual angioscopy include ‘‘fly through’’ coronary angiography [6] and qualitative assessment of carotid artery stenosis [7]. The prerequisites for good angioscopic VRT imaging are the reconstruction of thin overlapping slices and the existence of a strong interface between the Hounsfield Units The British Journal of Radiology, October 2006

Received 6 January 2006 Revised 12 April 2006 Accepted 19 May 2006 DOI: 10.1259/bjr/40749658 ’ 2006 The British Institute of Radiology

of the structure of interest and its surroundings. For virtual angioscopy this interface is achieved with good contrast opacification of the blood vessels. In this paper we show the pVRT appearance of the pulmonary arteries, both in normal cases and in cases of acute and chronic thromboembolic disease. These are correlated with the MPR reformats. We postulate that this post-processing non-invasive technique has the potential to be used as a problem solving tool in the differential diagnosis of chronic thromboembolic disease versus periarterial lymph nodes and bronchiolar impaction. We also briefly discuss a possible role in medical education and the visualization of disease processes.

Methods Records were compiled of cases where a thoracic CT or pulmonary CT demonstrated acute or chronic pulmonary emboli, periarterial lymph nodes or a normal pulmonary arterial tree. These cases were protected on a workstation to enable future study.

CT scanning CT pulmonary angiograms were performed using a Siemens Somatom Sensation 16 (Siemens AG Medical Systems, Germany). 100 ml of intravenous contrast was administered at a rate of 3 ml s21 using an automated injector pump. Timing for the scan was influenced by the use of bolus triggering software (Siemens CARE Bolus) with a threshold of 100 HU and a region of interest positioned over the main pulmonary artery. Once triggered, a volume scan was undertaken following a 6 s delay through the whole chest caudocranially, using 779

C Hoskins and M Carpenter

a slice collimation of 16 mm 6 0.75 mm. Following raw data reconstruction (slice width of 1.0 mm, reconstruction interval of 0.7 mm) the resultant overlapping axial images were transferred to a workstation.

Data processing Case images were reviewed using the interactive axial, sagittal and coronal MPR display with any additional reformatting (angulation, curved MPR, magnification) undertaken when required. Examinations were first preselected for adequate contrast enhancement of the pulmonary arteries, which was judged subjectively. MPR images of normal arteries, arteries with adjacent lymphatic tissue and arteries with features of acute or chronic emboli were created and saved. Chronic embolic features included mural thrombus and web formation. From the selected images, pVRT or ‘‘fly through’’ views were acquired utilizing the standard post-processing software supplied with the scanner. The standard isotropic data set is manipulated to display 3D imagery by assigning different colour and degrees of transparency to different voxel Hounsfield Units within the volume. Whilst this rendering can be altered interactively by the user with regards to colour, opacity and brightness, our study employed the default setting. Further manipulation of the pVRT images is performed, locating areas of interest within the arteries by moving the virtual camera display on the MPR images. Subsequent views of the arterial lumen and emboli were obtained and saved.

image of this area using different virtual camera positions within the arterial lumen. Figure 2 shows a sagittal multiplanar reformat demonstrating an acute embolus appearing as a filling defect in the pulmonary arterial branch. The pVRT image clearly shows the occluding clot and an adjacent patent vessel. Figure 3 shows a chronic embolus appearing as a flattened eccentric defect in contrast filling at an obtuse angle with the vessel wall on the right side. There is a web in the left interlobar pulmonary artery. The pVRT image demonstrates the intraluminal web. Figure 4 shows chronic mural embolus in the left pulmonary artery on both coronal and sagittal MPRs, with the sagittal MPR demonstrating the position of the virtual camera. The pVRT image shows the intra-arterial nature of the embolus. Figure 5 shows both axial and coronal views of a follow up scan in a patient with long term perivascular lymphatic tissue unchanged from previous scans. PVRT imaging confirms that this perivascular tissue is not intra-arterial.

Discussion

Figure 1 shows an axial reconstruction from a CT pulmonary angiogram with a diagrammatic representation of the virtual camera observing a normal main pulmonary arterial bifircation; together with a pVRT

With the advent of 16 (and more) slice CT and subsequent acquisitions of near perfect isotropic volume data, there has been an explosion in the use of postprocessing techniques. These include maximum intensity projections (MIP), SSD, VRT, pVRT and MPR reconstructions which have benefited from a remarkable increase in resolution as well as software improvements, which have allowed them to be used easily and quickly with the minimum of training. Post-processing with CT pulmonary angiography includes the standard multiplanar reformats as well as curved or other reformats individually tailored to the arterial branch under study. MIP algorithms have been most commonly used for the assessment of vascular structures, extracting the highest attenuating voxels along a specified direction and adding this to the projected

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(b)

Results

Figure 1. Axial CT pulmonary angiography of normal main pulmonary bifurcation with volume rendered (pVRT) image. 780

The British Journal of Radiology, October 2006

Short communication: Virtual arterioscopy in pulmonary emboli

(a)

(b)

Figure 2. Sagittal multiplanar reconstruction of acute embolus with corresponding pVRT image.


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(c) Figure 3. Multiplanar reconstruction and perspective volume rendering (pVRT) images of chronic emboli and webs.

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Short communication: Virtual arterioscopy in pulmonary emboli

(a)

(b)

(c)

(d)

Figure 4. Multiplanar reconstruction and perspective volume rendering (pVRT) images of chronic embolus.

image, providing a display demonstrating the blood vessels well but suppressing the background tissue. Resultant images are rotatable, allowing elimination of overlap. However, unlike VRT and pVRT there is no sense of depth as the MIP image is essentially a collapsed 3D structure imposed onto a 2D surface without perspective [4]. Furthermore, as high attenuation suppresses low attenuation, this reconstruction technique can obscure low attenuation thrombus. Chronic embolic features such as webs and mural thrombus are also difficult to see even with relatively thin slice MIP imaging [3]. Characteristic CT features of chronic pulmonary thromboembolic disease include webs, bands, intimal irregularities and eccentric flattened defects at an obtuse angle with the vessel wall. Thrombi can be simulated by periarterial lymph nodes [8]. The British Journal of Radiology, October 2006

While to our knowledge this volume rendering imaging has not been studied in relation to pulmonary embolic disease, we feel it may prove useful in problem solving cases where the differential lies between intravascular but perimural pathology, such as chronic embolus, and extravascular pathology, such as bronchiolar impaction or periarterial lymphatic tissue, situated adjacent to the arterial wall. It may also prove useful in the display of the intravascular findings of chronic PE such as mural thrombus and webs, which are often extremely subtle [3]. Volume rendering can play a role in medical education with the improved visual presentation of disease processes. Several CT image reformatting approaches such as sagittal/coronal, oblique, curves and variable thickness viewing helps to orientate referring clinicians 783


(a)

(b)

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Figure 5. Multiplanar reconstruction of perivascular lymphatic tissue with confirmatory perspective volume rendering (pVRT) image of arterial lumen.

and radiologists to particular anatomical structures and pathology, allowing selective display and enhancement of relevant findings [3, 5]. However, these methods still require the user to think in sections whereas volume rendered imagery provides realistic three-dimensional pictures of intraluminal disease processes, which if studied on an interactive CT workstation can provide excellent learning opportunities to health care students.

References 1. Schoepf UJ, Goldhaber SZ, Costello P. Spiral computed tomography for acute pulmonary embolism. Circulation 2004;109:2160–7. 2. Reidel M. Diagnosing pulmonary embolism. Postgrad Med J 2004;80:309–19.

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3. Gruden JF, Tigges S, Baron MG, Pearlman H. MDCT pulmonary angiography: image processing tools. Semin Roentgenol 2005;40:48–63. 4. Edelman RR, Rabin DN. 3D imaging technologies. J Imaging Technology Management. www.ImagingEconomics.com, 2001. 5. Cody DD. AAPM/RSNA physics tutorial for residents: topics in CT. Imaging and Therapeutic Technology 2002;22:1255–68. 6. Van Ooijen PM, Oudkerk M, Van-Geuns RJ, Rensing BJ, DeFeyter PJ. Coronary artery fly through using electron beam computed tomography. Circulation 2000;102:E6–10. 7. Do-Yeon Kim, Jong-Wong P. Visualisation of the internal carotid artery using MRA images. Magn Reson Imaging 2005;23:27–33. 8. Han D, Kyung SL, Franquet T, Muller L, Sung Kim T, Hojoong K, et al. Thrombotic and nonthrombotic pulmonary arterial embolism: spectrum of imaging findings. RadioGraphics 2003;23:1521–39.



SHORT COMMUNICATION

Breast radiotherapy in women with pectus excavatum (funnel chest): is the lateral decubitus technique an answer? A dosimetric study M A BOLLET, MD, F CAMPANA, MD, Y M KIROVA, J-C ROSENWALD, PhD and A FOURQUET, MD

MD,

R DENDALE,

MD,

M-G SALIOU,

MD,

Department of Radiation Oncology, Institut Curie, Paris, France ABSTRACT. Breast radiotherapy is a technical challenge in women with pectus excavatum. We aim to assess isocentric lateral decubitus (ILD) technique as a means to irradiate breasts for patients with pectus excavatum. Four women presenting with leftsided breast cancers and found to have pectus excavatum were offered breastconserving treatments. Post-operative breast radiotherapy was indicated (50 Gy) in two patients, with an additional boost to the tumour bed (16 Gy). Both ILD and supine techniques were simulated. We report the dosimetric comparison of these techniques and the acute skin toxicity of ILD radiotherapy. ILD permitted the same breast dosehomogeneity as the supine technique while decreasing breast thickness by 4.5–6.8 cm. The width of lung and/or heart receiving . 20 Gy ranged between 2.1 cm and 4.3 cm with the supine technique and between 0.5 cm and 1.1 cm with ILD. The estimated percentage of ipsilateral lung receiving . 20 Gy ranged from 21% to 34% with the supine technique and from 0% to 5% with ILD. Acute skin toxicity was scored 1 for all patients at completion of ILD radiotherapy. ILD is an effective breast radiotherapy technique for patients with pectus excavatum that preserves the underlying heart and lung from unnecessary toxicity.

Breast conserving treatment is considered as standard management for early-stage breast cancer when technically feasible. Post-operative radiotherapy to the breast is an important part of the treatment and has been shown to significantly diminish the rate of local relapse, the rate of breast cancer mortality and the risk of overall mortality [1]. However, the irradiation of normal tissues, especially of the underlying lung and left ventricle, can lead to lethal complications explaining the increase in non-cancer related deaths. The risk of late toxicities such as radiation-induced pneumonitis or coronary morbidity has been related to the volume of, respectively, underlying lung and left ventricle receiving a dose higher than 20–25 Gy [2–4]. The standard breast radiation technique uses tangential fields delivered to the breast in a supine position [5]. In the case of women presenting with pectus excavatum, a congenital deformation of the anterior chest wall leading to a deep depression of the sternum, also described as funnel chest, this supine technique exposes too much of the underlying lung and/or left ventricle in the treatment fields. To try and comply with the two contradictory breast radiotherapy requirements of breast dose homogeneity and lung/heart preservation presents Address correspondence to: Marc A Bollet, MD, Institut Curie, Radiotherapy Department, 26, rue d’Ulm, 75005 Paris, France. E-mail: [email protected].


Received 6 January 2006 Revised 3 May 2006 Accepted 30 May 2006 DOI: 10.1259/bjr/23839243 ’ 2006 The British Institute of Radiology

a particularly difficult technical challenge that, up to now, has been dealt with by using sophisticated threedimensional, intensity-modulated radiotherapy (IMRT) or proton therapy techniques [6–8]. In this study we present a convenient solution: the isocentric lateral decubitus technique (ILD). Breast irradiation in the lateral decubitus set-up was originally developed for women with large, pendulous breasts and has been used to treat thousands of patients [9]. The need to comply with modern requirements has prompted our department to adapt the technique to an isocentric set-up as described by Campana et al [10]. To ensure a good reproducibility of position, the alignment of the patient is verified by using laser projections marked on the patient’s skin. The position of the epoxy breast-support is also marked on the patient’s skin. Day-to-day reproducibility is verified using the distance from the top of the treated breast to the corner of the epoxy breast-support. Portal imaging controls are regularly performed.

Methods and materials Four women presenting with cancer of the left breast and found to have pectus excavatum were offered breast-conserving treatments at the Institute Curie from October 2002 to September 2004. Surgery consisted of 785

M A Bollet, F Campana, Y M Kirova et al

lumpectomy with either axillary lymph node clearance or sentinel node excision. Breast irradiation alone was indicated (50 Gy at ICRU point in 25 fractions over 5 weeks) with, in the cases of patients A and B who were younger than 50 years old at the time of diagnosis, an additional boost to the tumour bed (16 Gy at ICRU point in 8 fractions over 1.5 weeks). The breast ICRU point was located at the central slice at the mid-bridge for tangents [11, 12]. Both ILD and supine breast radiation techniques were evaluated at the time of simulation in order to decide on the optimal treatment using dose distribution as a criterion. For practical reasons we decided to focus our interest on the whole-breast irradiation, common to the four patients. Full descriptions of the ILD technique and the supine technique can be obtained in the articles of Campana et al and Perez et al, respectively [5, 10]. They both comply with the recommendations of the EORTC quality assurance in conservative treatment of early breast cancer [13]. A schematic presentation of the ILD technique is reproduced in Figure 1. For illustrative purposes Figure 2 shows patient A in both the supine (Figure 2a) and the ILD (Figure 2b) positions. In short, the ILD technique consists of treating a patient in a lateral decubitus position with the treated breast resting on an individually chosen breast support made of a thin layer of carbon fibre and the contralateral breast brought upwards, out of the field, by elastic straps. For both techniques we used a simulator – CT (Varian Ximascan) to produce localization films of the fields and at least three CT slices, i.e. one at the central beam axis and one in each mid-field (z50; z5+25% of the field size; z5–25% of the field size). The CT images were transferred to the ISIS 3D treatment planning system. After having checked the position of the beams, dose distributions were calculated, without heterogeneity correction, using either 60-cobalt (source–axis distance (SAD) 80 cm) or 4 MV X-ray (SAD 100 cm). Whenever needed, standard wedges were added in order to

obtain the best possible dose distributions for both techniques. Measures on CT slices were done with the ISIS 3D measurement tool. The breast thickness was defined as the distance between the entry and the exit points of the fields at the central axis. The width of lung and/or left ventricle that received a dose equal to or higher than 20 Gy was measured on all CT-scan slices, on an axis perpendicular to that of the lateral field (Figure 3a). Figure 3 shows an example of two CT-scan slices from the same patient (patient B) in both the supine technique (Figure 3a) and the lateral decubitus technique (Figure 3b). The central lung distance (CLD) was measured in centimetres as the perpendicular distance from the field edge to the posterior part of the anterior chest wall at the centre of the field (Figure 4) [14]. The maximal heart distance (MHD) was measured in centimetres as the maximal width of heart in the tangent fields (Figure 4) [15]. The percentage of the ipsilateral lung volume and of heart receiving 20 Gy or more, known respectively as IPV20 and HV20, were estimated using Kong’s formulae [15]: IPV20 (%)~9:8|CLD (cm)

4:9

HV20 (%)~6:1|MHD (cm)

1:3

The acute skin toxicity of the breast alone radiotherapy was assessed on the last day of treatment (50 Gy for patients C and D and 66 Gy for patients A and B) and scored according to the Radiation Therapy Oncology Group (RTOG) acute radiation morbidity scoring criteria [16].

Results Dosimetry study Both techniques fulfilled the recommendations of the ICRU 62 report with a dose heterogeneity in the planning

Figure 1. Schematic presentation of the isocentric lateral decubitus (ILD) technique for radiotherapy of the left breast.

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Short communication: ILD breast irradiation – pectus excavatum

(a)

(b)

Figure 2. Pictures of the same patient with pectus excavatum at simulation in both (a) the supine and (b) the lateral decubitus set-ups.

target volume within a range corresponding to 95–107% of the prescription dose. In the case of patient D, the dose coverage of the breast with the supine technique was optimal when only 48 Gy were prescribed at the ICRU point in 24 fractions. Breast thickness ranged between 3.6 cm and 6.4 cm with the ILD technique and between 9.4 cm and 12 cm with the supine technique. The internal tangent fields’ simulation films of both techniques are shown in Figure 4. An overview of the results of the dosimetry study is reported in Table 1. The central lung distance ranged from 2.6 cm to 4 cm with the supine technique and from 0 cm to 1 cm with the ILD technique. The estimation of ILV20 ranged from 21% to 34% with the supine technique and from 0% to 5% with the ILD technique. The maximal heart distance ranged between 0 cm and 1.7 cm and the estimation of the HV20 from 0% to 9% with the supine technique. They were both zero with the ILD technique. The maximal width of lung and/or left ventricle that received 20 Gy or more ranged between 2.1 cm and 4.3 cm with the supine technique and between 0.5 cm and 1.1 cm with the ILD technique.

Acute toxicity No patient needed to have her radiotherapy suspended. Acute skin toxicity at the completion of treatment was mild and scored 1 (RTOG) for all four patients. The British Journal of Radiology, October 2006

Discussion Breast conserving treatments for breast cancers can sometimes be denied to women with pectus excavatum, where radiotherapy would involve too high a risk of pulmonary or cardiac toxicity. As previously discussed, the ILD is a simple, reproducible breast irradiation technique that can easily be implemented in radiotherapy departments [10]. More than 500 patients have already been treated using this technique in our centre. Most often, this technique was decided on in the case of patients with large pendulous breasts who needed breast alone irradiation. In this study, we address the possibility of using this technique for women with pectus excavatum who need breast-alone irradiation. We performed dosimetric studies on four women and report here the acute toxicity of this technique. Because the set-up device we used was too cumbersome to fit into our dosimetric CT-scan, we could not perform whole lung and heart scanning and thus dose–volume histogram comparisons were not available. We could nevertheless perform a number of CT slices that accurately reflected the dosimetry of both the supine and the ILD breast radiotherapy techniques. In addition, we used the two-dimensional information available from the simulation field to estimate the percentage of the ipsilateral lung and heart receiving 20 Gy or more [15]. The formula we used was generated by Kong et al from data of 40 patients simulated for breast radiotherapy (22 left side) in a supine position with a prescribed dose of 787

M A Bollet, F Campana, Y M Kirova et al

(a)

(b)

Figure 3. CT-scan slices at the level of the central axis (z50) of the same patient with pectus excavatum taken in (a) the supine and in (b) the lateral decubitus set-ups. Dosimetry of whole-breast radiotherapy was made for a prescribed dose of 50 Gy at midthickness using 4 MV photons.

46.8 Gy at the central slice at the mid-bridge for tangents. This formula is to be handled with caution when it comes to estimating the lung and heart volumetric doses in our patients. First, because we prescribed 50 Gy, i.e. 7% more than in Kong et al series. Second, because this formula was not meant to be applied to ILD. We assumed that, if anything, the results using this formula would underestimate the difference between the supine and the ILD techniques in terms of volumes of lung and heart receiving 20 Gy or more as the craniocaudal lengths of either lung or heart within the simulation field were diminished in the lateral position (data not shown).

Pectus excavatum is a rare condition and therefore this study suffers from a small number of patients making a formal statistical comparison of the two techniques inappropriate. We can nevertheless observe that, as expected, the change of position from the supine to the lateral decubitus position, with the breast spread out on an individually chosen dedicated breast support, made the thickness of breast smaller by 41–62% and thus also made the dose in depth more homogeneous. The breast took the shape of a parallelogram making it easier to obtain a homogeneous dose with the use of a single wedge (Figure 3b). The most important benefit is that the breast

Figure 4. Simulation films of the internal fields of the four patients presenting pectus excavatum both with the isocentric lateral decubitus (ILD) and the standard supine (Sup) radiotherapy techniques. Each graduation represents a centimetre. Central lung distance (CLD) is the perpendicular distance from the tangential field edge to the posterior part of the anterior chest wall at the centre of the field. Maximal heart distance (MHD) is the maximal width of heart in the fields. 788


Short communication: ILD breast irradiation – pectus excavatum Table 1. Dosimetric study comparing the isocentric lateral decubitus (ILD) and the supine breast radiotherapy techniques in four women presenting with pectus excavatum. Dosimetries of whole-breast radiotherapy were made for a prescribed dose of 50 Gy at mid-thickness using either 60-Cobalt gammas or 4 MV photons Patient

A B C D*

Breast thickness (cm)

Central lung distance (cm)

20 Gy ipsilateral lung volume (%)

Maximal heart distance (cm)

20 Gy heart volume (%)

Maximal dose to breast (Gy)

Supine

ILD

Supine

ILD

Supine

ILD

Supine

ILD

Supine

ILD

Supine

ILD

10.9 12.0 9.4 11.4

6.4 5.2 3.6 5.6

2.6 4.0 3.2 2.7

1.0 0.0 0.0 0.4

21 34 26 22

5 0 0 0

0.0 1.7 0.0 0.8

0.0 0.0 0.0 0.0

0 9 0 4

0 0 0 0

52 54 53 54

52 53 52 52

*For the dosimetry study of this patient, we chose to reduce the prescribed dose to 48 Gy in 24 fractions with the supine technique in order to keep the maximum dose below 54 Gy (this patient has however been treated with the ILD technique to a total dose of 50 Gy in 25 fractions).

tissue is located at a greater distance from the chest wall, making it easier to encompass the whole breast within the radiation field without taking in any of the heart and less than 1 cm of the underlying lung. The resultant estimated percentage of ipsilateral heart and lung are thus reduced to a maximum of 0% and 5%, respectively. In contrast to this, the standard supine position had central lung depths of at least 2.5 cm resulting in the estimated volume of lung receiving 20 Gy always being greater than 20%. The protocols in use at the Institut Curie define the maximum depth of lung and/or left ventricle receiving greater than 20 Gy to be 2 cm, thus the standard supine technique would be unacceptable. IMRT has been proposed as a means to decrease the volume of the ipsilateral lung receiving a radiation dose that is above the tolerance threshold [6, 7]. Fogliata et al have suggested that three-field conformal radiotherapy technique permitted a reduction of the ipsilateral lung volume receiving more than 20 Gy from 24% with twofield tangential treatment to less than 20%. This reduction was the same whether using IMRT or not [8]. The drawback of IMRT and the 3D conformal technique described by Fogliata et al, contrary to the ILD technique, is that it increases the volume of surrounding normal tissues (the heart, spinal cord, and contralateral breast and lung) receiving low-dose irradiation compared with the supine technique [6]. The other advantage the ILD technique has over a three-dimensional radiotherapy approach is its cost-effectiveness, both financially and time-wise. However, we have to bear in mind that the ILD technique is not designed to treat the breast in the case of a treatment with radiotherapy to the lymph node bearing areas; in such cases, IMRT with or without active breathing control, or even proton therapy [8], could be considered. The acute skin toxicity of the ILD technique was good, with no need for a suspension of treatment and only a grade I RTOG skin toxicity in all of our four patients of whom two were treated with a boost of up to 66 Gy.

Conclusions The ILD technique is an easy, cost-effective means to offer breast preservation to patients with pectus excavatum, without exposing the underlying heart and, more particularly, ipsilateral lung to unnecessary toxicity. The British Journal of Radiology, October 2006

Acknowledgments We thank J Y Kristner for designing the treatment setup device, Chantal Dauphinot and Yann Brunet for their help in the dosimetry study and all the members of the Groupe Sein at the Institut Curie for their participation in the care of the patients.

References 1. Clarke M, Collins R, Darby S, et al. Effects of radiotherapy and of differences in the extent of surgery for early breast cancer on local recurrence and 15-year survival: an overview of the randomised trials. Lancet 2005;366:2087–106. 2. Lind PA, Pagnanelli R, Marks LB, et al. Myocardial perfusion changes in patients irradiated for left-sided breast cancer and correlation with coronary artery distribution. Int J Radiat Oncol Biol Phys 2003;55:914–20. 3. Graham MV, Purdy JA, Emami B, et al. Clinical dosevolume histogram analysis for pneumonitis after 3D treatment for non-small cell lung cancer (NSCLC). Int J Radiat Oncol Biol Phys 1999;45:323–9. 4. Yu X, Prosnitz RR, Zhou S, et al. Symptomatic cardiac events following radiation therapy for left-sided breast cancer: possible association with radiation therapy-induced changes in regional perfusion. Clin Breast Cancer 2003;4:193–7. 5. Perez CA, Garcia DM, Kuske RR. Breast: Stage Tis, T1, and T2 tumors. In: Company JL, editor. Principles and practice of radiation oncology. Second Edition. Philadelphia: JB Lippincott Company, 1994:877–947. 6. Teh BS, Lu HH, Sobremonte S, et al. The potential use of intensity modulated radiotherapy (IMRT) in women with pectus excavatum desiring breast-conserving therapy. Breast J 2001;7:233–9. 7. Thilmann C, Zabel A, Kuhn S, et al. Inversely planned intensity modulated radiotherapy for irradiation of a woman with breast cancer and funnel chest. Strahlenther Onkol 2002;178:637–43. 8. Fogliata A, Bolsi A, Cozzi L. Critical appraisal of treatment techniques based on conventional photon beams, intensity modulated photon beams and proton beams for therapy of intact breast. Radiother Oncol 2002;62:137–45. 9. Fourquet A, Campana F, Rosenwald JC, et al. Breast irradiation in the lateral decubitus position: technique of the Institut Curie. Radiother Oncol 1991;22:261–5. 10. Campana F, Kirova YM, Rosenwald JC, et al. Breast radiotherapy in the lateral decubitus position: a technique to prevent lung and heart irradiation. Int J Radiat Oncol Biol Phys 2005;61:1348–54.

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M A Bollet, F Campana, Y M Kirova et al 11. ICRU. Report 62: Prescribing, recording, and reporting photon beam therapy (supplement to ICRU Report 50). Bethesda, MD: International Commission on Radiation Units and Measurements; 1999. 12. ICRU. Report 50: Prescribing, recording, and reporting photon beam therapy. Washington DC: International Commission on Radiation Units and Measurements; 1993. 13. Bartelink H, Garavaglia G, Johansson KA, et al. Quality assurance in conservative treatment of early breast cancer. Report on a consensus meeting of the EORTC Radiotherapy and Breast Cancer Cooperative Groups and the EUSOMA (European Society of Mastology). Radiother Oncol 1991;22:323–6.

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14. Bornstein BA, Cheng CW, Rhodes LM, et al. Can simulation measurements be used to predict the irradiated lung volume in the tangential fields in patients treated for breast cancer? Int J Radiat Oncol Biol Phys 1990;18:181–7. 15. Kong FM, Klein EE, Bradley JD, et al. The impact of central lung distance, maximal heart distance, and radiation technique on the volumetric dose of the lung and heart for intact breast radiation. Int J Radiat Oncol Biol Phys 2002;54:963–71. 16. Cox JD, Stetz J, Pajak TF. Toxicity criteria of the Radiation Therapy Oncology Group (RTOG) and the European Organization for Research and Treatment of Cancer (EORTC). Int J Radiat Oncol Biol Phys 1995;31:1341–6.



REVIEW ARTICLE

Hypoxia in head and neck cancer 1

A Y ISA, MBChB, MRCS, M Phil, 2T H WARD, and 5J J HOMER, MD, FRCS (ORL-HNS)

PhD,

3

C M L WEST,

BA, PhD,

4

N J SLEVIN,

MBChB, FRCP, FRCR.BW

1

Department of Surgery, Christie Hospital, Manchester, 2Drug Development Group, Paterson Institute for Cancer Research, Manchester, 3Academic Department of Radiation Oncology, University of Manchester, Manchester, 4Department of Clinical Oncology, Christie Hospital, Manchester and 5Departments of Surgery, Christie Hospital and University Department of ORL-HNS, Manchester Royal Infirmary, Manchester, UK ABSTRACT. A high level of hypoxia in solid tumours is an adverse prognostic factor for the poor outcome of cancer patients following treatment. This review describes the status of research into finding a practical method for measuring hypoxia and treating hypoxic tumours. The application of such methodology would enable the selection of head and neck cancer treatment based on an individual’s tumour oxygenation status. This individualization would include the selection not only of surgery or radiotherapy, but also of novel hypoxia-modification strategies.

Hypoxia, the inadequate supply of blood-borne oxygen, has been linked with a poor tumour response to radiotherapy [1–4]. This resistance relates, in part, to the radiobiological effects of hypoxia. However, hypoxia is also a marker of an aggressive tumour phenotype and is associated with a poor outcome following surgery [5, 6]. Furthermore, hypoxia is implicated in the resistance of tumours to some chemotherapeutic agents [7]. This review discusses the biology of tumour hypoxia, methods for measuring tumour hypoxia, possible clinical applications of hypoxia measurements and potential hypoxia modification approaches for patients with head and neck cancer.

The biology of tumour hypoxia Tumour cells require a host vasculature for their supply of nutrients and oxygen, but oxygen cannot diffuse further than around 150 mm through tissues. As tumour growth outstrips its vasculature, the cells become hypoxic [8]. This hypoxia, termed chronic or diffusion-limited hypoxia, occurs adjacent to areas of necrosis. Tumour hypoxia also arises from the transient or intermittent occlusion of tumour blood vessels, called acute or perfusion-limited hypoxia. Although the mechanisms behind the development of acute hypoxia are unknown, it may arise from the blocking of vessels This work was supported by the Wolfson Foundation, the Christie Hospital Endowment Fund and the National Translational Cancer Research Network. Address correspondence to: Mr J J Homer, University Department of Otolaryngology-Head and Neck Surgery, Manchester Royal Infirmary, Oxford Road, Manchester M13 9DL, UK.


Received 21 July 2005 Revised 19 April 2006 Accepted 22 May 2006 DOI: 10.1259/bjr/17904358 ’ 2006 The British Institute of Radiology

by circulating blood and tumour cells, blood vessel collapse from high tumour interstitial pressure, and/or the interruption of tumour blood flow. All these factors could occur because of the chaotic, disorganized and fragile nature of tumour compared with normal tissue vasculature [9, 10]. Thomlinson and Gray first proposed the existence of hypoxia in human tumours in 1955 [11] and confirmed the radioresistance of hypoxic mammalian cells, which was described in 1936 by Mottram [12]. Thomlinson and Gray suggested that the presence of quiescent but viable hypoxic cells, capable of re-oxygenation during fractionated radiotherapy, would limit the success of treatment. This suggestion spurred research into finding methods for measuring tumour hypoxia and for overcoming hypoxic cells in patients undergoing radiotherapy. It is only in the last decade, however, that hypoxia has emerged as a key factor in driving malignant progression. Hypoxia is involved in the transcriptional regulation of a number of genes. Hypoxia-inducible factor-1a (HIF-1a) and HIF-2a are transcription factors that mediate cell responses to hypoxia. Cells express HIF-1a continuously, but the protein degrades rapidly under normoxic conditions. In hypoxia, the protein is stabilized and induces the transcription of a number of genes including those involved in angiogenesis, glycolysis, pH control and oxygen delivery. HIF-1a transcribed genes include the key angiogenic growth factor, vascular endothelial growth factor (VEGF) [13], as well as glucose transporter 1 (Glut-1) [14] and the pH regulator, carbonic anhydrase IX (CAIX) [15]. Thus, cellular response to hypoxia strives to prevent cell death. HIF-1a upregulation occurs within 2 min of hypoxia, accumulating rapidly over 30 min to peak after another 30 min [16]. 791

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Clinical importance of hypoxia The biological effect of radiation depends on the degree of oxygenation, and hypoxic cells are approximately three-fold more resistant than well-oxygenated cells. This oxygen effect is due to the interaction between oxygen and the free radicals produced when radiation is absorbed in tissues. Radiation absorbed in tissues produces highly reactive, short-lived free radicals, which produce double strand breaks in DNA leading to cell death. The oxygen increases the damage produced by radiation by increasing the lifetime of the free radicals. Due to the short life-span of the free radicals, oxygen needs to be present at the time of irradiation to be effective [17]. Studies showed poorer locoregional control and survival in patients with hypoxic compared with oxygenated head and neck squamous cell cancers treated with radiation [1–4]. There is evidence that hypoxia limits the effectiveness of not only radiotherapy but also surgery. In studies measuring tumour oxygenation using polarographic oxygen microelectrodes, tumours treated with primary surgery with or without radiation had a poor outcome [5, 6]. Hockel et al found the median partial pressure of oxygen (pO2) to be the strongest independent predictor of overall and disease-free survival in patients with locally-advanced uterine cervical cancer irrespective of treatment modality. Similar findings reported with high-grade non-metastatic soft tissue sarcomas [6] indicate the involvement of factors other than hypoxia-mediated radioresistance. There is evidence that hypoxia enhances genetic instability in tumour cells and selects for tumour cell populations with diminished apoptotic potential, and increased aggressiveness and metastatic potential [17–19]. Hypoxia has also been implicated in the resistance of tumours to some chemotherapeutic agents, including those used in the treatment of head and neck cancers [7]. Hypoxic cells are resistant to 5-fluorouracil, doxorubicin, bleomycin and cisplatin [20–23]. Although the mechanisms behind this drug resistance are not understood, reduced cellular proliferation, low pH and hypoxiainduced alterations in gene expression may play a role.

Measuring hypoxia Various methods are under development for measuring tumour hypoxia in cancer patients. There is a need for a method that is practical, fast and reliable, i.e. suitable for routine clinical application. Although intratumour variability in oxygenation is a potential confounding factor affecting many of the measurement strategies, tumour-totumour variability is greater [17]. The ability to differentiate between necrotic and hypoxic areas is also important, but not possible with all techniques. Whilst the ultimate result of hypoxia is necrosis (and therefore necrotic cells might be a measure of the process of hypoxia), it is the measurement of hypoxic cells that remain viable that are of the greatest clinical importance.

Oxygen microelectrodes The current resurgence of interest in tumour hypoxia stems from data obtained during the 1990s using 792

computerized Eppendorf oxygen microelectrodes. Advantages of Eppendorf over older style electrodes were the use of fine needles to minimize tissue trauma, an automatic stepper motor to enable rapid movement through tissue and to avoid tissue compression artefacts and computerization to enable the quick collection of multiple measurements. The electrodes are used with or without image guidance, depending on the location and size of the tumour. Le et al showed that median pO2 readings from advanced head and neck cancer nodal metastases were significantly lower than normal subcutaneous tissue (14.6 mmHg vs 51.2 mmHg; p,0.001) and in each patient, median tumour metastases pO2 was consistently lower than that of normal subcutaneous tissue. 40% of the study group had median pO2 measurements of less than 10 mmHg [24]. Nordsmark et al [3] found in their study of pre-treatment oxygenation in neck metastasis of advanced squamous cell carcinoma of the head and neck that the fraction of pO2,2.5 mmHg was significant as a continuous variable of local failure following radiotherapy, but median pO2 was not. Another study of head and neck squamous cell carcinoma (HNSCC) treated with chemoradiotherapy/ radiotherapy showed that the hypoxic subvolume defined as the percentage of tumour multiplied by percentage of pO2-values below 5 mmHg were significant in multivariate analysis for poorer overall survival where no correlation was found with median pO2 [18]. The approach, however, is limited to tumours that are accessible. Also, oxygen electrodes cannot differentiate viable hypoxic regions from necrotic areas [25, 26].

Comet assay The comet assay uses individual tumour cells dissociated from needle biopsies. It is a sensitive method of measuring DNA strand breaks in tumour cells after single doses of 3.5–10 Gy. The approach is based on the fact that ionizing radiation produces around three-fold more damage in well-oxygenated compared with hypoxic cells (i.e. it measures radiobiologically hypoxic cells) [27, 28]. The comet assay was compared with the oxygen microelectrode system in a group of patients with heterogeneous advanced tumours (including nonHNSCC). Aquino-Parsons et al found a correlation between hypoxic fraction measured by comet assay and percentage of pO2 values ,5 mmHg measured by the oxygen microelectrode (r250.46, p,0.001). In their study, tumours defined as hypoxic with a median pO2,10 mmHg were found to have .20% radiobiologically hypoxic cells as measured by the comet assay [28]. Another study compared the two methods in node positive Stage III–IV HNSCC. The majority had Stage IV disease and were randomized to cisplatin-based induction chemotherapy followed by concurrent chemoradiotherapy with or without tirapazamine. A weak but statistically significant negative correlation between oxygen microelectrode median pO2 and 1 min median tail moment as a measure of DNA damage postirradiation from the comet assay was found (r25 0.08, p50.05). There was a significant correlation between response and comet median tail moment (p50.001); The British Journal of Radiology, October 2006

Review article: Hypoxia in head and neck cancer

however, no correlation was found between response and median pO2 readings [24]. The lack of correlation between median pO2 and response is not unexpected as Nordsmark et al had described this in an earlier study [3]. The use of median pO2 instead of percentage pO2,5 mmHg may also explain the negative correlation between the oxygen microelectrode and the comet assay. The advantage of the comet assay lies in its ability to measure radiobiologically hypoxic cells only and not necrosis. Radiation doses in excess of 3.5 Gy are required to produce sufficient DNA damage to distinguish a hypoxic subpopulation, thus limiting its application to patients receiving larger fractions or doses. Contamination of the fine needle aspirate sample by circulating white blood cells (which are unirradiated) could falsely reduce the median tail moment making the tumour seem more hypoxic [28, 29].

Using nitroimidazoles as hypoxic markers Nitroimidazoles are nitro aromatic compounds that bind to hypoxic cells. Nitroimidazoles diffuse easily into hypoxic cells due to their high solubility and low metabolism [27]. A non-therapeutic dose of a nitroimidazole is administered, systemically prior to biopsy or resection of a tumour, and hypoxic cells in histological sections can be identified using immunohistochemistry (IHC), flow cytometry and immunofluorescence. Nitroimidazoles used as hypoxic markers include misonidazole, pimonidazole, etanidazole, EF5 and nitroimidazole-theophylline [17, 30]. An advantage of the approach is its applicability to tumours inaccessible to oxygen electrodes. However, a biopsy might not be representative of the heterogeneity of hypoxia within a whole tumour. Although no relationship was found between pimonidazole binding and oxygen electrode data in carcinoma of the cervix [31], a study in head and neck cancer showed that high pimonidazole binding did predict adverse treatment outcome. Locoregional tumour control was statistically significantly lower for patients who had hypoxic tumours with 2-year control rates of 48% vs 87% for tumours with high and low pimonidazole binding levels, respectively [32]. This finding raises the possibility that measurements of pimonidazole binding in head and neck cancer might be useful as a selection tool for hypoxia-modifying treatments. However, pimonidazole staining in highly differentiated or keratanized tumour tissue has been described. The question as to whether these areas are hypoxic has been raised. This is especially pertinent in head and neck cancers that have a considerable amount of keratinization as pimonidazole staining of keratinized but non-hypoxic areas would lead to an overestimation of the level of hypoxia. Janssen et al found in 25% of their head and neck cancer series up to 30% of staining was in well-differentiated areas [33]. Further investigation is required.

Measuring endogenous markers of hypoxia Hypoxia-inducible proteins are under investigation as potential endogenous markers of hypoxia. The The British Journal of Radiology, October 2006

expression of HIF and the proteins it regulates transcriptionally can be measured using IHC [34]. The advantage of the approach for routine clinical use is that, as for nitroimidazole probes, only a biopsy is required. Disadvantages are that a single biopsy might not be representative of a whole tumour and that the proteins investigated can be upregulated by factors other than hypoxia, i.e. they are not hypoxia specific. Potential endogenous markers of hypoxia investigated include HIF-1a, HIF-2a, Glut-1 and CAIX. Strong HIF-1a expression was associated with a significantly poorer outcome following radiotherapy in oropharyngeal tumours expressing HIF-1a [35]. High HIF-1a and HIF-2a expression was associated with poor outcome in patients with advanced squamous cell carcinoma of the head and neck treated with concurrent carboplatin chemoradiotherapy [36]. Complete response was lower in tumours expressing HIF-1a and/or HIF-2a 49% vs 86% (p50.004) and high expression of HIF-1a and HIF-2a was significantly associated with poor local relapse-free survival (p50.003 and p50.003, respectively) and overall survival (p50.05 and p50.001, respectively) in univariate analysis. HIF-2a over expression was also an adverse prognostic factor for locoregional control (p50.002) and overall survival (p50.0004) in head and neck cancers treated with accelerated radiotherapy [37]. However, the expression of HIF-1a in surgically treated head and neck squamous cell carcinoma was associated with improved diseasefree survival (p50.016) and overall survival (p50.027) where no difference in outcome was seen with HIF-2a expression [34]. A more recent study of T1/2 squamous cell carcinomas of the floor of mouth confirmed the positive prognostic effect of HIF-1a expression in surgically treated patients. In this series of 85 patients, HIF-1a expression was associated with a significantly improved 5-year survival rate (p,0.01) and a significantly increased disease-free period (p50.01) [38]. It remains to be established whether the importance of the expression a hypoxia-inducible marker that is upregulated by factors other than hypoxia will vary with disease stage, but clearly this is an area that needs further study. Glut-1 is a ubiquitously expressed facilitative glucose transporter that is over-expressed in a number of tumours [39]. Although Glut-1 is upregulated by a variety of agents and conditions, it has potential as an intrinsic marker of hypoxia because of dual control of expression in hypoxic conditions via HIF-1 and reduced oxidative phosphorylation [14]. Its expression is heterogeneous in cell cytoplasm and membrane, and overexpression has been shown in tumours including those of the breast [40], lung [41], thyroid [42], cervix [39], hypopharynx [43] and oral cavity [44]. Glut-1 staining tends to localize in necrotic and perinecrotic areas of tumours [39]. Over expression of Glut-1 was associated with a poor prognosis in oral squamous cell carcinoma [44, 45] and hypopharyngeal carcinoma [43]. For example, survival times were 138 months and 60 months for patients with surgically treated oral squamous cell carcinoma expressing low versus high levels of Glut-1, respectively [44]. Recent experiments on human tumourderived xenografts revealed that the chemosensitivity of certain alkylating agents may be influenced by Glut-1 expression [46]. However, there is only a very weak [39] or no [47] relationship between tumour Glut-1 793


expression and oxygen variables (pO2, hypoxic fractions 2.5 and 5). The latter, along with the lack of prognostic significance of Glut-1 expression in multivariate analysis, has raised questions on the suitability of Glut-1 as an endogenous hypoxia marker [47]. CAIX is a transmembrane glycoprotein that is induced by hypoxia via the HIF-1 pathway and may be a useful endogenous marker of hypoxia [15]. Expression of CAIX in hypoxic tumours is localized to perinecrotic regions and is thought to confer a survival advantage by maintaining intracellular pH [15]. CAIX expression correlated with tumour oxygenation status measured using oxygen electrodes in cervical squamous cell carcinomas and is associated with poor outcome following radiotherapy [48]. CAIX expression is associated with poor complete response rate (40% vs 70%, p50.02) to chemoradiotherapy in advanced squamous head and neck cancer [49]. Interestingly, the co-expression of HIF1a and CAIX were associated with a poorer progressionfree survival (p50.04) in chemoradiated locally advanced nasopharyngeal cancer than the expression of either marker alone [50]. The combined expression of CAIX and Glut-1 in HNSCC treated radiotherapy with or without chemotherapy was also associated with poorer local control (p50.02) and disease-free survival (p50.04) where expression of either marker alone was not significantly correlated with outcome [51]. More recently, interest has grown in osteopontin – an extracellular matrix protein involved with tumour cell invasion, migration, angiogenesis and tumour growth [52, 53]. Osteopontin is secreted in plasma making it an attractive, easy method of assessing hypoxia. Plasma osteopontin levels were inversely correlated with pO2 (p50.003, r5 –0.42) in HNSCC and osteopontin was an independent predictor for freedom from relapse and survival on multivariate analysis [54]. The prognostic significance of osteopontin in HNSCC treated by radiotherapy was supported by Overgaard et al [55]. Moreover, the prognosis of patients with high levels of osteopontin was improved by addition of hypoxia modifier nimorazole. Locoregional failure (RR 0.19 [95% CI 0.08–0.44]) and disease specific mortality (RR 0.25, 95% CI 0.11–0.59) was more frequent in patients with high concentration of osteopontin assigned placebo than those assigned nimorazole [55]. Hypoxic regions within solid tumours will be not only oxygen deficient, but also acidic and nutrient depleted. Recent studies indicating that glucose concentration [56] and pH [57] affect the expression of endogenous markers of hypoxia raise concern over the generalized use of these proteins or genes as hypoxic markers. In vitro experiments using FADU human pharyngeal carcinoma and HT1080 human fibrosarcoma cells showed a lack of hypoxic HIF-1a accumulation in glucose-depleted conditions. Serum-depleted conditions also caused decreased hypoxic HIF-1a accumulation in FADU cells despite the presence of glucose [56]. Even though Vordermark et al [56] showed no effect of pH on HIF1a expression, Sorensen et al [57] found that the expression of CA9, GLUT1 and Osteopontin (OPN) genes were inhibited in hypoxia when extracellular pH was reduced to 6.3. As CA9 and GLUT1 are HIF-1 regulated genes it is likely that HIF-1 will be affected by pH as well. A possible explanation for the findings of 794

Vordermark et al may be that the lowest pH level of 6.7 was not acidic enough to show any inhibitory effect. The time course and the levels to which the genes were upregulated were different, as was the oxygen concentration at which upregulation was maximal, e.g. after 24 h exposure CA9 upregulation was maximal at 1% oxygen whereas GLUT1 and OPN were maximal at 0.01% and 0% oxygen, respectively [57]. More work is required to reveal any other factors that may effect these proteins or genes that have until now been loosely termed endogenous markers of hypoxia. Perhaps the term ‘‘markers of poor prognosis’’ may be more apt. It may be that instead of utilizing a single protein, tumours should be tested for multiple markers to create a hypoxia-associated molecular profile.

Non-invasive imaging 2-Fluoro-2-deoxy-D-glucose (FDG) positron emission tomography (PET) utilizes the concept that most tumours exhibit accelerated glycolysis allowing 18Flabelled FDG to be trapped in tissues with a higher metabolic rate than normal tissues [58]. There are several studies with contradictory findings with regards the correlation of Glut-1 expression and 18F-FDG uptake in tumours [59–62]. Most studies compared standardized uptake ratios or values from pre-operative 18F-FDG PET scans and compared them with immunohistochemical Glut-1 expression from a section of the tumour or a biopsy specimen. Interestingly, a recent article based on a rat tumour model showed a good correlation between 18 F-FDG uptake and Glut-1 expression (r50.829, p,0.001) at a micro-regional level [60]. Even so, the ability of 18F-FDG to detect tumour hypoxia is questionable [63, 64]. The imaging of radiolabelled nitroimidazoles by single-photon emission computer tomography (SPECT) and PET as a non-invasive measure of hypoxia is also being studied, amongst them 18F-fluoromisonidazole (18FMISO) [63], 18F-fluoroerythronitroimidazole (18F-FETNIM) [65], 18F-fluoroetanidazole (18F-FETA) [66], and 18F-EF3 [67]. Metal labelled hypoxia markers are also currently being developed, e.g. Cu ATSM [68]. Preliminary clinical testing of some of these markers has been promising with small series indicating a prediction of radiotherapy outcome [65, 69]. Advantages of the PET based approach include the ability to assess hypoxia of the whole tumour using a non-invasive technique that evaluates viable regions within the tumour. Deep lesions that are otherwise inaccessible to the microelectrodes can be studied and sequential scans can indicate response to treatment. A disadvantage is the limited spatial resolution. Cross sectional imaging methods are also being studied, in particular MRI. The methods are not hypoxia specific, but the wide availability of MRI is an advantage of the approach. Dynamic contrast-enhanced MRI has been the most widely studied method. The method yields parameters that reflect, amongst other parameters, tumour perfusion and vascularity. There is some evidence that the approach can yield data that reflect tumour oxygenation [70], and predict radiotherapy outcome in patients with carcinoma of the cervix and advanced head and neck cancer [71, 72]. The British Journal of Radiology, October 2006

Review article: Hypoxia in head and neck cancer

Hypoxia modification strategies There are a number of approaches under investigation to improve the oxygenation status of tumours. Oxygen delivery to a tumour can be raised by increasing the oxygen content of inspired air or raising haemoglobin levels. There are also hypoxia specific cytotoxins such as bioreductive drugs. For tumours treated with radiation, hypoxic cell radiosensitizers and radiotherapy modulation strategies are being developed.

Increasing oxygen delivery One of the earliest hypoxia modification strategies studied was hyperbaric oxygen. Although breathing oxygen during radiotherapy was used to some success in the past, delivery was complex and patient tolerance poor. Carbogen (2–5% carbon dioxide and oxygen) was suggested as an alternative to pure oxygen because of it causing vasodilation, an increase in respiratory drive and a right shift in the oxygen-haemoglobin dissociation. Patients breathe carbogen through a face mask during their radiotherapy. Accelerated radiotherapy combined with carbogen and nicotinamide (ARCON) is being investigated in patients with head and neck cancer. The accelerated radiotherapy targets rapidly proliferating tumour cells, the carbogen maximizing arterial pO2 and nicotinamide is a vasodilator that aims to minimize diffusion-limited hypoxia. The radiosensitizing effect of ARCON can also be expected in normal tissue hence the need for reduced doses to the spinal cord and laryngeal cartilage [73, 74] and the importance of its selective use based perhaps on hypoxia measurement. So far, results with ARCON in clinical trials in squamous cell carcinoma of the head and neck have been mixed [74, 75]. How much additional benefit with ARCON is due to hypoxia reduction as opposed to reduction in tumour clonogenic repopulation by accelerated fractionated radiotherapy has not been ascertained.

Raising haemoglobin levels Anaemia is associated with a poor outcome following chemoradiotherapy and radiotherapy [18, 76, 77]. The detrimental effect of anaemia on radiotherapy outcome has led to the routine use of blood transfusions prior to radiotherapy in patients with low haemoglobin levels. An alternative approach under investigation is the use of erythropoietin. Erythropoietin increases red cell production and theoretically reduces tumour hypoxia. In a retrospective study of patients with oropharyngeal cancer treated with chemoradiotherapy prior to surgical resection, Glaser et al found that a low pre-operative haemoglobin level was an independent adverse prognostic factor for locoregional control and survival [76]. This was reversed with the use of recombinant human erythropoietin during chemoradiotherapy. However, a recent double-blind, placebo-controlled randomized trial using erythropoietin in head and neck cancer patients undergoing radiotherapy did not show any improvement in treatment outcome and actually suggested poorer locoregional progression-free survival [78]. The British Journal of Radiology, October 2006

Further carefully controlled trails are required to study this more closely.

Hypoxia specific cytotoxins Bioreductive drugs are reduced under hypoxia or by reducing enzymes to cytotoxic metabolites [79]. Quinone (e.g. mitomycin C, porfiromycin), nitroimidazole (e.g. RSU-1069) and N-oxide (e.g. tirapazamine) compounds can be bioreduced to cytotoxic species. Although mitomycin C was used to treat head and neck cancer patients before it was known to be a bioreductive agent, it is not hypoxia selective and there is little differential cell killing between oxygenated and hypoxic cells. Of the new agents developed for their preferential toxicity toward hypoxic cells, tirapazamine is probably of most interest. Tirapazamine exhibits selective cytotoxicity for hypoxic cells. It is reduced in hypoxia to a highly reactive radical, which is capable of causing single and double strand DNA breaks. In normoxia this radical is back oxidized to the inert parent compound. This cycle is thought to confer hypoxic selectivity [80]. Tirapazamine can be given with other cytotoxic agents or with radiotherapy. A phase I trial of fractionated radiotherapy with concomitant tirapazamine and cisplatin in advanced oropharyngeal cancer, found dose limiting toxicity in the form of febrile neutropenia. This was overcome by omitting tirapazamine in weeks 5 and 6. Measurement of hypoxia by 18F-misonidazole PET scanning revealed a decrease in tumour hypoxia following treatment [81]. A recent randomized phase II trial comparing tirapazamine, cisplatin and radiation versus fluorouracil, cisplatin and radiation in 120 stage III and IV HNSCC patients found a trend in favour of the tirapazamine arm for both locoregional control (84% vs 66%) and failure-free survival (55% vs 44%). This became statistically significant for locoregional control following adjustment for known prognostic factors [82]. Gene therapy strategies are also under investigation to target hypoxic cells in solid tumours. These include using anaerobic bacteria, e.g. Clostridium Spp as a vector to deliver genetic material preferentially to hypoxic tissue at sufficient levels to cause a therapeutic effect [79]. HIF-1a downregulation by intratumoural delivery of antisense HIF-1a plasmid showed promising results [83].

Hypoxic cell radiosensitizers As discussed above, nitroimidazoles can diffuse into hypoxic cells due to their high solubility and low metabolism. The radiosensitizing effects of these compounds are related to their electron affinity, mimicking the oxygen effect, increasing DNA damage and restoring radiosensitivity. The use of first generation nitroimidazoles (e.g. misonidazole) was limited due to intolerable side effects including irreversible peripheral neuropathy. However, nimorazole, a 5-nitroimidazole derivative, has fewer side effects [84]. A phase III trial in Denmark showed significantly increased locoregional control in supraglottic and pharyngeal tumours treated with nimorazole and conventional radiotherapy [77]. 795


Improving the delivery of radiotherapy Radiotherapy can be modulated to deliver more radiation to hypoxic tumours. Intensity-modulatedradiation therapy (IMRT) and three-dimensional conformal radiation therapy (3D-CRT) [85] enables small irregular volumes to be ‘‘painted’’ with radiotherapy. With better methods of detecting hypoxic microregions in the future, possibly with PET and SPECT, it may be possible to select hypoxic regions within tumour for treatment with higher radiation doses using these techniques.

The future: individualizing treatment according to hypoxic status The choice of treatment for patients with head and neck cancer is often a matter of clinician and centre preference. There is an unmet need for biological parameters to individualize treatments. The various ways of measuring hypoxia provide potentially promising ways of predicting response to radiotherapy. This may allow clinicians to choose primary surgery or (chemo)radiotherapy on a scientific, individualized basis, rather than by protocol and acumen. This individualization of treatment has potential to improve patient management in two ways. First, a decrease in radiation failures would reduce the need for salvage surgery, which is associated with greater surgical morbidity than primary surgery [86]. Second, there would be a more appropriate selection of patients for organ preserving chemoradiotherapy. Furthermore, there are number of specific hypoxia modification strategies that can be applied to patients on an individualized basis. Progress in the development of a tool for measuring tumour hypoxia should enable the future informed clinical choice between primary (chemo)radiotherapy, (chemo)radiotherapy with hypoxia modification strategies or radical surgery for locally advanced tumours.

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Is routine chest radiography a useful test in the follow up of all adult patients with soft tissue sarcoma? 1

H K LORD, MRCP, 2D M SALTER, 1 G R KERR, MSc

MD, FRCPath, FRCPE,

1

R H MACDOUGALL,

FRCS, FRCR, FRCPE

and

1

Department of Clinical Oncology, Edinburgh Cancer Centre, Crewe Road, Edinburgh EH4 2XU and Department of Pathology, Royal Infirmary of Edinburgh, 51 Little France Crescent, Edinburgh EH16 4SA, UK 2

ABSTRACT. Following treatment for localized soft tissue sarcoma the risk of relapse is either locally or in the lungs. In Edinburgh patients are reviewed every 6 months with a chest X-ray (CXR). The radiation exposure over a 10 year follow up remains small, but it is unclear if all patients, irrespective of the initial grade of their primary tumour, require this. To determine the pick up rate of lung metastases by routine CXR over a 10 year period and to review the primary histology. Adult patients on routine follow up between 1994 and 2004 were identified and the notes of those with lung metastases reviewed. Data was collected on their initial histology, and date and method of diagnosis of lung metastases. 179 patients were under follow up. 24 (13%) developed lung metastases. For 2, notes were not found. 6 (27%) had metastases diagnosed by routine CXR, 9 (41%) had metastases diagnosed by non routine CXR and 7 (32%) had metastases diagnosed by CT. On review of histology none were grade 1, 4 (18%) were grade 2 and 18 (82%) were grade 3. 155 patients received. 6 monthly CXR for 10 years with no detection of lung metastases. Lung metastases occurred in a minority of patients (13%) and most (82%) occurred in patients with grade 3 tumours. No patients with grade 1 tumours developed lung metastases. Thus routine CXR may be appropriate on grade 3 tumours, but not on lower grade tumours where other risk factors are absent.

Soft tissue sarcoma in adulthood is relatively rare. Following treatment for localized disease, the risk of relapse is either locally or in the lungs. Prognostic indicators for recurrence are well documented [1–6] and include age at diagnosis, tumour depth, tumour size, histological type, histological grade, positive surgical margins and tumour site. All patients in our centre are followed up for 10 years regardless of initial pathology, with 6 monthly chest radiographs (CXR). This is justified because surgical intervention may be curative if lung metastases are diagnosed early. The effective radiation dose from a single CXR is quoted as between 0.02 mSv and 0.1 mSv, which is the equivalent to between 2.4 days and 10 days of background radiation. Thus whilst the overall radiation dose over 10 years from 6 monthly CXR remains small, it was unclear if CXR was a useful screening tool in this setting, and if it was indicated in all patients regardless on initial pathology. This retrospective audit was therefore performed to answer these questions.

Methods and materials Adult patients newly diagnosed with a localized primary soft tissue sarcoma and all patients on routine follow up between 1994 and 2004 were identified on the departmental database and the notes of those with lung metastases were reviewed. The British Journal of Radiology, October 2006

Received 9 December 2005 Revised 16 February 2006 Accepted 10 March 2006 DOI: 10.1259/bjr/69175634 ’ 2006 The British Institute of Radiology

Data were collected on the initial grade of tumour and the method of diagnosis of lung metastases. The pathology was reviewed where possible and graded according to Federation Nationale des Centre de Lutte Contre le Cancer (FNCLCC) classification.

Results A total of 179 patients were under follow up during this 10 year period. 24 (13%) developed lung metastases, and 22 sets of notes were retrieved. Two sets of notes were unavailable. Histology was reviewed for 21 patients and known from documentation for 1 patient. None of the 22 cases with pulmonary metastases had grade 1 tumours. 4 cases (18%) had grade 2 tumours and 18 (82%) had grade 3 tumours at initial diagnosis (Table 1). Six (27%) had their lung metastases diagnosed by routine screening CXR, 9 (42%) had their lung metastases diagnosed on a non-routine CXR, prompted by symptoms, and 7 (32%) had their metastases diagnosed by a CT scan, performed as part of re-staging for local recurrence or performed to investigate a decline in general health (Table 2). 155 patients received 6 monthly CXR for 10 years and remained relapse free. Total patient years at risk were 512.24, equating to 1 patient developing lung metastases for every 21 years of follow up. 799

H K Lord, D M Salter, R H M MacDougall and G Kerr Table 1. Number of patients developing lung metastases according to grade of primary Tumour grade

No. developing lung metastases

Grade 1 Grade 2 Grade 3

0 4 18

Discussion The small size of this cohort prevents meaningful statistical analysis, but the trend shown is in keeping with previously published data. Whilst this audit has not identified the metastatic potential for grade 1 tumours, there are data that this potential exists. In a paper by Le Doussal [1] the 5 year risk of metastases for grade 1 and 2 tumours together was 20%, although it is not clear how many were grade 1. In a paper by Le [2] looking at prognostic factors in head and neck sarcomas, which included 65 patients, the 5 year cause specific survival was 100% for grade 1 tumours. This is in concordance with a series by Willers et al [7] demonstrating similar outcomes for grade 1 tumours. However, the interplay of prognostic factors beyond tumour grade alone is relevant. Mandard et al [8] performed a multivariate analysis of 109 patients with soft tissue sarcoma. Four prognostic groups were identified depending on the presence or absence of necrosis, invasive tumour, adequate surgical margins and size of tumour less than or greater than 5 cm. Three patients with FNCLCC grade 1 tumours fell into a poor prognostic due to tumour size greater than 5 cm and/or inadequate surgical margins and 2 of these developed metastatic disease from which they died. A multivariate analysis performed by Heise [9] also demonstrated tumour site to be relevant, with retroperitoneal and mediastinal disease associated with poorer outcomes compared with disease in the extremities or head and neck. Thus it is clear that the prognosis of an individual patient is dependent on factors other than tumour grade. The usefulness of routine CXR to diagnose metastatic disease in soft tissue sarcoma has been less well documented. Kane [10] has summarized the data demonstrating a lack of cost effectiveness for the use of CT, and highlights how regular clinic visits and thorough clinical examination detect the majority of recurrent disease. Early pulmonary disease is frequently asymptomatic however and, if detected, can still be cured by surgical intervention. The low cumulative dose of radiation received from 6 monthly CXR makes this a safe, simple and appropriate first tool [11]. The optimal follow up for patients with soft tissue sarcoma remains unknown [12], but in this series routine CXR detected 27% of the pulmonary metastatic cases and thus should be recommended. Whilst it may be omitted in patients with grade 1 tumours, other prognostic

800

Table 2. Method of diagnosis of lung metastases Method of diagnosis

Number diagnosed

Routine CXR Non routine CXR CT

6 9 7

CXR, chest radiograph.

factors should also be borne in mind, and the decision based on the overall prognostic grouping, rather than grade alone.

References 1. Le Doussal V, Coindre JM, Leroux A, Hacene K, Terrier P, Bui NB, et al. Prognostic factors in patients with localised primary malignant fibrous histiocytoma: a multicenter study of 216 patients with multivariate analysis. Cancer 1996;77:1823–30. 2. Le QT, Fu KK, Kroll S, Fitts L, Massullo V, Ferrell L, et al. Prognostic factors in adult soft tissue sarcomas of the head and neck. Int J Radiat Oncol Biol Phys 1997;37:975–84. 3. Coindre JM, Terrier P, Guillou L, Le Doussal V, Collin F, Ranchere D, et al. Predictive value of grade for metastases development in the main histologic types of adult soft tissue sarcomas: a study of 1240 patients from the French Federation of Cancer Centers Sarcoma Group. Cancer 2001;91:1914–26. 4. Vraa S, Keller J, Nielson OS, Sneppen O, Jurik AJ, Jensen OM. Prognostic factors in soft tissue sarcomas: the Aarhus experience. Eur J Cancer 1998;34:1876–82. 5. Collin CF, Friedrich C, Godbold J, Hajdu S, Brennan MF. Prognostic factors for local recurrence and survival in patients with localised extremity soft tissue sarcoma. Semin Surg Oncol 1988;4:30–7. 6. Pezzi ME, Rawling MS, Esgro JJ, Pollock RE, Rohmsdahl MM. Prognostic factors in 227 patients with malignant fibrous histiocytoma. Cancer 1992;69:2098–103. 7. Willers H, Hug EB, Spiro IJ, Efird JT, Rosenberg AE, Wang CC. Adult soft tissue sarcomas of the head and neck treated by radiation and surgery or radiation alone: patterns of failure and prognostic factors. Int J Radiat Oncol Biol Phys 1995;33:585–93. 8. Mandard AM, Petiot JF, Marnay J, Mandard JC, Chasle J, de Ranieri E, et al. Prognostic factors in soft tissue sarcoma; a multivariate analysis of 109 cases. Cancer 1989;63:1437–51. 9. Heise HW, Myers MH, Russel WO, Suit HD, Enzinger FM, Edmonson JH, et al. Recurrence free survival time for surgically treated soft tissue sarcoma patients. Cancer 1986;57:172–7. 10. Kane JM. Surveillance strategies for patients following surgical resection of soft tissue sarcomas. Curr Opin Oncol 2004;16:238–321. 11. Chang AE, Schaner EG, Conkle DM, Flye MW, Doppman JL, Rosenberg SA. Evaluation of computed tomography in the detection of pulmonary metastases: a prospective study. Cancer 1979;43:913–6. 12. Sakata K, Johnson FE, Beitler AL, Kraybill WG, Virgo KS. Extremity soft tissue sarcoma patient follow up; tumour grade and size affect surveillance strategies after potentially curative surgery. Int J Oncol 2003;22:1335–43.



Multiple hyperechoic testicular lesions are a common finding on ultrasound in Cowden disease and represent lipomatosis of the testis 1

J WOODHOUSE,

MRCS

and 2M M FERGUSON,

FDSRCPS

1

Department of Urology, Christchurch Hospital, Christchurch and 2Department of Stomatology, University of Otago, Dunedin, New Zealand

ABSTRACT. Cowden disease (CD) is a genetic disease associated with multiple hamartomas and malignant neoplasms. During investigations for possible subnormal fertility, a series of eight males with CD underwent ultrasound scanning of their testes. Our findings detail the seven adult patients that were found to have multiple and bilateral testicular lesions on ultrasound. These lesions in Cowden’s patients represent a newly described testicular pathology – lipomatosis of the testis. Here we detail the radiological findings. Ultrasound findings showed multiple (estimated 40+) discreet lesions randomly scattered thoughout the parenchyma of each affected testis. The lesions themselves had heterogeneous echo texture, but all being hyperechoic and with a variable size from 1 mm to 6 mm. One patient underwent MR examination of the lesions in which a high T1 signal was seen, but no useful characteristic features were identifiable.

Cowden disease (CD; also known as Multiple Hamartoma Syndrome), is marked by increased cellular proliferation of ectodermal, mesodermal and endodermal tissues [1, 2]. These neoplasms can be benign or malignant. In particular, cancer of the breast, thyroid, endometrium and skin has been associated with CD. Diagnosis is based on operational criteria laid out by the ‘‘International Cowdens syndrome consortium’’ and supported by genetic testing. Mucocutaneous lesions comprise the predominant feature and are a principal way of suspecting the diagnosis. Phenotypic expression is variable and thus some cases probably remains undiagnosed. Despite it being a dominantly inheritable condition, the prevalence of the disease remains relatively low with more recent estimates at 1 per 250 000 [3]. A singular lesion within the parenchyma of a testis is not an uncommon finding with ultrasound scanning, but when found they are regarded as highly suspicious of malignancy. Multiple lesions of the testes are infrequent and usually represent microlithiasis. Multiple testicular lesions in the context of CD has been previously reported [4], but only in the context of a single patient. No evidence was provided to establish the histopathology in this case. Our findings are suggestive that these lesions are a very frequent occurrence in CD. This ultrasound finding is now known to represent a new histopathological entity called lipomatosis of the testis [5]. Here we detail the radiological aspects of this new condition. Address correspondence to: Professor Martin M Ferguson, Department of Stomatology & Health Sciences, University of Otago, P.O. Box 647 Dunedin, New Zealand.


Received 5 October 2005 Revised 22 March 2006 Accepted 30 March 2006 DOI: 10.1259/bjr/50628431 ’ 2006 The British Institute of Radiology

Methods All male patients with CD known to the authors were involved in this study, which was initially directed as a study into fertility issues and CD and subject to local ethics approval. The study was prospective and there were no exclusions made. All of the eight males had clinical features of CD and the diagnosis was confirmed by identification of the mutation on chromosome 10q.23. Ages ranged from 16 years to 58 years (mean 5 38 years). The patients were interviewed and examined with attention initially on issues surrounding possible low fertility. Structural assessment of the testes in all eight was performed with ultrasound. Scanning was performed with a range of machines in four separate institutions from both the private and public sectors. Scanners included Philips HDI 5000 using high frequency linear probes. Exact scanner models were unable to be determined for all scans performed. All sonographers held the diploma of Australian Society of Ultrasound in Medicine and Biology. All scans were reviewed retrospectively by consultant radiologists at Christchurch Public Hospital. In some cases, some scans were assessed in real time by those consultants. Following identification of anatomical testicular lesions, all patients had serum tumour marker assays performed (alpha fetoprotein, beta human chorionic gonadotropin, and lactate dehydrogenase) and were offered interval ultrasound. Two patients had had previous testicular ultrasound for comparison. Overall, five of the eight had repeated scans with intervals ranging from 5 months to 89 months for comparison. One patient underwent MRI (General Electric Sigma Horizon LX 1.5 Tesla, Wisconsin, USA) with T1, T1 fat saturated, T1 fat saturated postgadolinium enhancement (repetition time (TR) 600 ms, 801

J Woodhouse and M M Ferguson

echo time (TE) 14 ms, field of view (FOV) 10 cm, slice thickness 3 mm, gap 0.5 mm, phase 224, frequency 384, electronic train length 3; 3 inch surface coil) and T2 sequences (TR 3500 ms, TE 96 ms, FOV 10 cm, slice thickness 3 mm, gap 0.5 mm, phase 192, frequency 256, electronic train length 12; 3 inch surface coil), attempting to further characterize these lesions. Four of eight subjects elected to undergo open testicular biopsy under general anaesthetic due to lack of any definitive pathological correlation.

Results Of the cohort, all were normal on urogenital physical examinations with none of the lesions described being palpable. One patient previously had a testicular lump observed by ultrasound and discharged as benign, and this was still palpable and non tender. Another patient reported generalized low grade testicular tenderness. All were of normal sexual development. All patients had at least one ultrasound and five had repeat interval scans. Ultrasound showed that all had testes of normal size and contour, but the seven adults had multiple (estimates of 40+ per testes) hyperechogenic foci of approximately 1–6 mm size within the parenchyma of both testicles (Figure 1). Most lesions were small and non-shadowing. Only in one testis were acoustic shadows seen (Figure 2). The lesions were randomly scattered throughout the testis in all cases. The lesions did not demonstrate increased vascularity on Doppler imaging. There were no other focal parenchymal abnormalities seen. Five of seven had interval repeat ultrasound for comparison (range 5–89 months). Within the limits of ultrasound analysis, all lesions were reported as remaining stable. However, owing to the number of lesions in each testis it was not possible to categorically determine this.

Figure 2. Longitudinal ultrasound of testis in one case demonstrated that some of the lesions displayed acoustic shadowing. The lesions here are similarly hyperechoic. No hypoechoic regions or calcifications are seen.

Two patients had undergone previous testicular ultrasound performed for other reasons available for retrospective comparisons. One patient, 3 years previously, had four serial scans for a suspicious scrotal lump. These had shown a calcified nodule on the surface of the testicle, but with poorly defined regions of echogenicity within the parenchyma. The original interpretation was that these regions had represented microlithiasis. Films were reviewed in light of our findings and the opinion was that these lesions were not microlithiasis. Another patient had had two scans carried out 7 years previously. Identical small echogenic nonshadowing foci were present at that point. With both patients, as initial scans had been for other reasons, it was impossible to comment with certainty as to the stability of the lesions; although it did seem that these lesions had not progressed significantly. In an effort to further characterize the lesions noninvasively, one patient underwent MRI of the scrotum. MRI showed multiple (greater than 15) tiny focal highsignal areas within both testes, probably corresponding to the lesions seen on the ultrasound images, although these were very poorly visualized in all sequences used. Lesions were best seen on the T1 series (Figure 3). Insufficient detail was gained to be diagnostic.

Discussion

Figure 1. Longitudinal ultrasound view of the right testis demonstrates well defined multiple non-shadowing small round lesions that are hyperechoic. The lesions are of various sizes and are non-confluent. No hypoechoic regions or calcifications are seen.

802

CD is a rare disease and as such our cohort of eight patients represents a relatively large population of CD patients. The identification of the above described multiple bilateral testicular lesions on ultrasound and MR was initially documented by Lindsay [4] in a single CD patient, and has not been documented outside the context of CD. In this single case, the lesions were assumed to be hamartoma based on the lesions having a fat component on MRI, but no histopathological proof was offered for this diagnosis. Our series of CD patients were all impalpable incidentalomas found while The British Journal of Radiology, October 2006

Ultrasound of testicular lesions in Cowden disease can represent lipomatosis

Figure 3. Sagittal T1 weighted spin echo sequence (repetition time (TR) 600 ms, echo time (TE) 14 ms, 3 mm slice thickness, field of view (FOV) 10 cm) of the left testis. The scan demonstrates tiny hyperintense areas (white arrows) within the testis parenchyma consistent with fat.

investigating subfertility, probably explaining why they have not been extensively documented previously. Our group has established that these lesions represent lipomatosis of the testis in the four patients who underwent biopsy. Biopsy microscopic findings showed ‘‘interstitial lipomatosis consisting of nests of adipocytes within the testicular interstitium. No intratubular calcification or intratubular germ cell neoplasia was seen’’. There appeared to be no detrimental effect on spermatogenesis, nor on testicular biochemical function [5]. The diagnosis of malignancy was considered before diagnostic biopsy was performed. Serial scanning using ultrasound could not absolutely qualify the stability of the lesions as lesions were too numerous to enable the identification of individual lesions for comparative assessment on repeated scans. MRI unfortunately proved inadequate to be useful diagnostically. Hence biopsy was a diagnostic option that was offered to all these patients and testicular biopsies were performed on some of these patients (but not on all, once it was established that the lesions seen on ultrasound were not tumours). The appearances seen in our series represent benign testicular lipomatosis (hamartomatous growths). Ultrasound findings of primary testicular malignancies tend to be hypoechoic and homogeneous (seminoma) or are well defined with cystic elements and calcifications (teratoma). Lymphoma, leukaemia, and metastatic disease is uncommon and has variable appearance on ultrasound from focal hypoechoic lesions to a diffusely altered echogenicity [6, 7] – not similar to the findings of testicular lipomatosis seen in our cohort, nor the lesions described by Lindsay [4]. Benign testicular lesions are uncommon and are difficult to distinguish from malignancy on ultrasound. Because of this, they often result in orchidectomy. The high signal on MR T1 imaging supported a fatty component to the lesions and does lend itself to a diagnosis of hamartomas (as suggested by Lindsay) or lipomas as is the diagnosis. Testicular lipomas are


reported, although they are very rare and have only been described only as singular lesions [8, 9]. Microlithiasis is the most commonly documented testicular pathology occurring as a multiple entity, and consists of multiple randomly scattered punctuate lesions in the testicular parenchyma (with variable acoustic shadowing) of approximately 1–3 mm in size representing calcified deposits in the lumen of seminiferous tubules. As they represent calcifications, they have a very high echogenicity and are thus distinctly different to the lesions we find in the Cowden’s testes. Testicular lipomatosis is a novel entity. We have detailed how the ultrasound appearances differ distinctively from other pathologies. These lesions are not described outside the context of CD and thus are apparently pathognomic for CD and could be used as a major diagnostic criterion for CD in adult males. A finding on ultrasound of these lesions should thus alert the radiologist to the diagnosis of CD. In our series, 7 of 8 of males with CD had these lesions. Phenotypic expression in CD often is not manifest until aged 20–30 years [10] and it remains to be seen if the eighth male – currently with no testicular lesions and aged 16 years – will eventually develop these lesions. The long term behaviour of these lesions is unknown to date.

Acknowledgments The Authors would like to acknowledge the kind help of Hamish Fraser, Andrew Lang, Diane Leighton and Tim Buckingham for their help and advice regarding this work. We would also like to acknowledge the staff of Christchurch Public Hospital Radiology department and the Southern Cross Radiology department where the majority of scans were performed.

References 1. Lloyd KM, Dennis M. Cowden’s disease. A possible new symptom complex with multiple system involvement. Ann Intern Med 1963;58:136–42. 2. Weary PE, Gorlin RJ, Gentry WC Jr, Comer JE, Greer KE. Multiple hamartoma syndrome (Cowden’s disease). Arch Dermatol 1972;106:682–90. 3. Nelen MR, et al. Novel PTEN mutations in patients with Cowden disease: absence of clear genotype-phenotype correlations. Eur J Hum Genet 1999;7:267–73. 4. Lindsay C, Boardman L, Farrell M. Testicular hamartomas in Cowden disease. J Clin Ultrasound 2003;31:481–3. 5. Woodhouse JB, Delahunt B, English SF, Fraser HH, Ferguson MM. Testicular lipomatosis in Cowden’s syndrome. Mod Pathol 2005;18:1151–6. 6. Hamm B. Differential diagnosis of scrotal masses by ultrasound. Eur Radiol 1997;7:668–79. 7. Howlett DC, Marchbank ND, Sallomi DF. Pictorial review. Ultrasound of the testis. Clin Radiol 2000;55:595–601. 8. Harper M, Arya M, Peters JL, Buckingham S, Freeman A, O’Donoghue EP. Intratesticular lipoma. Scand J Urol Nephrol 2002;36:223–4. 9. Honore LH. Fatty metaplasia in a postpubertal undescended testis: a case report. J Urol 1979;122:841–2. 10. Nelen MR, et al. Localization of the gene for Cowden disease to chromosome 10q22–23. Nat Genet 1996;13:114–6.

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Duplex ultrasound of the superior mesenteric artery in chronic pancreatitis 1

M HORNUM,

MD,

1

S LARSEN,

MD, PhD,

3

O OLSEN,

MD, PhD

and 2J F PEDERSEN,

MD, PhD

Departments of 1Gastroenterology and 2Radiology, Glostrup Hospital, Copenhagen and 3 Department of Surgical Gastroenterology, Amager Hospital, University of Copenhagen, Copenhagen, Denmark ABSTRACT. Blood flow in the superior mesenteric artery (SMA) increases after a meal due to a vasoactive effect of the decomposed food. In exocrine pancreatic insufficiency, the digestion of food is compromised. We used duplex ultrasound to test the hypothesis that blood flow in the SMA after a meal increases less in patients with pancreatic insufficiency than in control persons. We studied 16 patients with chronic pancreatitis, eight of them with exocrine insufficiency, and eight healthy volunteers. The resistive index (RI) in the SMA was determined before and after a liquid meal. The RI reflects the downstream circulatory resistance, giving a precise description of mesenteric hyperaemia. Both groups of patients with chronic pancreatitis unexpectedly had lower fasting RI than controls, 0.818 and 0.815 vs 0.851, p50.028 and p50.0030, respectively. Postprandialy there was significantly less decrease in RI (less increase in flow) in patients with exocrine insufficiency than in controls, 0.055 vs 0.099, p50.0047. There was a significant trend for a less pronounced postprandial decrease in RI with more impaired pancreatic function (p50.0036). Our study thus demonstrates a reduced postprandial increase in SMA flow in patients with exocrine pancreatic insufficiency, and suggests an increased fasting SMA flow in chronic pancreatitis. Further studies are needed to evaluate the possible role of the test-meal-induced shift in RI in the SMA and of a lower-than-normal fasting RI in the diagnosis and monitoring of chronic pancreatitis.

At ultrasound scanning combined with the pulsed Doppler technique we can non-invasively monitor alterations in blood flow velocity. Applied to the superior mesenteric artery (SMA), the method confirms that blood flow in this vessel increases in response to a meal [1–3]. The vasoactive components seem to be influenced by the digestive products of the diet, so that the effect of a meal on splanchnic blood flow may depend on the intraluminal digestion [1]. Patients with end stage chronic pancreatitis are characterized by maldigestion due to exocrine pancreatic insufficiency, but it is not known whether the reduced intraluminal food digestion is mirrored in the mesenteric blood flow. It was our hypothesis that these patients would show less postprandial increase in mesenteric blood flow than control persons, so that the postprandial shift in Doppler velocity pattern might be used in the diagnosis and monitoring of chronic pancreatitis. In this study we examine the velocity pattern in the SMA before and after a liquid test meal in patients with chronic pancreatitis with and without preserved exocrine function, and in healthy controls.

Methods and patients We studied 16 patients with chronic pancreatitis, eight with minimal or no exocrine pancreatic function and eight with reduced exocrine function, and eight 804


healthy volunteers (Table 1). The aetiology of chronic pancreatitis was alcohol in 14 patients and unknown in two. All patients had moderate or marked imaging findings according to the Cambridge classification [4] plus reduced exocrine pancreatic function judged from the intraduodenal meal-stimulated lipase concentration (Lundh test), or steatorrhoea. The eight patients with exocrine insufficiency were characterized by a Lundh test result below 10% of the normal lower limit [5], or excretion of more than 7 g of fat per day in faeces [6] (plus morphological changes). Four of them had diabetes but no signs of peripheral vascular disease, and seven had enzyme substitution. The patients with residual pancreatic exocrine function had no steatorrhoea and meal stimulated intraduodenal lipase concentration far over 10% of lower normal limit (plus morphological changes) and had no other known gastrointestinal disorder, and no diabetes or peripheral vascular disease. The controls had no known gastrointestinal disease and were not taking any kind of medication. All participants gave written informed consent to take part in the study, which was approved by the Regional Research Ethics Committee (KA 0129). The ultrasound examinations were performed with a Philips HDI 5000 Sono CT unit (Philips Medical Systems, Bothell, WA). The SMA was examined in its long axis in the sagittal plane. The sampling cursor was placed within the diameter of the vessel, 2–3 cm distal to its origin, and the angle between the ultrasound beam and The British Journal of Radiology, October 2006

Mesenteric blood flow and pancreatic insufficiency Table 1. Patients and control persons Chronic pancreatitis

No. of subjects Females/males Median age, years (range) Median body mass index kg m2 (range)

Normal control persons

Exocrine insufficiency

No exocrine insufficiency

8 2/6 55 (40–65)

8 4/4 51 (36–73)

8 3/5 44 (31–69)

21 (17–25)

24 (19–28)

23 (19–27)

the SMA was kept smaller than 60 ˚. The resistive index (RI) was calculated in accordance with the formula: peak-systolic velocity minus end-diastolic velocity divided by the peak-systolic velocity [7]. Each RI value is the average of three measurements. To monitor gastric emptying the antrum was localized on a sagittal image in front of the superior mesenteric vein [8]. The image was frozen and the antral area was measured by means of the built-in calliper system. After the meal the antral area was determined at 5–10 min intervals until the area approached baseline, and the time from the meal until the area had decreased to 150% of baseline was used as surrogate expression of gastric emptying time [8]. Patients and healthy volunteers were examined in random order. Any pancreatic enzyme substitution was discontinued for 72 h before the study. The examination commenced after an overnight fast, with the subjects resting in the supine position for 45 min. RI in the SMA was recorded at least twice during the last 30 min before the meal and the baseline antral area was determined. The subjects then ingested a test meal consisting of 74.8 g NAN 1 (Nestle´ Danmark A/S), containing 5.9% fat, 2.3% protein, and 11.6% carbohydrate, in 300 ml of water, with a total energy load of 1598 kJ, and SMA flow characteristics were recorded five times at 15 min intervals. Antral area was monitored as described. In most subjects, gastric emptying was not complete after 80 min. In these situations antral area was monitored for an additional 20 min. For the data analysis, we calculated the difference between baseline RI and the mean of the five post-meal RI determinations in each participant. This difference thus expresses the integrated RI response to the test meal. Results were analysed for statistical significance

using the Mann-Whitney two-sample rank sum test for unpaired data and Wilcoxon test for paired data. The Jonckheere-Terpstra test for ordered alternatives [9] was used to test the hypothesis that the postprandial shift in RI would be smaller with poorer exocrine function. Results are expressed as medians and ranges, and p,0.05 was considered statistically significant.

Results The fasting RI was significantly lower in patients with exocrine insufficiency and in patients with reduced exocrine function, median values being 0.82 in both groups as compared with 0.85 in healthy volunteers, p50.028 and p50.0030, respectively (Table 2, Figure 1). After the test meal, the RI decreased less in patients without and with residual exocrine function, 0.055 and 0.072, respectively, compared with controls, 0.099 (Table 2, Figures 2 and 3). The difference between patients with exocrine insufficiency and controls was statistically significant, p50.0047, indicating less postprandial increase in mesenteric blood flow in exocrine insufficiency, whereas there was no significant difference between patients with residual exocrine function and the control group. There was a significant trend for less postprandial decrease in RI (less decrease in downstream circulatory resistance) from controls through patients with chronic pancreatitis without clinical insufficiency to patients with manifest pancreatic insufficiency (p50.0036, Figure 3). There was no difference in the mean antral emptying time between the three groups (Table 2).

Discussion We used RI to characterize downstream circulatory resistance in the SMA. It is calculated from two Doppler shift frequencies measured at the same image, and is independent of the angle of insonation. The RI is shown to increase linearly with the peripheral resistance at a constant pressure [7, 10]. It therefore makes sense that RI in the SMA has been reported significantly lower in active than in inactive ulcerative colitis [11] and Crohn’s disease [12]. Median fasting RI in our control persons was 0.85 which is in agreement with results from other studies [11, 12]. Unexpectedly, both patients with chronic pancreatitis and patients with pancreatic insufficiency

Table 2. Median resistive index (RI) in the superior mesenteric artery before and after standard meal and antral emptying time after meal in eight patients with chronic pancreatitis and exocrine insufficiency, eight patients with chronic pancreatitis and preserved exocrine function and in eight healthy control persons Chronic pancreatitis

Fasting RI, median (range) Mean RI 0–80 min after meal, median (range) Decrease in RI after meal, median (range) Antral emptying time, median (range, min.)

Normal control persons

Exocrine insufficiency

No exocrine insufficiency

0.818 (0.753–0.850) 0.753 (0.691–0.869)

0.815 (0.757–0.838) 0.712 (0.695–0.761)

0.851 (0.821–0.893) 0.745 (0.676–0.801)

0.055 (–0.023–0.093)

0.072 (0.057–0.110)

0.099 (0.078–0.146)

77 (40 to .100)

73 (59 to .100)

94 (75 to .100)


805

M Hornum, S Larsen, O Olsen and J F Pedersen

Figure 1. Fasting resistance index (RI) in the superior

Figure 3. Difference between fasting and postprandial

mesenteric artery in eight patients with chronic pancreatitis and exocrine insufficiency, eight patients with chronic pancreatitis without manifest insufficiency and in eight healthy control persons.

resistance index (RI) in the superior mesenteric artery in eight patients with chronic pancreatitis and exocrine insufficiency, eight patients with chronic pancreatitis without manifest insufficiency and in eight healthy control persons.

had significantly lower fasting RI in the SMA than controls, 0.82 in both groups. This suggests a more dilated vascular bed peripherally in the SMA territory in the fasting state in patients with chronic pancreatitis. To our knowledge this has not been previously reported, and further studies are needed to confirm this incidental observation. It could be speculated that in chronic pancreatitis the mechanisms that regulate the flow in the SMA are adapted to a weaker stimulus from the digested food components.

Figure 2. Resistance index (RI) in the superior mesenteric artery before (baseline) and 0–80 min after test meal in eight patients with chronic pancreatitis and exocrine insufficiency ( – ), eight patients with chronic pancreatitis without manifest insufficiency (#–#), and in eight healthy control persons (m–m).

NN

806

After the meal, RI decreased less (downstream resistance decreased less) in patients with exocrine insufficiency compared with patients with preserved exocrine function and with healthy controls (Figure 2). This lower effect of a meal on the blood flow in SMA in patients with exocrine insufficiency has not been demonstrated before. The test meal contained protein, fat and carbohydrate. All three components increase the blood flow in SMA after isocaloric and iso-osmotic loads into the duodenum [13]. In a recent study, instillation into the duodenum of free fatty acids (product of hydrolysis of triglycerides) increased SMA blood flow in normal subjects [1] to the same degree as we observed in the present study using a mixed test meal. Our results indicate that digestive products are more active in modulating SMA blood flow than undigested food. This could be due to increased release of vasoactive gastrointestinal hormones such as glucagon-like peptide-2 (GLP-2) during intestinal absorption of digestive products [14]. Since fat in the duodenum provokes gallbladder emptying [15], another explanation for our findings could be that the bile rather than the food causes the vascular effects. However, intravenous cholecystokinin in doses known to induce gallbladder contraction causes no flow response in SMA [16], so luminal bile does not seem to be responsible for the increased flow. Also, it has been reported that the postprandial gastric motility in chronic pancreatitis is altered [17, 18] so that a difference in gastric emptying of the test meal could influence our results. We therefore monitored antral emptying and observed no significant difference in antral emptying rate between patients and controls. Any practical consequences of our findings remain to be defined. The modestly lower fasting RI in chronic pancreatitis is not impressive, the difference between medians in patients and controls being 0.03, but the majority of patients (11 of 16) actually had RI levels below the lowest level in our control persons (Figure 1). The British Journal of Radiology, October 2006

Mesenteric blood flow and pancreatic insufficiency

If other studies confirm this observation, it might be considered worthwhile to determine RI in the SMA routinely in all fasting upper abdominal studies, or at least in studies without an obvious diagnosis. An unusually low RI would suggest undetected abdominal pathology like inflammatory bowel disease or chronic pancreatitis. The study confirms our hypothesis of a reduced postprandial increase in SMA flow in patients with exocrine pancreatic insufficiency. Further studies are needed to evaluate whether the meal induced shift in RI in the SMA can be used in the diagnosis and monitoring of chronic pancreatitis. Such a test would have the virtue of being non-invasive, in contrast to the duodenal-tubebased measurement of meal-induced release of lipase.

Acknowledgments This study was supported by grants from the Danish Hospital Foundation for medical research. Region of Copenhagen, The Faroe Islands and Greenland (15/03). Special thanks to Tove Laursen for technical support.

References 1. Andersen TC, Pedersen JF, Nordentoft T, Olsen O. Fat and mesenteric blood flow. Scand J Gastroenterol 1999;34:894–7. 2. Dauzat M, Lafortune M, Patriquin H, Pomier-Layrargues G. Meal induced changes in hepatic and splanchnic circulation: a noninvasive Doppler study in normal humans. Eur J Appl Physiol Occup Physiol 1994;68:373–80. 3. Qamar MI, Read AE. Effects of ingestion of carbohydrate, fat, protein, and water on the mesenteric blood flow in man. Scand J Gastroenterol 1988;23:26–30. 4. Sarner M, Cotton PB. Classification of pancreatitis. Gut 1984;25:756–9. 5. Worning H, Mullertz S. pH and pancreatic enzymes in the human duodenum during digestion of a standard meal. Scand J Gastroenterol 1966;1:268–83. 6. Van de Kamer JH, ten Bokkel Huinink H, Weyers HA. Rapid method for determination of fat in feces. J Biol Chem 1949;177:347–55.


7. Spencer JA, Giussani DA, Moore PJ, Hanson MA. In vitro validation of Doppler indices using blood and water. J Ultrasound Med 1991;10:305–8. 8. Pedersen JF. A modified sonographic technique for assessment of gastric emptying of liquid. Acta Radiol 2003;44:340–2. 9. Siegel S, Castellan NJ. Nonparametric statistics for the behavioural sciences. 2nd edn. Boston: McGraw-Hill, 1988:216–22. 10. Maulik D, Arbeille P, Kadado T. Hemodynamic foundation of umbilical arterial Doppler waveform analysis. Biol Neonate 1992;62:280–9. 11. Maconi G, Imbesi V, Porro GB. Doppler ultrasound measurement of intestinal blood flow in inflammatory bowel disease. Scand J Gastroenterol 1996;31:590–3. 12. Yekeler E, Danalioglu A, Movasseghi B, Yilmaz S, Karaca C, Kaymakoglu S, et al. Crohn disease activity evaluated by Doppler ultrasonography of the superior mesenteric artery and the affected small-bowel segments. J Ultrasound Med 2005;24:59–65. 13. Sieber C, Beglinger C, Jager K, Stalder GA. Intestinal phase of superior mesenteric artery blood flow in man. Gut 1992;33:497–501. 14. Guan X, Stoll B, Lu X, Tappenden KA, Holst JJ, Hartmann B, et al. GLP-2-mediated up-regulation of intestinal blood flow and glucose uptake is nitric oxide-dependent in TPNfed piglets 1. Gastroenterology 2003;125:136–47. 15. Olsen O, Schaffalitzky de Muckadell OB, Cantor P, Erlanson-Albertsson C, Hansen CP, Worning H. Effect of trypsin on the hormonal regulation of the fat-stimulated human exocrine pancreas. Scand J Gastroenterol 1988;23:875–81. 16. Sieber C, Beglinger C, Jaeger K, Hildebrand P, Stalder GA. Regulation of postprandial mesenteric blood flow in humans: evidence for a cholinergic nervous reflex. Gut 1991;32:361–6. 17. Layer P, Ohe MR, Holst JJ, Jansen JBMJ, Grandt D, Holtman G, et al. Altered postprandial motility in chronic pancreatitis: role of malabsorption. Gastroenterology 1997;112:1624–34. 18. Vu MK, Vecht J, Eddes EH, Biemond I, Lamers CBHW, Masclee AAM. Antroduodenal motility in chronic pancreatitis: are abnormalities related to exocrine insufficiency? Am J Physiol Gastrointest Liver Physiol 2000;278:G458–66.

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Accuracy and precision of an external-marker tracking-system for radiotherapy treatments 1

E M DONOVAN, 3 R WILKS, PhD

PhD,

2

P BRABANTS, 1P M EVANS,

DPhil,

1

J R N SYMONDS-TAYLER,

MSc

and

1

Joint Department of Physics, Institute of Cancer Research and Royal Marsden Foundation Trust, Downs Road, Sutton, Surrey SM2 5PT, 2QADOS Ltd, 5 Lakeside Business Park, Swan Lane, Sandhurst, Berkshire GU47 9DN and 3Osiris Systems Ltd, Burraton Bungalow, Broadclyst, Exeter, Devon EX5 3DB, UK ABSTRACT. The purpose of this work was to determine the accuracy and precision of a real-time motion-tracking system (Osiris+) for the monitoring of external markers used on patients receiving radiotherapy treatments. Random and systematic errors in the system were evaluated for linear (1D), circular (2D) and elliptical (3D) continuous motions, and for a set of static positions offset from an origin. A Wellhofer beam data measurement system and a computer controlled platform (which could be programmed to give motion in 3D) were used to move a hemi-spherical test object. The test object had four markers of the type used on patients. Three markers were aligned in the central plane and a fourth was positioned out of plane. Errors were expressed as deviations from the planned positions at the sampled time points. The marked points on the test object were tracked for the linear motion case with a variation from the true position of less than ¡1 mm, except for two extreme situations. The variation was within ¡2 mm when the lights were dimmed and when the amplitude of the movement was ¡5.0 cm. The 2D circular motion was tracked with a standard deviation of 1 mm or less over four cycles. The sampling rates of the system were found to be 0.3– 0.4 s when it was monitoring actively and 1.5–1.6 s otherwise. The recorded Osiris+ measurements of known static positions were within ¡1 mm of the value from the computer controlled platform moving the test object. The elliptical motions in 3D were tracked to ¡1 mm in two directions (Y,Z), and generally to within ¡2 mm for the third direction (X); however, specific marked points could display an error of up to 5 mm at certain positions in X. The overall displacement error for the 3D motion was ¡1 mm with a standard deviation of 2.5 mm. The system performance is satisfactory for use in tracking external marker motion during radiotherapy treatments.

Portal imaging protocols for treatment verification and data analysis have been derived, and used to determine random and systematic errors in patient set up, based on bony anatomy and internal markers for a range of treatment sites, for example, breast [1, 2], prostate [3–5] and lung [6]. The data thus measured has been used to determine appropriate margins around organs at risk for radiotherapy planning [7–9]. This type of protocol is wellestablished and interest has moved towards real time monitoring of both external and internal organ movement. Information about external motion throughout the duration of the treatment may be used in combination with electronic imaging of internal structures to improve further the accuracy and precision of radiotherapy treatment. This may be of particular value in situations where the radiation delivery is gated to a physiological function such as breathing [10–12], or where there is potential for the treated organ to be mobile, an example being the breast. There is considerable interest in respiratory gating for thoracic tumours treated with radiotherapy. The clinical problem is significant as the large margins necessary to 808

Received 22 July 2005 Revised 22 December 2005 Accepted 24 January 2006 DOI: 10.1259/bjr/24917728 ’ 2006 The British Institute of Radiology

create a planning target volume (typically 1.5 cm) limit the prescribed dose to the tumour so that normal tissue tolerances are not exceeded. If these margins could be reduced by some control of the tumour motion, it might be possible to escalate the dose to the tumour [13, 14]. Systems which have been used to investigate thoracic tumour motion and respiratory gating include those based on external devices [15] and internal markers [16]. Some authors have evaluated the correlation of the signals from devices placed on the abdomen and thorax with tumour motion monitored with kilovoltage imaging [17, 18]. Others have explored the variation of internal markers in the tumour with imaging [19]. It appears from these studies that it is difficult to generalize on any relationship between external and internal motion and that this needs to be determined on an individual patient basis [16]. One component of any system required to investigate this is a means of tracking external markers, either to correlate to internal motion, or to provide a signal for gating of the linac. We have investigated the tracking capability of a system called Osiris+ (QADOS, Sandhurst, Berkshire) The British Journal of Radiology, October 2006

Assessment of radiotherapy motion tracking system

[20]. The Osiris+ equipment is a simple camera-based system which enables the user to check patient contours at the time of treatment and to monitor skin marks. Patients are positioned for treatment using skin tattoos, often highlighted by external, cross-shaped pen markers on their skin. The system has a function to monitor such marks. This function can be used to ensure that the patient has not moved during the radiation delivery and has the potential to provide a signal to be used in gating. It does not require any monitoring equipment to be fixed to the patient. It is one of a number of systems which provide a motion tracking function [15, 21, 22]. The measurement errors of any tracking system must be evaluated prior to any potential use in monitoring movement. For any tracking system to fulfil its purpose, the system measurement errors must be much lower than the movement to be measured. The purpose of this work was to determine the accuracy and precision of the Osiris+ system when tracking markers.

Methods and materials Osiris+ system The Osiris+ equipment may be installed in radiotherapy simulator and treatment rooms without modifications to the simulator or treatment equipment, or direct patient contact [20]. It consists of a set of wall mounted cameras which are used with the in-room alignment lasers to acquire patient contours at multiple levels along the patient, or alone, to track markers on the patient’s skin. When the system is used for acquiring external outlines, images of the patient in the set-up position are captured and the patient outline and the reference marks may be generated by means of an automatic outlining function. The Osiris+ system also has a real time movement monitoring function which allows skin marks to be tracked during irradiation using stereoscopy – the mathematical combination of two images taken using different camera positions to create three-dimensional information. The real time movement monitoring uses the marks present on the patient’s skin, or immobilization shell, and does not require any external markers to be adhered to the patient. Our evaluation of the system concentrated on this monitoring function. The system was installed in a linac treatment room. It had four cameras mounted on the walls, all along a line of sight to the linac isocentre. The user indicates the points to be tracked on still images of the target object. Each point must be seen by two cameras in order to be tracked in 3D. After the points have been selected, the system displays real time images from the camera acquired at a slow frame rate. The active motion tracking is initiated manually and the system acquires images at a faster frame rate for the duration of the tracking period. This is referred to as the ‘‘linac on’’ state in the software although the acquisition of motion data does not currently gate to the linac pulse production. The tracking time is determined by the user. A point tracking algorithm determines the coordinates of the selected points in space and time and stores these in a simple The British Journal of Radiology, October 2006

spreadsheet format. The image acquisition rates and the accuracy and precision of the system were all evaluated in the experiments described.

Description of point tracking algorithm A pattern matching function has been implemented in the Osiris+ system to enable the tracking of points. The function is a grey level and colour pattern matching library. The system is trained on a reference pattern and afterwards locates its occurrences in other images. The library works by superimposing the pattern over the image and comparing them by computing a (normalized) correlation score, i.e. measuring discrepancies between the pattern and the target image. For each point selected, a search region of interest 100 pixels square is defined, and a central region of interest of 15 pixels square is stored as a search pattern. When locking on points, a match is searched for only in the search region. This typically allows about ¡40 mm of search area – depending upon the calibration values for that camera. (If the whole image were searched for a match, the process would take too long, especially if several points are being tracked). If all positions had to be tried for a match, this would lead to an unacceptable running time. To alleviate this, a coarse-to-fine approach is used. This means that several search stages called reductions are performed. At the coarsest reduction, an approximate location is found quickly. Then the location is improved, using the next reductions, and working in a close neighbourhood. This arrangement drastically reduces the number of positions to be tried. At the final stage, additional processing can be done to achieve sub-pixel accuracy. Points are not ‘‘lost’’ if they are temporarily obscured, but the data associated with those points when obscured or out of range will be set to a null. Point matching enables a point selected from the image obtained from a camera to be ‘‘found’’ whenever the image is updated. The location of each point found is indicated by its x, y screen coordinates. For this location to be translated from ‘‘screen coordinates’’ (pixels) to a position in space (world coordinates – mm) it is necessary for the same point to be seen by at least two cameras. Osiris+ uses an x, y, z world coordinate system where x is lateral, y is vertical and z is longitudinal. The centre of each camera’s field of view is the isocentre of the machine. At the isocentre ‘‘z’’ is always considered to be 0, i.e. ‘‘Z0’’. ‘‘X ’’ and ‘‘Y’’ coordinates are considered to be in this Z0 plane. The calibration of the Osiris+ system results in the position of each camera being known relative to the isocentre and so, any x, y screen coordinate of a point lying in the Z0 plane seen by any one camera can mathematically be translated to its true world coordinates, X, Y mm in the Z0 plane only. However, if the same point is seen by another camera then it is possible to translate the screen coordinates for that point (e.g. x1, y1 for camera 1 and x2, y2 for camera 2) to world coordinates of X,Y and particularly, Z mm. The solution is to solve for vector or ‘‘skew’’ lines, which are lines in space that are not parallel. If the same points can be seen by camera 1 and camera 2, then for 809

E M Donovan, P Brabants, P M Evans et al

each point selected from camera 1 it is necessary to correlate that point with one of the points selected from camera 2. This is done by generating a skew line which passes through x1, y1 between Z50 and Z5100 (camera 1). For each point from camera 2, skew lines are similarly generated and the one which passes closest to that from camera 1 (,5 mm ) is considered to be derived from the same point, say, x2, y2 on camera 2. These two points are then correlated. The process is repeated for all points selected from camera 1 against all points selected from camera 2, ignoring those already correlated. It is then possible using direction cosines to determine the X, Y and Z coordinates for each pair of correlated points. The correlation process described occurs only once when the points to be tracked are defined. After that, every time a fresh set of images is acquired, a pattern match is attempted for each point from each camera. This matched point will have screen coordinates of x1’, y1’ etc., which may or may not be the same as the last time around. The actual position/movement of the point can then be determined in world coordinates. The repetition rate is dependant upon how many points are being tracked and how long it takes to find a match.

Determination of accuracy and precision A rigid hemi-spherical test object was used to evaluate the Osiris+ system in all of the experiments. The test object was marked, using a permanent ink pen, with three markers in the central plane and one out of plane marker, as shown in Figure 1. The markers were cross shaped of 1.5 cm length and 0.2 cm width. The Osiris+ system was used to track continuously throughout each movement set. Each mark on the test object was seen by at least two cameras and was manually indicated by the user on a static camera image. For the marker on the top of the test object, there was a systematic discrepancy in the position indicated by the user on the two camera views of 2 mm. The system recorded and tracked all marked positions separately, hence the data are recorded for five points where points 3 and 4 refer to the same marker. Once the positions are marked by the user, the system begins monitoring at a slow frame rate. Active monitoring at a faster frame rate is initiated by the user (this corresponds to the ‘‘linac on’’ state). In all experiments the active mode of Osiris+ was activated prior to the movements driven by two motorized systems. The output from the Osiris+ system uses a binary tag to indicate whether passive or active monitoring is being recorded. The active monitoring state is referred to as ‘‘linac on’’ in the system software. It is initiated manually rather than being gated to the linac pulse production. All experiments were performed without radiation. The two motorized systems used were a clinical beam data measuring system and an experimental programmable platform. The advantage of the clinical system was that its measurement integrity had been well validated. However, it was difficult to use this system to create more complex motions in 3D. The programmable platform allowed the input of any function and hence provided additional flexibility. 810

Figure 1. Photograph of the test object with markers. The photograph is taken from above the test object. The height of the object is 8.0 cm. There are three markers in a plane and one out-of-plane marker. One camera pair viewed the left-side, top and out-of-plane marker; a second camera pair viewed the top and right-side markers.

Using Wellhofer beam measurement system A Wellhofer beam measurement system (Wellhofer Scanditronix, Schwartzenbruck, Nurnberg, Germany) was used for the first set of experiments. This system was in routine clinical use and the positional accuracy of the detector carriage had been verified previously to be 0.1 mm. The test object was fitted into the detector holder of the system and hence could be moved around in 3D space. It was aligned with the sagittal and lateral room set up lasers using the three central plane markers, and that position was defined as the origin. The Wellhofer system was used to drive the test object in each of the three cardinal directions from the origin at the amplitudes and speeds given in Table 1 in a linear movement that was repeated five times. The Linear 1 conditions were taken as the reference conditions and one parameter was changed to give each of the linear movement conditions labelled 2 to 7. The coordinate system of the Wellhofer equipment was related to the Osiris+ and Elekta linac treatment room coordinate system in the following way: Z5Gun2Target (GT) (longitudinal) direction, Y5vertical, X5transverse (lateral) direction. The Osiris+ system is designed to be used with maximum room illumination; one of the tests was carried out with dimmed lighting to test the ability of the system to track marks under more extreme conditions. A further experiment moved the phantom in the Z–Y (GT/vertical) plane in an approximation to a circle. The coordinates of movement are shown in Figure 2 and the The British Journal of Radiology, October 2006

Assessment of radiotherapy motion tracking system Table 1. Summary of the experimental parameters for the linear motion using the Wellhofer beam data measurement system Experiment number

Direction

Speed (mm s21)

Amplitude (cm)

Number of points tracked

Room light conditions

Linear Linear Linear Linear Linear Linear Linear

Z (GT) Z (GT) Z (GT) Z (GT) Z (GT) X (Lateral) Y (Vertical)

10.5 10.5 10.5 0.87 10.5 10.5 10.5

¡1.0 ¡1.0 ¡1.0 ¡1.0 ¡5.0 ¡1.0 ¡1.0

5 5 3 5 5 5 5

On Off On On On On On

1 2 3 4 5 6 7

GT, gun–target axis of the linear accelerator.

Wellhofer system was programmed to drive between the positions in a straight line at a speed of 10.5 mm s21 with no pause between sections. This sequence of movements was repeated four times.

Using programmable computer controlled platform with motion in X,Y,Z The computer-controlled platform (Time and Precision, Basingstoke, UK) has three orthogonal axes driven by leadscrews and stepper motors. Each axis has an incremental encoder which allows the position to be known at a resolution of 0.5 mm, even if a motor skips steps. The motors are microstepped to give a drive resolution of 0.5 mm. A versatile three-axis motion controller (Model DMC-2130; Galil Motion Control, Rocklin, CA) generates the step and direction signals to the three motor drives, according to the specified trajectory. The controller also keeps track of the actual position of the axes via the encoders, but the steppers are run open-loop. The difference between commanded position and actual position is less than 0.02 mm.

Figure 2. Coordinates of circular motion in Y–Z plane (Vertical: gun–target (GT) plane) using the Wellhofer beam data system. The British Journal of Radiology, October 2006

The test object was fixed on the platform and aligned to the room lasers using the three markers in the central plane to define the origin position. The platform was used to drive the test object to 3 points in X,Y,Z from this origin: ¡1 cm, ¡2 cm, ¡4 cm. After each set of movements in the three cardinal directions, the test object was returned to the origin. This experiment was designed to assess the absolute positional accuracy of the system. The platform was used to drive the test object in four elliptical trajectories in three-dimensional space. To generate the ellipses, the controller was run in its ‘‘electronic cam’’ mode, in which the trajectories of the three axes are cyclically synchronised to a virtual cam axis. Commands were sent to the controller from a laptop via a serial link. These trajectories corresponded to an ellipse of eccentricity 0.6 tilted at 45 ˚ in both horizontal and vertical planes. The details are listed in Table 2.

Results Using the Wellhofer beam measurement system Linear motion: positional data These data were analysed by evaluating the difference between the true position and time, as given by the speed and period of the Wellhofer equipment, with that recorded by the Osiris+ system. Figure 3 provides a summary of the standard deviation of the positional data difference for the seven experiments. For each experiment the data are separated into the X,Y,Z components of movement and are presented as the average over all of the tracked points and over all five cycles. Each of these experiments involved motion in one of the cardinal directions only. The direction of motion of the test object is indicated by the black bar on the chart. Data were collected for all three directions. The largest standard deviations are found for experiment 2, when the system was tracking the points under dimmed lighting, and for experiment 5 where the motion had a large amplitude of ¡5.0 cm. Neither of these two conditions resulted in standard deviations of greater than 2 mm; the other experiments showed that the system could track the points with standard deviations of less than 1 mm. The conditions in which the system tracked the motion with the greatest precision were when the motion was slow (0.87 mm s21) and here the standard deviations did not exceed 0.4 mm. Table 3 shows the standard deviation 811

E M Donovan, P Brabants, P M Evans et al Table 2. Summary of elliptical motion tracking experiment

Table 3. SD (mm) averaged over all tracked points and each

parameters carried out using the computer controlled platform to drive the test object

direction of motion for each of the linear motions of the test object driven by the Wellhofer beam data system

Experiment number

Eccentricity

Semi-major axis amplitude (cm)

Period (s)

Experiment number

SD (mm)

Ellipse Ellipse Ellipse Ellipse

0.6 0.6 0.6 1.0

2.0 1.0 1.0 1.0

6 2 4 4

Linear Linear Linear Linear Linear Linear Linear

0.64 1.43 0.55 0.33 1.14 0.49 0.42

1 2 3 4 (circle)

data averaged over all directions and for all points for all the experimental conditions.

Circular motion: positional data The data of Figure 4 show the standard deviation for the 2D circular motion in the Y–Z plane (GT, vertical) for each of the repeat cycles. The data are given for each tracked point averaged over the Y and Z positions recorded by the Osiris+. The standard deviation values were determined from the difference between the recorded values at each Osiris+ monitoring time point and the expected Y, Z values interpolated from the Wellhofer driven movements between points. The standard deviation does not exceed 1 mm and the average value over all cycles and points is 0.66 mm. Values of the X coordinate were recorded also, although the movement was not in this direction. The average value of one standard deviation over all cycles and points is 0.5 mm. Temporal data: linear and circular motion The data from both linear and circular movements of the test objects have been used to determine the sampling rates of the Osiris+ system. Figure 5 shows that these are 0.3– 0.4 s whilst the system is monitoring actively (i.e. ’’linac on’’ state) and 1.5–1.6 s in the ‘‘linac off’’ state.

Using the programmable computer controlled platform with motion in X,Y,Z Stationary point measurement Figure 6 summarizes the results obtained from the Osiris+ system when the test object was driven to

1 2 3 4 5 6 7

specific, known displacements from its origin using the computer controlled platform. The Osiris+ system was assessed by determining the deviation of the recorded values from the absolute position coordinates, as given by the platform. Figure 6 gives the absolute differences in millimetres between the platform values and the recorded data from Osiris+. All the differences were within ¡1 mm. The test object was returned to the origin between each set of measurements and no variation in this recorded value was found throughout the experiment.

Elliptical motion in 3D Data from the analysis of elliptical movements 1 and 3 are used as an example of the ability of the Osiris+ system to track motion in 3D. The results for the ellipse experiments 2 and 4 follow the same pattern; as these add no further information, the data are not given. As an example, Figure 7 shows the movement in the X direction (transverse) for ellipse 1. The true, calculated trajectory and the trajectories recorded by the Osiris+ system are shown for each of the five points. The absolute difference between the exact positions on the elliptical trajectories and the recorded positions were determined for each point for each of the X,Y,Z directions. These data are given in Figure 8 for ellipse 1. Figure 8b,c shows that the absolute errors in Y and Z are within ¡1 mm for the majority of the positions, but the data in Figure 8a show that for points 1,2 and 3 the absolute error values are only within ¡4 mm for ellipse motion 1 in the X direction.

Figure 3. Summary of errors for linear motion. The plot shows the deviation from the exact to tracked position averaged over all five cycles of the linear movement. 812



Figure 4. Summary of errors for circular motion. The values for each point are the average of the deviations from the calculated to tracked position.

Table 4 summarizes four parameters from the analysis of the elliptical movements 1 and 3. For each tracked point: the average displacement; the standard deviation of the variation and the minimum and maximum displacements are given. These are all averaged over the three cycles of the motion. The average value of the displacement is within ¡1 mm with an average standard deviation within 2.5 mm. The largest magnitudes of the minimum and maximum displacements are 24.72 mm and +4.0 mm, both for the ellipse 1 motion. In this experiment, the test object was moving at a fast speed and with the largest amplitude.

Discussion The tracking of patient movement during a fraction of a radical radiotherapy treatment requires a monitoring system with sufficient accuracy and precision. One potential application of such a system is to use the signal as a trigger to gate a linac for delivery based on the respiratory cycle as described by Berson et al [10]. Realtime monitoring of patient movements also has an application in situations where control of breathing is used to reduce an organ at risk dose, as suggested by Lu et al [23] and reported by Remouchamps et al [24]. We

have investigated the Osiris+ system in order to determine whether its performance characteristics were sufficient for this type of patient monitoring. The systems used for respiratory controlled studies as described in the literature tend to be either of the Active Breathing Control (ABC) device [11] (now from Elekta Oncology Systems, Crawley, UK) or the Varian Medical Systems RPMTM equipment (Varian Medical Systems, Palo Alto, CA) [10, 12]. These use a flow volume (the ABC device) or a marker on the patient’s chest (Varian RPM) to monitor respiratory motion and breath-hold over time, in one dimension. The studies by Berson et al [10] and Pedersen et al [12] show that the trace of the external marker on a patient’s chest has an amplitude of between 5 mm and 19 mm. It would be realistic to expect a monitoring system to detect movements of the external surface of a patient down to 2 mm. Yan et al [22] examined the Novalis Body system (BrainLAB Inc., Germany) which uses a combination of infrared and X-ray imaging for tracking markers. In an experiment similar to the absolute position checking, they moved a phantom known distances (2 mm, 5 mm, 10 mm and 20 mm) and determined the positional error of the system in X, Y and Z. They quote the average errors in the lateral direction as 0.6¡0.3 mm; those in the vertical direction as 0.7¡0.2 mm and those in the longitudinal

Figure 5. Summary of sampling intervals. Data are from linear and circular motion.


813

E M Donovan, P Brabants, P M Evans et al

Figure 6. Summary of errors for all points at all positions for the test object moved using the programmable platform.

direction as 0.5¡0.2 mm. The results from the Osiris+ system given in Figure 6 compare well with these data. Yan et al did not investigate the continuous tracking of markers on a moving object. Our assessment of the Osiris+ system covered a range of movements in one, two and three dimensions and different speeds. Figure 3 and 4 demonstrate that the errors in tracking for the sawtooth and 2D circular motions were within ¡1 mm for most conditions. Two situations, which were more extreme, were the linear 2 and 5 experiments. The former used the Osiris+ with the lights dimmed and the latter had a ¡5 cm amplitude, which is much larger than would be found in practice. Here the errors rose above 1.5 mm. Shirato et al [21] evaluated the performance of a Real-Time TumourTracking system (EXL-20DP; Mitsubishi Electronics Co., Ltd., Tokyo, Japan) using a phantom moving in a circle over a range of speeds from 6 mm s21 to 40 mm s21. They

quote the accuracy of the system as better than 1.5 mm; a similar level of accuracy to the Osiris+. The largest displacement errors were in the tracking of the points when the test object was driven in the elliptical movements 1 and 3, but only in the X direction (transverse) and only for three of the points. Here the error could reach ¡5 mm for ellipse 1 and ¡3 mm for ellipse 3 at the extremes of the travel. The marker at the top of the test object was seen by two different pairs of cameras and recorded twice as point 3 and point 4. Only the point 3 X direction errors were large for this marker. The reason is unclear and this is under further investigation. The differences in the phase of the data (shown in Figure 7) may indicate that the cameras were measuring at different time points as data points 1, 2 and 3 were recorded by one camera pair and the data from points 4, 5 were from the other camera pair. The mean

Figure 7. Data recorded from Osiris+ for displacement in X axis with time compared with programmed displacement for platform movement Ellipse 1. 814



Figure 8. (a) Absolute error in position in the X direction for ellipse 1 for all points. (b) Absolute error in position in the Y direction for ellipse 1 for all points. (c) Absolute error in position in the Z direction for ellipse 1 for all points.

displacement error in the X direction was 20.23 mm ¡ 1.7 mm for Ellipse 1. The mean error in Y was 0.12 mm ¡0.5 mm and that in the Z direction 20.09 mm ¡ 0.6 mm. For most situations investigated, the system The British Journal of Radiology, October 2006

performance was sufficient to meet the requirement to track movement with a minimum threshold of 2 mm and the imaging rate of one every 0.4 s was sufficient for good monitoring accuracy. 815

E M Donovan, P Brabants, P M Evans et al Table 4. Summary of error analysis data for elliptical movement 1 and 3 using the programmable platform Average (mm) Ellipse number

SD (mm)

Minimum (mm)

Maximum (mm)

1

3

1

3

1

3

1

3

20.3 20.01 0.02 20.34 0.19 0.06 20.07 0.03 20.04 20.12 0.26 20.35 20.31 0.14 20.15

0.48 20.43 0.76 0.56 20.58 0.45 0.7 20.51 0.54 0.56 20.36 0.51 0.39 20.54 0.43

2.37 0.52 0.6 2.32 0.39 0.73 2.28 0.32 0.6 0.81 0.59 0.6 0.56 0.47 0.65

1.54 0.37 0.39 1.43 0.35 0.51 1.6 0.39 0.41 0.69 0.55 0.57 0.62 0.45 0.62

24.72 20.93 21.47 24.52 20.6 21.47 24.52 20.81 21.47 22.12 20.99 21.89 21.52 20.72 21.43

21.88 21.35 20.21 21.68 21.19 20.81 21.98 21.39 20.45 20.54 21.81 20.52 20.45 21.81 20.72

3.61 1.27 1.19 3.31 1.17 1.19 4.0 0.92 0.95 1.5 1.89 0.78 0.74 1.11 0.93

3.15 0.41 1.47 3.05 0.21 1.37 3.38 0.26 1.47 2.03 0.45 1.65 1.83 0.44 1.75

Direction/point

X1 Y1 Z1 X2 Y2 Z2 X3 Y3 Z3 X4 Y4 Z4 X5 Y5 Z5

The errors are derived from the difference between the calculated position from the input function to the platform and that recorded by the Osiris+ tracking software at the sampled time points.

Conclusion Assessment of the Osiris+ system found that continuous linear, circular and elliptical motions were tracked with a standard deviation of 1 mm in the difference between measured and true positions of the test object markers. Under more extreme conditions, such as dimmed lighting and large amplitude of motion, this increased to 2 mm. The recorded values of static, known displacements by the system were also within ¡1 mm. The sampling rates of the system were 1.4–1.5 s prior to the command to track and 0.3–0.4 s during the tracking time. It is recommended that the system be used with the lights on and that regular calibrations are carried out to maintain its performance. The system performance is considered satisfactory to enable its use (i) in assessing external motion based on skin markers; (ii) in conjunction with electronic portal imaging to determine the nature, if any, of external-tointernal organ movement correlation and (iii) to provide a signal for respiratory gating.

References 1. Das IJ, Cheng C-W, Fosmire H, Kase KR, Fitzgerald TJ. Tolerances in setup and dosimetric errors in the radiation treatment of breast cancer. Int J Radiat Oncol Biol Phys 2000;26:883–90. 2. Lirette A, Pouliot J, Aubin M, Larochelle M. The role of electronic portal imaging in tangential breast irradiation: a prospective study. Radiother Oncol 1995;37:241–5. 3. Alasti H, Petric MP, Catton CN, Warde PR. Portal imaging for evaluation of daily on-line setup errors and offline organ motion during conformal irradiation of carcinoma of the prostate. Int J Radiat Oncol Biol Phys 2001;49:869–84. 4. Aubry JF, Beaulieu L, Girouard LM, Aubin S, Tremblay D, Laverdiere J, et al. Measurements of intrafraction motion and interfraction and intrafraction rotation of prostate by three-dimensional analysis of daily portal imaging with radiopaque markers. Int J Radiat Oncol Biol Phys 2004;60:30–9.

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5. Chung PWM, Haycocks T, Brown T, Cambridge Z, Kelly V, Alasti H, et al. On-line aSi portal imaging of implanted fiducial markers for the reduction of interfraction error during conformal radiotherapy of prostate carcinoma. Int J Radiat Oncol Biol Phys 2004;60:329–34. 6. Van de Steene J, Van den Heuvel F, Bel A, Verellen D, De Mey J, Noppen M, et al. Electronic portal imaging with on-line correction of setup error in thoracic irradiation: clinical evaluation. Int J Radiat Oncol Biol Phys 1998;40:967–76. 7. McKenzie A, van Herk M, Mijnheer B. Margins for geometric uncertainty around organs at risk in radiotherapy. Radiother Oncol 2002;62:299–307. 8. Ekberg L, Holmberg O, Wittgren L, Bjelkengren G, Landberg T. What margins should be added to the clinical target volume in radiotherapy treatment planning for lung cancer? Radiother Oncol 1998;48:71–7. 9. De Boer HCJ, Van Os MJH, Jansen PP, Heumen BJM. Application of the no action level (NAL) protocol to correct from prostate motion based on electronic portal imaging of implanted markers. Int J Radiat Oncol Biol Phys 2005;61:969–83. 10. Berson AM, Emery R, Rodriguez L, Richards GM, Ng T, Sanghavi S, et al. Clinical experience using respiratory gated radiation therapy: comparison of free-breathing and breath-hold techniques. Int J Radiat Oncol Biol Phys 2004;60:419–26. 11. Wong JW, Sharpe MB, Jaffray DA, Kini VR, Robertson JM, Stromberg JS, et al. The use of active breathing control (ABC) to reduce margin for breathing motion. Int J Radiat Oncol Biol Phys 1999;44:911–9. 12. Pedersen A, Korreman S, Nystrom H, Specht L. Breathing adapted radiotherapy of breast cancer: reduction of cardiac and pulmonary doses using voluntary inspiration breathhold. Radiother Oncol 2004;72:53–60. 13. Starkschall G, Forster KM, Kitamura K, Cardenas A, Tucker SL, Stevens CW. Correlation of gross tumor volume excursion with potential benefits of respiratory gating. Int J Radiat Oncol Biol Phys 2004;60:1291–7. 14. Engelsman M, Sharp GC, Bortfeld T, Onimaru R, Shirato H. How much margin reduction is possible through gating or breath hold? Phys Med Biol 2005;50:477–90. 15. Denissova SI, Yewondwossen MH, Andrew JW, Hale ME, Murphy CH, Purcell SR. A gated deep inspiration



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breath-hold radiation therapy technique using a linear position transducer. J Appl Clin Med Phys 2005;6:61–70. Berbeco RI, Nishioka S, Shirato H, Chen GT, Jiang SB. Residual motion of lung tumours in gated radiotherapy with external respiratory surrogates. Phys Med Biol 2005;50:3655–67. Ford EC, Mageras GS, Yorke E, Rosenzweig KE, Wagman R, Ling CC. Evaluation of respiratory movement during gated radiotherapy using film and electronic portal imaging. Int J Radiat Oncol Biol Phys 2002;52:522–31. Hoisak JDP, Sixel KE, Tirona R, Cheung PCF, Pignol J-P. Correlation of lung tumor motion with external surrogate indicators of respiration. Int J Radiat Oncol Biol Phys 2004;60:1298–306. Onimaru R, Shirato H, Fujino M, Suzuki K, Yamazaki K, Nishimura M, et al. The effect of tumor location and respiratory function on tumor movement estimated by realtime tracking radiotherapy (RTRT) system. Int J Radiat Oncol Biol Phys 2005;63:164–9.


20. Wilks RJ. An optical system for measuring surface shapes for radiotherapy planning. Br J Radiol 1993;66:351–9. 21. Shirato H, Shimizu S, Kunieda T, Kitamura K, van Herk M, Kagel K, et al. Physical aspects of a real-time tumor-tracking system for gated radiotherapy. Int J Radiat Oncol Biol Phys 2000;48:1187–95. 22. Yan H, Yin FF, Kim JH. A phantom study on the positioning accuracy of the Novalis Body system. Med Phys 2003;30:3052–60. 23. Lu H-M, Cash E, Chen MH, Chin L, Manning WJ, Harris J, et al. Reduction of cardiac volume in left-breast treatment fields by respiratory maneuvers: a CT study. Int J Radiat Oncol Biol Phys 2000;47:895–904. 24. Remouchamps VN, Vicini FA, Sharpe MB, Kestin LL, Martinez AA, Wong JW. Significant reductions in heart and lung doses using deep inspiration breath hold with active breathing control and intensity-modulated radiation therapy for patients treated with locoregional breast irradiation. Int J Radiat Oncol Biol Phys 2003;55:392–406.

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Normalized data for the estimation of fetal radiation dose from radiotherapy of the breast 1

B BRADLEY,

BSc,

2

A FLECK,

BSc, MSc

and

2,3

E K OSEI,

BSc, MSc, PhD

1

Department of Systems Design Engineering, University of Waterloo, 200 University Avenue West, Waterloo, Ontario, 2Department of Medical Physics, Grand River Regional Cancer Center, 835 King Street West, Kitchener, Ontario and 3Department of Physics, University of Waterloo, 200 University Avenue West, Waterloo, Ontario, Canada ABSTRACT. There can be several reasons why a pregnant patient may receive a radiological examination. It could have been a planned exposure, or the exposure might have resulted from an emergency when a thorough evaluation of pregnancy was impractical. Sometimes the pregnancy was unsuspected at the time of the examination and, with younger women being diagnosed with breast cancer, the likelihood of this will increase in radiotherapy departments. Whatever the reason, when presented with a pregnant patient who has received a radiological examination involving ionizing radiation, the dose to the fetus should be assessed based on the patient’s treatment plan. However, a major source of uncertainty in the estimation of fetal absorbed dose is the influence of fetal size and position as these change with gestational age. Consequently, dose to the fetus is related to gestational age. Various studies of fetal dose during pregnancy have appeared in the literature. Whilst these papers contain many useful data for estimating fetal dose, they usually contain limited data regarding the depth and size of the fetus within the maternal uterus. We have investigated doses to the fetus from radiation therapy of the breast of a pregnant patient using an anthropomorphic phantom. Normalized data for estimating fetal doses that takes into account the fetal size (gestational age: 8–20 weeks post-conception) and depth within the maternal abdomen (4–16 cm) for different treatment techniques have been provided. The data indicate that fetal dose is dependent on both depth within the maternal abdomen and gestational age, and hence these factors should always be considered when estimating fetal dose. The data show that fetal dose can be underestimated up to about 10% or overestimated up to about 30% if the dose to the uterus is assumed instead of the actual fetal dose. It can also be underestimated up to about 23% or overestimated up to about 12% if a mean depth of 9 cm is assumed, instead of using the actual depth of the fetus within the maternal abdomen. Multisegments sMLC technique showed consistently lower fetal doses compared with all the wedged plans employed.

Peripheral dose (PD), which is an inevitable consequence of radiotherapy, is mainly due to the radiation that is scattered within the patient and also the scattered and leakage radiation from the head of the machine and the collimator assembly. The magnitude of the scattered radiation from within the patient will mainly depend on the beam energy, distance from the radiation field edge, field size and depth. For a pregnant patient, such dose is of greater concern due to the fact that the developing fetus is more susceptible to the adverse effects of ionizing radiation, especially when the absorbed dose to the fetus may be high. However, using additional shielding can minimize the radiation dose due to scatter and leakage radiation from the machine head and collimator assembly [1, 2]. Address correspondence to: Ernest K Osei, Department of Medical Physics, Grand River Regional Cancer Center, 835 King Street West, Kitchener, Ontario, Canada. E-mail: [email protected].

818


Although radiotherapy treatment during a known pregnancy (especially during the 8–15 weeks gestational age period) should be avoided [3], it is not always possible to postpone treatment for the full duration of the pregnancy. Under such circumstances, it is essential to be able to estimate the absorbed dose (and hence risk) to the fetus if the patient must undergo treatment. Modifications to the plan and shielding techniques can then be incorporated where necessary to achieve an optimum balance between risk to the fetus and benefit to the mother, although the fetus also receives an indirect benefit. There are also occasions when pregnancy may not be known before the onset of a therapeutic treatment, and sometime later the patient realises she is pregnant. Under such a situation too, the fetal absorbed dose (and hence risk) as a result of the radiation treatment received should be assessed by a medical physicist. With younger women being diagnosed with breast cancer, the likelihood increases that a patient may be pregnant when The British Journal of Radiology, October 2006

Estimation of fetal dose from radiotherapy of the breast

treatment is indicated or that inadvertent treatment may take place for women who subsequently learn that they are pregnant. Fetal dose estimation is a difficult and time-consuming procedure that requires extensive dosimetric measurements and sometimes the construction of a specific phantom to simulate the patient’s geometry [3, 4]. A review of the scientific literature regarding the measurement of peripheral dose for the estimation of dose to the fetus from radiotherapy procedures has revealed an extensive collection of data [1, 5–14]. Mazonakis developed a method similar to the one described here for estimating fetal dose during brain radiotherapy. Using an anthropomorphic phantom, fetal absorbed dose was determined at two gestational ages and depths [4]. Rincon [9] has also estimated the dose to the fetus using peripheral dose to the uterus for two standard breast radiotherapy treatments (a non-wedged plan and a 15 ˚ physical wedged plan). Point dose measurements were taken in a phantom from the field edge to the uterus, and the fetal depth was assumed to be 15 cm. However, there are very limited data in the literature that take into account the different fetal positions (depth within the maternal abdomen) and gestational age (size) when estimating dose, although it has been shown that these fetal parameters are significant factors when estimating dose [2, 15]. Therefore, a source of data, which accounts for the different fetal parameters (i.e. fetal depth and size), would be very useful when estimating the fetal absorbed dose and hence risk. This paper presents the experimental procedure and results of fetal dose estimation from radiotherapy treatment of the breast. Data are presented for two different photon beam energies (6 MV and 15 MV), four different physical wedge angles (15 ˚, 30 ˚, 45 ˚ and 60 ˚), and a segmented multi-leaf collimator (sMLC) plan. Fetal sizes (gestational age) considered range from 8 weeks to 20 weeks post-conception, and fetal depth within the maternal abdomen ranges from 4 cm to 16 cm. By treating the fetus as a three-dimensional volume at various gestational ages and depths from the anterior surface of the mother’s abdomen, data are generated for a more accurate estimation of the mean dose to the fetus from radiotherapy of the breast. Fetal dose is assumed to be a whole body exposure of the fetus within the maternal abdomen. There are three scenarios where the data presented in this paper can be of benefit to the medical and paramedical personnel involved with fetal dose estimation. For the case of a patient who has already completed radiotherapy treatment and discovers later that she was pregnant during the time of treatment, the data would be useful in estimating the dose to the fetus and hence risk using the patient’s treatment plan parameters. Such inadvertent fetal exposures would be very rare if modern patient safety legislation is implemented well in radiological departments. For the case of a patient who, during the course of treatment, discovers that she is pregnant and there is clinical indication for her to complete treatment, the data would be useful to estimate the fetal dose from the treatment already received. The patient’s treatment plan could then be modified and optimized in order to minimize the fetal dose. Appropriate abdominal shielding could also be The British Journal of Radiology, October 2006

employed for the rest of the treatment to further reduce the fetal dose and hence risk and the actual dose measured for the rest of the treatment. Lastly, for the case of a patient who is pregnant and wishes to have radiotherapy treatment and there is also clinical indication for her to undergo treatment, the data would be useful to estimate the fetal dose without taking into account any form of shielding (maximum dose). A detailed study of the actual treatment plan could then be conducted, and the plan might be modified and optimized to reduce the fetal dose. Furthermore, appropriate abdominal shielding could then be employed to further reduce the fetal dose and risk from the treatment. Actual fetal dose should then be measured using the optimized treatment plan for risk assessment.

Materials and methods Experimental set-up The experimental set-up used for the estimation of fetal dose from radiation therapy of the breast is shown in Figure 1. An Alderson adult Rando anthropomorphic phantom (Alderson Research Labs, Stamford, CT) was used to represent the patient. The phantom consists of a human skeleton encased in tissue-equivalent material, with 33 transverse sections. It has an AP thickness of 20 cm, a width (shoulder-to-shoulder) of about 34 cm and a height of 90 cm. Radiographic film (Kodak X-OMAT TL, Rochester, NY) was used for all dosimetry measurements. The films were placed between each of the slices spanning the fetal region. The films spanned from slice 27 to 33 depending on the length of the fetus at the gestational age being studied (Figures 1 and 2). The film calibration curve was performed in a homogeneous medium. A sensitometric curve was generated for each batch of film at the time of measurement and used to convert optical densities into relative doses. The phantom was irradiated using the plans generated (6 MV and 15 MV (all wedges), and segmented MLC plan). The irradiated films were developed, scanned and analysed using a Vidar VXR-16 Dosimetry Pro scanner (Vidar, Herndon, VA) and RIT113 (v4) Radiation Therapy Dosimetry Software. Metal-oxide-semiconductor field-effect transistors (MOSFETs) were also used to measure point doses in order to verify the film dosimetry. The pegs along a central axis throughout the fetal region of the Rando phantom were removed and five high-sensitivity MOSFET dosemeters were placed with the sensitive region covering the holes (Figure 2). They were at a depth of approximately 11 cm from the anterior surface of the phantom and ranged from a distance of approximately 22.5 cm to 32.5 cm from the field edge. The phantom was placed in a supine position on the treatment couch in accordance with department protocol (Figure 1) and was ‘‘treated’’ with the same set-up for each treatment plan described below.

Treatment planning The phantom underwent a treatment planning CT scan. At the time of scanning, the phantom was placed in 819

B Bradley, A Fleck and E K Osei

Figure 1. The experimental setup for a typical two-field tangential breast treatment. The Rando phantom is placed in a supine position on the treatment couch and setup as per departmental protocol, and was ‘‘treated’’ with the same set-up for each treatment plan.

Figure 2. The films and Metaloxide-semiconductor field-effect transistors (MOSFETs) placement within the phantom showing the positioning of the films and MOSFETs. MOSFETs were located along a central axis within the fetal region. When measurements are being taken with film, the MOSFETs are removed and likewise, when measurements are being taken with the MOSFETs for comparison, the films are removed.

the treatment position for tangential breast radiation therapy adopted at the centre (i.e. supine position). A helical AcQSim CT scanner (Philips Medical Systems Cleveland Inc., Ohio, USA) was used to acquire contiguous 3 mm CT axial images. The CT data set was then transferred to a 3D treatment planning workstation (Pinnacle3; ADAC Laboratories, Milpitas, USA) for treatment planning. Using our departmental protocol for beam arrangements, a typical two-field tangential breast plan was developed. This standard plan was then modified to incorporate different photon energies (6 MV and 15 MV), wedge angles (15 ˚, 30 ˚, 45 ˚ and 60 ˚), and a multi-segment static MLC technique. The MLC leaves were fully retracted in all wedge plans. In total, 10 different treatment plans were generated for the study and they spanned the different plans usually used in the radiotherapy centre for the treatment of breast cancer.

Fetal parameters and dose calculation A chart of fetal dimensions was created for all gestational ages ranging from 8 weeks to 20 weeks post-conception, which are based on the mean of a series of measurements taken from the literature [16–19]. They gave the fetal crown–rump-length (CRL), biparietal diameter (BPD), abdominal circumference (AC), head circumference (HC) and fetal mass (wt) at the various 820

gestational ages, with statistical variation as 2 standard deviations from the mean. These data were used to create a three-dimensional rectangular volume to represent the fetus at the different gestational ages (Table 1). The density (r) of the fetus was assumed to be 1.061023 kg m23. For gestational ages ranging from 8 weeks to 12 weeks post-conception, the CRL was used for the length (l), the width (w) was estimated from the mean of the abdominal circumference (i.e. AC/p) and the anteroposterior (AP) thickness was calculated from: AP~

wt l|w|r

For gestational ages ranging from 13 weeks to 20 weeks post-conception, the biparietal diameter was used for the width, AP thickness was estimated from the mean abdominal circumference and the length was estimated in a similar way as above. Using the RTT113 software, a rectangular region of interest (ROI) (Figure 3) was selected on each film using the width and AP thickness dimensions of the fetus at the depth of interest. The mean dose and standard deviation for this region are recorded. This process is repeated for each subsequent film spanning the length of the fetus at the same depth and gestational age dimensions. From the complete set of mean doses for each film, the mean dose for the volume (gestational age) is estimated. The whole process is then The British Journal of Radiology, October 2006

Estimation of fetal dose from radiotherapy of the breast Table 1. Fetal dimensions at different gestational ages Gestational age (weeks)

Length (cm)

Width (cm)

AP thickness (cm)

Mass (g)

8 9 10 11 12

1.6 2.3 3.3 4.1 5.4

0.7 0.8 1.0 1.4 1.9

0.9 1.1 1.2 1.2 1.4

1.0 2.0 4.0 2.0 14.0

13 14 15 16 17 18 19 20

5.9 6.5 8.0 9.2 10.7 11.8 12.5 13.5

2.1 2.5 2.9 3.2 3.5 3.9 4.3 4.6

2.3 2.6 3.0 3.4 3.7 4.1 4.5 4.8

28.0 43.0 70.0 100.0 140.0 190.0 240.0 300.0

Calculations

Length5CRL Width5AC/p AP5calculated

Width5BPD AP5AC/p Length5calculated

CRL, crown–rump length; AC, abdominal circumference; AP, anteroposterior; BPD, biparietal diameter.

repeated for each gestational age (8–20 weeks) and depth (4–16 cm) of the fetus, and for all the films exposed from each treatment plan delivered.

Results and discussion In this paper, fetal dose is assumed to be a whole body exposure of the fetus within the maternal abdomen. The data collected and presented here were taken without any modification to the treatment plan setup to minimize the dose to the fetus. They would be useful for estimating dose to the fetus when a patient realises she is or was pregnant during or after the course of treatment. If there is a clinical indication for a pregnant patient to complete treatment when pregnancy is known, the data could be used to estimate the fetal dose without shielding and the setup could then be modified by the use of external shielding over the abdominal region of the pregnant patient during treatment to significantly reduce the fetal dose. The dimensions of the rectangular shape representing the fetus at different gestational ages are shown in

Table 1. Measured fetal doses normalized to the prescribed dose at the isocentre of the plan as a function of fetal depth within the maternal abdomen for different breast treatment plans are given in Tables 2–6. Figure 4 shows the relationship between the normalized fetal dose as a function of depth for all gestational ages using the multi-segment static MLC treatment plan and 6 MV photon beam. A similar relationship comparing the various treatment plans (sMLC, 15 ˚, 30 ˚, 45 ˚ and 60 ˚) and photon beam energies (6 MV and 15 MV) is shown in Figure 5. One of the most important parameters in the measurement of fetal dose from radiotherapy procedures is the distance from the radiation field edge to the point of measurement. The fetal dose decreases approximately exponentially with distance from the field edge. Published data [2, 4] show that the change in the peripheral dose with depth is small, although other data [1, 8] show greater change for a Co-60 treatment unit. The results of fetal dose estimation from radiotherapy of the breast show that fetal dose is dependent on both fetal depth within the maternal abdomen and gestational age (Figure 4). In general for all the techniques used and for

Figure 3. Region of interest representing a fetus at a gestational age of 20-weeks post-conception and at a depth of 8 cm from the abdominal surface of the maternal abdomen. The British Journal of Radiology, October 2006

821

B Bradley, A Fleck and E K Osei Table 2. Normalized mean fetal dose, as a percentage of prescribed dose (PD), for a typical two-field tangential beam plan using multi-segment static MLC for both 6 MV and 15 MV photon energies. Data are presented for various gestational ages (GA) and fetal depth within the maternal abdomen Photon energy

6 MV

15 MV

GA (weeks)

Fetal depth (cm) 4

5

6

7

8

9

10

11

12

16

8 9 10 11 12 13 14 15 16 17 18 19 20

0.103 0.104 0.106 0.107 0.110 0.112 0.114 0.119 0.125 0.133 0.140 0.144 0.151

0.102 0.103 0.105 0.106 0.109 0.111 0.113 0.118 0.124 0.131 0.138 0.142 0.149

0.100 0.101 0.103 0.104 0.107 0.109 0.111 0.116 0.121 0.129 0.135 0.139 0.146

0.099 0.099 0.100 0.102 0.105 0.106 0.108 0.113 0.118 0.126 0.132 0.136 0.142

0.097 0.097 0.099 0.100 0.103 0.104 0.106 0.110 0.115 0.122 0.128 0.132 0.138

0.095 0.096 0.097 0.098 0.100 0.101 0.103 0.108 0.112 0.119 0.124 0.128 0.133

0.093 0.093 0.094 0.096 0.098 0.099 0.100 0.105 0.109 0.115 0.119 0.123 0.128

0.090 0.091 0.092 0.093 0.095 0.096 0.098 0.101 0.105 0.110 0.114 0.117 0.122

0.087 0.088 0.089 0.090 0.093 0.094 0.095 0.098 0.101 0.106 0.109 0.111 0.115

0.068 0.069 0.070 0.071 0.073 0.073 0.074 0.076 0.079 0.081 0.083 0.085 0.087

8 9 10 11 12 13 14 15 16 17 18 19 20

0.088 0.089 0.090 0.091 0.093 0.095 0.095 0.099 0.102 0.105 0.109 0.111 0.115

0.086 0.087 0.089 0.090 0.092 0.093 0.094 0.097 0.100 0.103 0.106 0.108 0.110

0.084 0.085 0.087 0.088 0.090 0.091 0.092 0.095 0.098 0.101 0.103 0.105 0.107

0.082 0.083 0.085 0.086 0.088 0.089 0.090 0.093 0.096 0.099 0.101 0.102 0.105

0.080 0.081 0.083 0.084 0.086 0.087 0.088 0.091 0.093 0.096 0.098 0.100 0.102

0.077 0.078 0.080 0.082 0.084 0.085 0.086 0.088 0.091 0.093 0.095 0.097 0.099

0.075 0.076 0.078 0.079 0.081 0.082 0.083 0.086 0.088 0.091 0.093 0.094 0.096

0.072 0.073 0.075 0.076 0.079 0.079 0.080 0.083 0.085 0.087 0.089 0.090 0.092

0.069 0.070 0.072 0.073 0.075 0.076 0.077 0.079 0.081 0.084 0.085 0.086 0.088

0.054 0.054 0.055 0.056 0.057 0.057 0.058 0.060 0.061 0.063 0.064 0.065 0.067

Table 3. Normalized mean fetal dose, as a percentage of prescribed dose (PD), for a typical two-field tangential beam 15 ˚ wedged plan for both 6 MV and 15 MV photon energies. Data are presented for various gestational ages (GA) and fetal depth within the maternal abdomen Photon energy

6 MV

15 MV

822

GA (weeks)

Fetal depth (cm) 4

5

6

7

8

9

10

11

12

16

8 9 10 11 12 13 14 15 16 17 18 19 20

0.325 0.331 0.342 0.350 0.365 0.370 0.378 0.396 0.412 0.433 0.450 0.460 0.476

0.323 0.330 0.341 0.349 0.364 0.369 0.377 0.395 0.411 0.432 0.448 0.459 0.474

0.321 0.328 0.338 0.346 0.361 0.367 0.374 0.393 0.409 0.430 0.446 0.456 0.471

0.319 0.326 0.336 0.344 0.358 0.364 0.371 0.389 0.405 0.425 0.441 0.451 0.466

0.315 0.322 0.332 0.340 0.354 0.359 0.366 0.384 0.399 0.419 0.434 0.444 0.459

0.308 0.315 0.325 0.333 0.347 0.353 0.360 0.377 0.392 0.411 0.426 0.435 0.449

0.301 0.307 0.317 0.325 0.339 0.344 0.351 0.368 0.382 0.401 0.415 0.424 0.438

0.292 0.298 0.308 0.316 0.329 0.334 0.340 0.357 0.371 0.389 0.402 0.411 0.423

0.280 0.287 0.296 0.304 0.317 0.322 0.328 0.344 0.357 0.374 0.387 0.395 0.406

0.218 0.221 0.227 0.232 0.241 0.245 0.250 0.263 0.274 0.289 0.300 0.308 0.319

8 9 10 11 12 13 14 15 16 17 18 19 20

0.205 0.209 0.216 0.221 0.230 0.234 0.239 0.251 0.261 0.275 0.286 0.293 0.304

0.203 0.207 0.214 0.219 0.228 0.232 0.236 0.248 0.258 0.271 0.281 0.288 0.297

0.198 0.203 0.209 0.215 0.224 0.228 0.232 0.244 0.254 0.266 0.276 0.283 0.292

0.193 0.198 0.204 0.209 0.218 0.222 0.227 0.238 0.248 0.261 0.270 0.277 0.286

0.188 0.192 0.198 0.204 0.213 0.216 0.221 0.232 0.242 0.254 0.264 0.270 0.279

0.181 0.186 0.192 0.198 0.207 0.210 0.215 0.226 0.235 0.247 0.256 0.262 0.271

0.175 0.180 0.186 0.191 0.200 0.204 0.208 0.219 0.228 0.239 0.248 0.254 0.262

0.167 0.172 0.179 0.184 0.193 0.196 0.200 0.211 0.219 0.230 0.238 0.243 0.251

0.157 0.162 0.169 0.174 0.183 0.186 0.190 0.200 0.208 0.219 0.227 0.231 0.238

0.120 0.122 0.125 0.128 0.134 0.137 0.140 0.147 0.153 0.162 0.169 0.173 0.180


Estimation of fetal dose from radiotherapy of the breast Table 4. Normalized mean fetal dose, as a percentage of prescribed dose (PD), for a typical two-field tangential beam 30 ˚ wedged plan for both 6 MV and 15 MV photon energies. Data are presented for various gestational ages (GA) and fetal depth within the maternal abdomen Photon energy

6 MV

15 MV

GA (weeks)

Fetal depth (cm) 4

5

6

7

8

9

10

11

12

16

8 9 10 11 12 13 14 15 16 17 18 19 20

0.430 0.438 0.451 0.462 0.481 0.488 0.497 0.522 0.544 0.573 0.595 0.610 0.632

0.431 0.440 0.453 0.464 0.482 0.490 0.499 0.524 0.546 0.575 0.597 0.611 0.633

0.430 0.439 0.452 0.463 0.482 0.489 0.499 0.524 0.545 0.574 0.596 0.610 0.632

0.428 0.437 0.449 0.460 0.479 0.486 0.496 0.521 0.542 0.571 0.592 0.607 0.628

0.424 0.432 0.445 0.455 0.474 0.481 0.491 0.515 0.536 0.564 0.586 0.600 0.621

0.417 0.425 0.438 0.448 0.467 0.474 0.483 0.507 0.528 0.555 0.576 0.590 0.611

0.407 0.416 0.428 0.441 0.457 0.464 0.473 0.497 0.517 0.544 0.564 0.575 0.597

0.395 0.404 0.416 0.427 0.445 0.452 0.460 0.484 0.503 0.529 0.548 0.560 0.579

0.381 0.390 0.402 0.413 0.430 0.437 0.445 0.467 0.485 0.509 0.527 0.539 0.556

0.302 0.307 0.314 0.320 0.332 0.337 0.344 0.361 0.376 0.396 0.412 0.422 0.437

8 9 10 11 12 13 14 15 16 17 18 19 20

0.295 0.301 0.309 0.316 0.329 0.334 0.340 0.356 0.370 0.388 0.403 0.412 0.426

0.292 0.298 0.307 0.314 0.327 0.331 0.337 0.353 0.366 0.384 0.397 0.406 0.419

0.286 0.293 0.302 0.309 0.322 0.327 0.333 0.349 0.362 0.379 0.391 0.400 0.412

0.281 0.287 0.297 0.304 0.317 0.322 0.328 0.343 0.356 0.372 0.384 0.392 0.403

0.275 0.281 0.290 0.298 0.311 0.315 0.321 0.336 0.348 0.364 0.375 0.383 0.394

0.267 0.273 0.283 0.290 0.302 0.307 0.313 0.327 0.339 0.354 0.366 0.373 0.383

0.259 0.265 0.274 0.281 0.293 0.298 0.305 0.318 0.329 0.344 0.354 0.361 0.371

0.249 0.255 0.265 0.272 0.284 0.288 0.293 0.307 0.318 0.331 0.342 0.348 0.357

0.238 0.244 0.253 0.261 0.272 0.276 0.281 0.294 0.304 0.317 0.326 0.332 0.341

0.184 0.188 0.194 0.199 0.207 0.211 0.215 0.225 0.233 0.244 0.252 0.257 0.265

Table 5. Normalized mean fetal dose, as a percentage of prescribed dose (PD), for a typical two-field tangential beam 45 ˚ wedged plan for both 6 MV and 15 MV photon energies. Data are presented for various gestational ages (GA) and fetal depth within the maternal abdomen Photon energy

6 MV

15 MV

GA (weeks)

Fetal depth (cm) 4

5

6

7

8

9

10

11

12

16

8 9 10 11 12 13 14 15 16 17 18 19 20

0.334 0.341 0.350 0.359 0.373 0.379 0.386 0.405 0.421 0.443 0.459 0.470 0.486

0.330 0.337 0.347 0.355 0.370 0.375 0.382 0.401 0.417 0.438 0.454 0.465 0.480

0.324 0.331 0.341 0.349 0.364 0.370 0.377 0.395 0.411 0.432 0.447 0.458 0.473

0.318 0.324 0.334 0.343 0.357 0.363 0.370 0.388 0.403 0.424 0.439 0.449 0.464

0.312 0.318 0.328 0.336 0.350 0.355 0.362 0.380 0.395 0.414 0.429 0.439 0.454

0.304 0.311 0.320 0.328 0.342 0.347 0.354 0.371 0.385 0.405 0.419 0.429 0.443

0.297 0.303 0.312 0.320 0.333 0.338 0.344 0.361 0.375 0.394 0.408 0.417 0.431

0.287 0.293 0.302 0.310 0.322 0.327 0.333 0.349 0.363 0.381 0.395 0.403 0.416

0.273 0.279 0.288 0.296 0.308 0.313 0.319 0.335 0.348 0.366 0.378 0.386 0.399

0.208 0.211 0.217 0.222 0.231 0.236 0.241 0.254 0.266 0.281 0.293 0.301 0.313

8 9 10 11 12 13 14 15 16 17 18 19 20

0.299 0.304 0.312 0.319 0.331 0.335 0.341 0.356 0.370 0.387 0.401 0.410 0.424

0.294 0.299 0.308 0.315 0.326 0.331 0.336 0.351 0.364 0.380 0.392 0.400 0.413

0.287 0.292 0.301 0.308 0.320 0.324 0.330 0.344 0.356 0.372 0.384 0.392 0.404

0.279 0.284 0.293 0.299 0.311 0.315 0.321 0.335 0.347 0.363 0.375 0.383 0.394

0.271 0.276 0.284 0.291 0.302 0.306 0.312 0.326 0.338 0.353 0.365 0.373 0.384

0.262 0.267 0.275 0.282 0.293 0.298 0.303 0.317 0.329 0.344 0.355 0.362 0.373

0.254 0.259 0.267 0.274 0.285 0.289 0.295 0.308 0.320 0.334 0.345 0.352 0.362

0.244 0.250 0.258 0.265 0.276 0.280 0.286 0.299 0.309 0.323 0.333 0.339 0.349

0.233 0.239 0.248 0.255 0.266 0.269 0.274 0.287 0.297 0.310 0.320 0.326 0.334

0.183 0.187 0.193 0.198 0.206 0.209 0.213 0.224 0.232 0.244 0.252 0.258 0.267


823

B Bradley, A Fleck and E K Osei Table 6. Normalized mean fetal dose, as a percentage of prescribed dose (PD), for a typical two-field tangential beam 60 ˚ wedged plan for both 6 MV and 15 MV photon energies. Data are presented for various gestational ages (GA) and fetal depth within the maternal abdomen Photon energy

6 MV

15 MV

GA (weeks)

Fetal depth (cm) 4

5

6

7

8

9

10

11

12

16

8 9 10 11 12 13 14 15 16 17 18 19 20

0.381 0.389 0.400 0.409 0.425 0.431 0.439 0.460 0.478 0.502 0.520 0.532 0.550

0.377 0.384 0.395 0.404 0.421 0.427 0.435 0.456 0.474 0.497 0.516 0.527 0.545

0.370 0.378 0.389 0.398 0.415 0.421 0.429 0.450 0.468 0.491 0.510 0.522 0.539

0.362 0.370 0.381 0.390 0.407 0.413 0.421 0.443 0.460 0.484 0.502 0.514 0.532

0.355 0.363 0.374 0.383 0.399 0.405 0.413 0.434 0.452 0.475 0.493 0.505 0.523

0.348 0.355 0.366 0.375 0.391 0.397 0.405 0.425 0.443 0.466 0.484 0.495 0.513

0.341 0.348 0.358 0.366 0.382 0.388 0.395 0.415 0.433 0.455 0.473 0.484 0.501

0.331 0.338 0.348 0.356 0.371 0.377 0.384 0.404 0.420 0.442 0.459 0.470 0.486

0.317 0.324 0.334 0.343 0.358 0.363 0.371 0.389 0.405 0.426 0.442 0.452 0.467

0.245 0.251 0.259 0.266 0.278 0.283 0.289 0.305 0.319 0.336 0.351 0.360 0.373

8 9 10 11 12 13 14 15 16 17 18 19 20

0.339 0.345 0.353 0.361 0.373 0.378 0.384 0.401 0.416 0.435 0.450 0.460 0.475

0.333 0.339 0.348 0.355 0.368 0.372 0.378 0.400 0.409 0.427 0.436 0.451 0.465

0.326 0.331 0.340 0.347 0.360 0.365 0.371 0.388 0.401 0.420 0.434 0.443 0.457

0.318 0.324 0.332 0.339 0.352 0.357 0.363 0.379 0.393 0.411 0.425 0.434 0.448

0.311 0.316 0.325 0.332 0.344 0.349 0.355 0.371 0.384 0.402 0.416 0.425 0.438

0.303 0.308 0.317 0.324 0.336 0.341 0.349 0.363 0.376 0.393 0.406 0.415 0.427

0.294 0.300 0.309 0.316 0.329 0.333 0.339 0.354 0.367 0.383 0.396 0.404 0.415

0.285 0.291 0.300 0.308 0.320 0.324 0.329 0.344 0.356 0.372 0.383 0.391 0.402

0.271 0.277 0.287 0.294 0.307 0.311 0.317 0.331 0.343 0.358 0.369 0.376 0.386

0.213 0.218 0.225 0.231 0.241 0.245 0.250 0.263 0.273 0.286 0.296 0.303 0.313

all gestational ages, the fetal dose decreases with increase in depth. The data show that fetal dose estimations based on using a constant fetal depth of 9 cm would either overestimate or underestimate the dose, depending on the depth of interest. For example, dose could be underestimated up to about 10% at a depth of 4 cm and overestimated as high as about 30% at a depth of 16 cm (using 6 MV, sMLC plan) if a depth of 9 cm is used instead of the actual fetal depth. Fetal dose is also dependent on the gestational age of the fetus. For all fetal depths, as gestational age increases, fetal dose also increases. The data show that if fetal dose were to be estimated as the dose to the uterus, the fetal dose would again be either overestimated or underestimated at any given depth depending on the gestational age of interest. A uterine mass of 66.3 g and a length, width and AP thickness of 7.5 cm, 3.5 cm and 2.5 cm, respectively, correspond most closely to the dimensions of a fetus at 15 weeks post-conception. Therefore, for dose estimates below 15 weeks postconception (i.e. uterus dose), the fetal dose would be overestimated by up to about 12%, and for estimates above 15 weeks, the fetal dose would be underestimated by up to about 23% (using 6 MV with the sMLC plan). The degree to which the uterus dose differs from fetal dose is dependent upon the difference in gestational age, as well as other parameters such as depth and orientation of the fetus within the maternal abdomen. According to the literature [1, 6, 7], the presence of a wedge in the beam could increase the fetal dose by a 824

factor of about 2–4 and our data give a similar result ranging from a factor of about 2–5 for all the wedges. The data show that the multi-segment sMLC technique results in consistently lower fetal doses compared with all the wedged techniques and could significantly reduce fetal doses. This reduction in the fetal dose with the use of a multi-segment (sMLC) plan as against that of a wedged plan may be as a result of a combination of a decrease in scatter and improved collimation when MLC leaves are used in conjunction with the jaws to shape the treatment field. The MLC leaves were fully retracted in all plans employing a wedge. The data indicate that fetal dose is dependent on both depth within the maternal abdomen and gestational age, and hence these factors should always be considered when estimating fetal dose. The measured fetal doses are in good agreement with other data obtained from the literature [1, 2, 4, 5, 9]. Comparison was made between these data and others previously published, although most of the published data refer to fetal doses at one specific gestational age and depth. Rincon et al [9] reported a fetal dose (assuming dose to the uterus) of about 40 mGy at a depth of 15 cm using a 6 MV photon beam and a prescription of 50 Gy at the isocentre. Assuming a fetal size of 15 weeks post-conception to be that of the uterus (as used in the Rincon study), 6 MV photon beam and sMLC plan, we estimated a fetal dose of about 38 mGy at a depth of 16 cm. The difference in the estimated doses may be as a result of differences in fetal depth and irradiation geometry used. For a fetus at 2–6 weeks The British Journal of Radiology, October 2006

Estimation of fetal dose from radiotherapy of the breast

Figure 4. Normalized mean fetal dose, as a percentage of prescribed dose (PD), for a typical two-field tangential beam plan using multi-segment static MLC for 6 MV photon beam as a function of fetal depth within the maternal abdomen. Data are presented for various gestational ages (GA).

gestational age at a depth of 9 cm, Antypas et al [5] reported relative fetal doses in the range of 0.079% to 0.085% of the prescribed dose. Using the same fetal depth and 6 MV photon beam and sMLC plan, we estimated a fetal dose of about 0.095% of the prescribed dose for a gestational age of 8 weeks. The difference in the estimated doses may be due to differences in gestational age and irradiation geometry employed.

Conclusion Normalized data for converting the prescribed dose at the isocentre from medical exposure of a pregnant woman undergoing radiation treatment of the breast, to absorbed dose to the fetus, have been presented. These data may be useful for estimating absorbed dose to the fetus from breast cancer radiotherapy of the mother and take into account the dependence of fetal age on gestational age (fetal size), fetal depth, wedge dose and photon beam energy. The results show that fetal dose is dependent on both gestational age and depth within the maternal abdomen, and hence these factors should be The British Journal of Radiology, October 2006

taken into account when estimating dose. The data collected and presented here were taken without any modification to the treatment setup to minimize the dose to the fetus. They would be useful for estimating dose to the fetus when a patient realises she is or was pregnant during or after the course of treatment. Again, if there is a clinical indication for a pregnant patient to complete treatment when pregnancy is known, the data could be useful for estimating the fetal dose and the treatment setup can then be modified and optimized, and also external shielding over the abdominal region of the pregnant patient could be employed during treatment to significantly reduce the fetal dose. The sources of radiation dose to the fetus are mainly due to the radiation that is scattered within the patient and also the scattered and leakage radiation from the head of the machine and the collimator assembly. Whereas little can be done to reduce the internal scatter within the patient contributing dose to the fetus, the other sources of radiation exposure contributing dose to the fetus could be reduced significantly by the application of external shielding over the abdominal region of the pregnant patient. 825

B Bradley, A Fleck and E K Osei

Figure 5. Normalized mean fetal dose, as a percentage of prescribed dose (PD), for a typical two-field tangential beam plan as a function of fetal depth within the maternal abdomen. Data are presented for various treatment plans (sMLC, 15 ˚, 30 ˚, 45 ˚ and 60 ˚ wedged plans) and for 6 MV and 15 MV photon energies. The gestational age is 15 weeks post-conception.

References 1. Stovall M, Blackwell CR, Cundiff J, Novack DH, et al. Fetal dose from radiotherapy with photon beams: Report of AAPM Radiation Therapy Committee Task Group No. 36. Med Phys 1995;22:63–82. 2. Islam MK, Saeedi F, Al-Rajhi N. A simplified shielding approach for limiting fetal dose during radiation therapy of pregnant patients. Int J Radiat Oncol Biol Phys 2001;49:1469–73. 3. ICRP-60. International Commission on Radiological Protection (ICRP). 1990 Recommendations of the ICRP. ICRP publication 60, Annals of the ICRP 21, No 1-3, Pergamon Press, Oxford, 1991. 4. Mazonakis M, Damilakis J, Varveris H, Theoharopoulos N, Gourtsoyiannis N. A method of estimating fetal dose during brain radiation therapy. Int J Radiat Oncol Biol Phys 1999;44:455–9. 5. Antypas C, Sandilos P, Kouvaris J, Balafouta E, et al. Fetal dose evaluation during breast cancer radiotherapy. Int J Radiat Oncol Biol Phys 1998;40:995–9. 6. Fraass B, Van de Geijn J. Peripheral dose from megavolt beams. Med Phys 1983;10:809–18. 7. Diallo I, Lamon A, Shamsaldin A, Grimaud E, de Vathaire F, Chavaudra J. Estimation of the radiation dose delivered

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8.

9.

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to any point outside the target volume per patient treated with external beam radiotherapy. Radiother Oncol 1996;38: 269–71. Van der Giessen P. A simple and generally applicable method to estimate the peripheral dose in radiation teletherapy with high energy x-rays or gamma radiation. Int J Radiat Oncol Biol Phys 1996;35:1059–68. Rincon CM, Jerez Sainz I, Farree IM, Espana Lopez ML, Franco PL, et al. Evaluation of the peripheral dose to uterus in breast carcinoma radiotherapy. Radiat Prot Dosim 2002;101:469–71. Nuyttens JJ, Prado KL, Jenrette JM, Williams TE. Fetal dose during radiotherapy: clinical implementation and review of the literature. Cancer (Radiothe´rapie) 2002;6: 352–7. Prado KL, Nelson SJ, Nuyttens JJ, Williams TE, Vanek KN. Clinical implementation of the AAPM Task Group 36 recommendations on fetal dose from radiotherapy with photon beams: a head and neck irradiation case report. J Appl Clin Med Phys 2000;1:1–7. Sneed PK, Albright NW, Wara WM, Prados MD, Wilson CB. Fetal dose estimates for radiotherapy of brain tumors during pregnancy. Int J Radiat Oncol Biol Phys 1995;32:823–30.


Estimation of fetal dose from radiotherapy of the breast 13. Antolak JA, Strom EA. Fetal dose estimates for electronbeam treatment to the chest wall of a pregnant patient. Med Phys 1998;25:2388–91. 14. Ngu SL, Duval P, Collins C. Fetal radiation dose in radiotherapy for breast cancer. Aust Radiol 1992;36:321–2. 15. Osei EK, Faulkner K. Fetal position and size data for dose estimation. Br J Radiol 1999;72:363–70. 16. Doublet PM, Benson CD, Nadel AS. Improved birth weight table for neonates developed from gestations dated by early ultrasonography. J Ultrasound Med 1997;16:241.


17. Hadlock FP, Shah YP, Kanon DJ, Lindsey JV. Fetal crownrump length: reevaluation of relation to menstrual age (5–18 weeks) with high-resolution real-time US. Radiology 1992;182:501–5. 18. Chitty LS, Altman DG, Henderson A, Campbell S. Charts of fetal size: 3. Abdominal measurements. Br J Obstet Gynaecol 1994;101:125–31. 19. Hadlock FP, Deter RL, et al. Estimating fetal age: computerassisted analysis of multiple fetal growth parameters. Radiology 1984;152:497–501.

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Dosimetric and treatment planning considerations for radiotherapy of the chest wall M M ASPRADAKIS,

PhD, MIPEM,

H M McCALLUM,

PhD, MIPEM

and N WILSON,

DCR(T)

Regional Medical Physics Department, Newcastle General Hospital, Westgate Road, Newcastle upon Tyne NE4 6BE, UK ABSTRACT. Radiotherapy treatment planning calculations of the chest wall are complex due to missing tissue, the thin chest wall and the presence of lung. The accuracy of calculated dose is dependent on the type of algorithm employed. This work evaluates a collapsed cone (CC) and a pencil beam (PB) convolution model for radiotherapy planning of the chest wall. Various irradiation geometries simulating the chest wall have been examined and calculations were compared with measurements with an ionization chamber in epoxy resin water substitute and in low-density lung substitute blocks. A retrospective treatment planning study comprising 6 patients was carried out to evaluate the differences in the dose distributions and monitor units predicted by the two algorithms. The calculated dose in unit density medium was within ¡1% for the CC model and up to ¡2% for PB. In low density medium and under full scatter conditions, CC overestimated the dose by 1% whereas PB overestimated the dose by 9%. In the tangential irradiation geometry with water and lung media, the PB overestimated dose to the isocentre by up to 10%, whereas the dose from CC was within 3%. From the treatment planning study calculated monitor units (MU) and doses were consistent with the experimental findings. The CC model is more accurate for radiotherapy treatment planning of the chest wall and especially when there is significant involvement of lung tissue.

Breast cancer affects up to 1 in 10 women in the Western World and is the most common female cancer in the UK in terms of incidence and mortality [1]. Adjuvant radiotherapy to the chest wall has an established role in reducing the risk of locoregional recurrence following mastectomy. A meta-analysis of all randomized controlled trials published in 1987 showed that post mastectomy radiotherapy was associated with a 66% reduction in the risk of locoregional recurrence [2]. Post mastectomy radiotherapy may be indicated either due to inadequate surgical resection of tumour margins due to lymphovascular involvement or due to a tumour size greater than 5 cm. Radiotherapy treatment planning calculations for the chest wall are complex due to missing tissue and the presence of lung tissue within the treatment field. Accurate calculation of the dose distribution is important as doses to neighbouring organs at risk, such as heart, lungs and contralateral breast, needs to be minimized. Treatment planning systems (TPS) would, ideally, provide both accurate relative dose distributions and monitor unit settings for such irradiation geometries. The Helax-TMS (version 6.1a) TPS (Nucletron B.V., The Netherlands) has two dose calculation algorithms available for external photon beam planning, the pencil beam (PB) and the collapsed cone (CC) algorithms. Dose Current address for M M Aspradakis: Klinik fu¨r Radio-Onkologie, Universita¨ts Spital Zu¨rich, Ra¨mistrasse 100, Zu¨rich 8091, Switzerland. Address correspondence to: Dr Hazel M McCallum.

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Received 21 October 2005 Revised 14 February 2006 Accepted 20 February 2006 DOI: 10.1259/bjr/26575438 ’ 2006 The British Institute of Radiology

calculations for patients on Helax-TMS use density information from CT based on the conversion of Hounsfield numbers (HN) to material density, thus accounting for the attenuation properties of different tissues, as well as using information on the material composition. The two algorithms differ primarily in how they model radiation transport and calculate dose in heterogeneous media, with the CC algorithm better approximating the dose directly to the medium [3–5]. Their characteristics and dosimetric accuracy in simple and complex phantom irradiation geometries have been well documented [6–9]. It has been shown that there is little or no clinically significant difference in the calculation of dose by the two algorithms when no large non-unit density heterogeneities are present and under full scatter conditions, such as in the irradiation of the pelvis [9]. This excludes cases of pelvic irradiation with high density metal prosthesis [10]. In cases of treatment planning in the thorax region (for oesophagus and lung), differences between CC and PB in the calculation of dose (predicted monitor units (MU)), planning target volume (PTV) coverage and minimum dose to the PTV has been reported [11]. Such differences can potentially be clinically significant and affect tumour control. At tangential beam irradiations for a range of beam energies (4 MV, 6 MV, 15 MV) on a homogeneous medium, experimental verification of the Helax-TMS algorithms using a 15 cm615 cm square field had shown the CC algorithm to be accurate to within ¡2% of the measured dose and to model closely the reduction in dose due to the missing tissue [9]. From PB the dose The British Journal of Radiology, October 2006

Photon dose calculations in chest wall radiotherapy

close to phantom boundaries was overestimated by 4–5% at 4 MV, but less so at higher energies. The difficulties encountered with treatment planning of the chest wall in particular have not been reported, nor has the dosimetric accuracy of treatment planning calculations for such extreme irradiation geometries with the presence of lung and/or the plan normalization point within thin tissue or within lung been investigated. It was the purpose of this work to investigate the dosimetric accuracy of the Helax-TMS convolution/ superposition algorithms for treatment planning of the chest wall at the Northern Centre for Cancer Treatment (NCCT), in Newcastle upon Tyne, UK. The dosimetric differences between the Helax-TMS CC and PB algorithms were examined for clinically relevant scenarios of chest wall irradiations. A planning study comprising six patients was also undertaken to support changes in clinical practice.

Materials and methods Dosimetric verification of Helax-TMS algorithms for chest wall irradiation Experimental set-up All measurements were carried out with a 6 MV photon beam on a Siemens Primus accelerator (Siemens,

Erlangen, Germany). This is the most frequently used beam energy at NCCT for treatment of the breast and chest wall. Three experimental geometries relevant to chest wall irradiation were examined. The WT1-tissue and lung substitute blocks (Epoxy resin water substitute blocks with elemental composition: H(8.09); C(67.22); N(2.4); O(19.84); Ca(2.32); Cl(0.13) and relative electron density of 1.02 (mass density 1.04 g cm23) [12]. Lung equivalent blocks of elemental composition: H(8.38); C(60.50); N(1.68); O(17.28); Cl(0.15); Si(0.84); Mg(11.17) and relative electron density of 0.25 (mass density of 0.26 g cm23) [13, 14]) used in the measurements were manufactured by ScanPlus (St Bartholomews, London). The measurement system used was an NE2571 0.6 cm3 cylindrical ionization chamber and its associated electrometer with a valid calibration factor (traceable to the primary standard at the National Physical Laboratory) to convert the ionization reading to dose in water. The irradiation geometries used are shown in Figure 1. The ionization chamber was positioned in all cases at the isocentre. In Figure 1a, the lateral extent of the phantom normal to the beam central axis varied from 1 cm to 3 cm and a measurement was also made at full scatter conditions with a lateral thickness of 10 cm. In Figure 1b the phantom comprised lung and tissue equivalent blocks. Measurements were carried out with the chamber in the tissue equivalent block at 1 cm, 1.5 cm and 2 cm from the edge of the lung heterogeneity and the phantom boundary varied from 1 cm to 3 cm

Figure 1. Experimental geometries employed for measurements and/or calculations in this work. All measurements were carried out with a 6 MV photon beam, gantry angle of 270 ˚ and collimator angle of 90 ˚. The British Journal of Radiology, October 2006

829

M M Aspradakis, H M McCallum and N Wilson

and at full scatter conditions at 10 cm in the lateral direction. Figure 1c was analogous to Figure 1b, but the calculation point was in lung equivalent material. The position of the normalization/prescription point with respect to the phantom boundary or the lung was chosen to be similar to that often selected in patient treatment plans. Simulations on Helax-TMS were carried out for the slice separations and dose matrix resolutions of 0.25 cm.

is placed within lung, a replacement perturbation correction was applied to the dose from the part of the field irradiating the lung medium (D2571 homogeneous lung ). Thus, if D2571 meas,heterogeneous is the measured dose in lung in the heterogeneous medium and p2571 lung,repl (Aeq ) is the replacement perturbation correction for the equivalent field size Aeq irradiating the low density medium, the dose to water in an infinitesimally small volume of water within the lung in the heterogeneous medium is calculated from [15]:

Low density dosemeter correction factor For the measurement of dose in a low density medium using the NE2571 ionization chamber and its absorbed 2571 dose to water calibration factor (Nwater,6MV ), the chamber reading is converted to dose in an infinitesimally small volume of water within the low density medium (which is what is also modelled by the treatment planning system), without accounting for the high replacement perturbation effect caused by the chamber and its high density graphite wall in the low density medium. The NE2571 chamber perturbs fluence differently in water and in low density lung and this difference should be accounted for when converting the ionization reading to dose. The application of a perturbation correction to the measurement in lung in this work was based on the work of Krieger and Sauer [15]. The replacement perturbation factor prepl(A) at field size A was defined as the ratio of dosemeter reading with zero chamber wall thickness to the reading at normal wall thickness. Krieger and Sauer [15] measured doses in Styrofoam (with mass density of 0.035 g cm23) for variable chamber wall thicknesses made of plastic (PMMA) and at different field sizes, using a PTW 31003 ionization chamber which has a PMMA wall and wall thickness of 0.0655 cm (or 0.078 g cm22) and mass density of 1.19 g cm23. From an exponential curve fit to this data, dose values were calculated at zero wall thickness. For the purpose of our work, we assumed that the atomic compositions of graphite and PMMA are similar, that there is a linear relationship for the perturbation correction with medium density and that differences in the geometric characteristics of the chambers have a small influence on the perturbation in low density medium. The dose values in Styrofoam measured using a chamber with a PMMA wall were converted to dose values in Styrofoam measured with a chamber with a graphite wall. The replacement perturbation factor for the NE2571 in lung was further derived from linear interpolation between replacement perturbation factors for this chamber in Styrofoam and in water (prepl(A) in water is unity). Using this replacement perturbation correction, ionization 2571 readings (Mlung ðAÞ) measured for a field size A with the NE2571 in a homogeneous lung medium, were converted to dose to water in an infinitesimally small volume of water within the low density lung using:

2571 D2571 heterogeneous (A)~Dmeas,heterogeneous (A)z

2571 2571 D2571 homogeneous-lung ðAÞ~Mlung,repl ðAÞ|Nwater,6MV 2571 |Mhomogeneous-lung ðAÞ

ð1Þ

In the case of a measurement in a heterogeneous medium comprising lung and water where the chamber 830

1{

1 p2571 lung,repl (Aeq )

!

ð2Þ |D2571 (A eq ) homogeneous-lung

Thus for the derivation of the perturbation correction in the heterogeneous geometry, the dose in a homogeneous lung medium under full scatter conditions was also measured.

Treatment planning study Standard breast planning technique at NCCT All patients at NCCT undergoing external beam radiotherapy treatment planning for the breast have a three-dimensional (3D) planning study based on images from helical CT scanning using a 5 mm slice reconstruction. Patients are immobilized in a reproducible position, lying supine on a PosiboardTM-2 breast backrest (manufactured by Sinmed BV) and with couch position registration. Both arms are positioned above the patient’s head, facilitating the tangential field arrangements of the treatment plan. Ball bearings are placed at stable points of the patient to define a CT reference point to aid positional verification during treatment delivery. A planning CT scan incorporating these markers is acquired with a 0.5 cm slice separation and this dataset is transferred to Helax-TMS. The patient external contour (skin) and organs at risk are outlined, namely involved lungs and heart, if appropriate. The isocentric breast planning technique at NCCT treats the breast with two tangential fields and, in some cases, a single supraclavicular field. Non-divergent beam edges are produced at the match plane with the supraclavicular field and at the posterior edge of the tangential fields [16–18]. The plan normalization/prescription point is always set at the isocentre. The generation of a dose distribution within Helax-TMS is a two stage process. A dose plan is generated using fast interactive optimization of open and wedged beams. Interactive dose calculations (in the Beam Modelling module on Helax-TMS) are possible only with the PB algorithm. A final dose calculation is generated once the plan is placed for non-interactive computation (in the Evaluation module). Non-interactive dose computations are also possible with the CC algorithm, but this algorithm cannot be employed during manual optimization. Although a full 3D image set is used in the computation of dose distributions, current clinical practice at NCCT for breast planning dictates that The British Journal of Radiology, October 2006


optimization and evaluation of treatment plans are carried out only on the slice containing the normalization point.

Definition of the planning target volume (PTV) It is not standard practice at NCCT to define a planning target volume (PTV) for planning of the breast or chest wall. For the purpose of this study, however, a PTV was defined to enable the comparison of dose– volume histograms (DVHs) from calculations employing different calculation algorithms. The PTV was defined by the geometrical beam limits of the superior and inferior field borders and the pleura, and drawn 5 mm inside the external outline. With Helax-TMS, a structure cannot be defined once a beam geometry has been added to the study dataset. In order to define a PTV retrospectively and for the purpose of this work, each patient study set (images and structures) was exported in DICOM RT format to the Exomio (Version 2.0) virtual simulation software. PTVs were added on Exomio and the patient structure set was re-imported into Helax-TMS. Dose calculations on Helax-TMS All six patient studies were planned with a 6 MV beam using the standard breast planning technique. The PTV for each patient was drawn based on the criteria described previously. Dose calculations were carried out with both algorithms, both with and without inhomogeneity correction. The option of no-inhomogeneity correction simply assigns all tissue in the patient geometry to unit mass density. This was carried out in order to exclude the influence of the lung tissue on the dose calculations. The dose calculation grid was 0.5 cm. Evaluation of plans For the evaluation of treatment plans, isodose distributions, monitor units (to deliver 1 Gy at the normalization point) and data from DVHs were compared. To aid comparisons and relate the changes observed for the individual chest wall geometry and to current clinical practice, geometric measurements were taken from all CT studies on the central planning slice containing the normalization point. These measurements, as shown in Figure 2 were: the distance from the normalization point to the skin surface dskin-norm point, the thickness of breast, i.e. the distance from the lung-tissue interface to the skin surface dbreast, and the maximum thickness of lung tissue involved in this central slice dmax lung. For three patients, the normalization point was in breast tissue (patients 1, 2, 3) and for the others, in the lung (patients 4, 5, 6). The differences in monitor units from open and wedged beams for the PB and CC algorithms were analysed and the minimum, maximum and mean percentage doses in the PTV and lung were recorded from the DVH data.

Results Experimental verification The CC algorithm modelled the dose at the isocentre in a homogeneous unit density medium under tangential irradiation (Figure 1a) with varying lateral thickness of The British Journal of Radiology, October 2006

Figure 2. Distances measured on the central planning slice containing the isocentre. dskin-norm point is the distance from the normalization point (isocentre) to the skin of the patient, dbreast is the thickness on the chest wall at the level of the isocentre and dmax lungthe maximum lung involvement in the field.

tissue more accurately than PB. The uncertainties in the measured values did not exceed 0.5%. Deviations from measurement for CC did not exceed ¡1%, whereas for PB and for lateral thickness equal to or less than 2.5 cm, these were greater than +1% (maximum of +2.3%). These results were consistent with previous findings and generally the differences between the two algorithms for this beam energy are not large [9]. Figure 3 shows the variation of prepl with field size for the NE2571 in lung at 6 MV for collimator settings greater than 5 cm65 cm (the use of this chamber for dosimetry in smaller field sizes is not appropriate due to its size in relation to the size of the beam). The uncertainty in the derivation of prepl, as quoted by Krieger and Sauer, is 2.3% at 5 cm65 cm and is reduced to 1.5% at 15 cm615 cm [15]. In the case of a homogeneous lung medium irradiated with a 10 cm610 cm field under full scatter conditions, it was found that the CC model predicted the dose to within 0.9% of the measurement, whereas the PB overestimated dose by +8.8%. The combined uncertainty in the dose measurements did not exceed 2.1%. In Figure 4a–c, it can be seen how the measured dose varies with distance from lung tissue and phantom boundary, when the chamber is in unit density medium (experiment shown in Figure 1b). As the distance from the phantom boundary decreases, the dose decreases and it decreases even more (by 1.6%) once its proximity to the low density inhomogeneity is equal to or less than 1.5 cm. In this case, the overall deviation from measurement for CC is ¡0.6%, where errors by PB vary from +2.5% to +7%. For tangential irradiation with the normalization point in lung, the combined uncertainties in the measurement are expected to be of the order of 3% due to the assumptions in the derivation of the perturbation 831


Figure 3. Perturbation correction factors for the NE2571 ionization chamber (with its graphite cap) in lung and styrofoam. The data from Krieger and Sauer [15] for the PTW 31003 chamber with its PMMA cap in Styrofoam are shown for comparison (with permission). ¤, PTW31003 with PMMA wall in styrofoam (Krieger & Sauer, with permission); &, NE2571 with graphite wall in styrofoam; , NE2571 with graphite in wall in lung.

N

correction [15]. The dose predicted by the PB model varied between +10.7% to +8.2%, depending on the effective lateral thickness from the calculation point (Figure 4d). The CC predicted values between +3.6% and +2.3% for increasing effective thickness. This is in agreement with other workers who, in heterogeneous media and within low density and close to an interface, have reported the CC model to be more accurate than PB

[3, 15, 19]. Here, the CC model predicted a higher dose than the measured dose, but one has to consider the uncertainty in the derivation of the perturbation correction for the measurement in lung. In the case of the heterogeneous medium with the point of measurement in lung and under complex missing tissue irradiation geometry, the CC predicts dose values close to measurement and within 3% in the case of lateral

Figure 4. Dose in water with varying distance from phantom edge and lung. (a) 10 mm; (b) 15 mm (c) 20 mm. (d) Dose in lung with varying distance from phantom edge. &, pencil beam; m, collapsed cone; , measurement.

N

832



Figure 5. Typical dose distributions on the central planning slice generated by (a) the pencil beam (PB) and (b) the collapsed cone (CC) dose algorithms. These data are for patient 5 of this study.

The analogous results for the calculations without an inhomogeneity correction are shown in Table 3. Figure 6 illustrates how the percentage differences in mean doses to PTV and lung calculated by PB and CC vary with breast tissue thickness, maximum lung thickness and distance from the normalization point to the skin. As seen from Table 2, the CC model predicted higher MU than the PB, the latter overestimating dose, and this was consistent with our experimental findings. It can be seen (Figure 6) that the greater the thickness of breast tissue the smaller the differences between the two algorithms. For patient 3 a thickness of 4.4 cm of breast tissue, resulted in, on average, 3 MU difference per beam between algorithms, a 4.1% difference to the dose at the isocentre from open fields and a 0.4% difference in the mean PTV dose. For patient 5 the thickness of breast tissue was 1.0 cm and this resulted in 5 MU difference between algorithms, thus 16.5% difference to the dose contributions at the isocentre from the open beams and 11% difference at the mean dose to the PTV. For patients with the normalization point placed in breast tissue (patients 1, 2 and 3) the differences between the two algorithms were the lowest.

electronic equilibrium (effective lateral thickness greater than the 15 mm).

Treatment planning study Figure 5 shows isodoses for patient 5 of this study on the CT slice containing the normalization point within lung. The isodose distribution calculated with CC follows the lung contour, predicting an overall lower dose to lung tissue and a relatively higher dose in tissue. This is because the energy released in the lung is scattered laterally to the adjacent breast and lung tissue and PB does not model this. Table 1 summarizes the patients included in this study and their measurements. The effective lateral distance from the normalization point to the skin is defined as the distance of tissue scaled with mass density. Table 2 summarizes the results from the two algorithms in terms of absolute difference in MU per beam, the average percentage difference in MU for the open fields only and the percentage difference in minimum, maximum and mean doses to the PTV and lung extracted from DVHs.

Table 1. Details and geometric measurements for patients included in this study Patient

Chest wall/ isocentre location

Breast tissue depth (cm): dbreast

Maximum lung depth (cm): dmax lung

Normalization point to skin (cm): dskin-norm point

Effective distance from normalization point to skin (g cm23). For the normalization point in lung: dbreast + (dskin-norm point2dbreast)60.26

1 2 3 4 5 6

Left/breast Left/breast Left/breast Right/lung Right/lung Right/lung

1.9 2.6 4.4 1.1 1.0 1.5

1.8 2.0 2.5 2.4 2.8 2.3

1.3 2.1 2.5 1.2 1.8 1.7

1.3 2.1 2.5 1.1 1.2 1.6


833

M M Aspradakis, H M McCallum and N Wilson Table 2. The difference in monitor units calculated between a pencil beam (PB) and collapsed cone (CC), their average percentage differences for the open fields and the percentage difference in dose to the planning target volume (PTV) and lung, with the inhomogeneity correction switched on Patient

1

Beam modulation

Open Wedged Open Wedged Open Wedged Open Wedged Open Wedged Open Wedged

2 3 4 5 6

Difference in MU PB-CC Medial field

Lateral field

22.8 20.7 23.5 – 22.4 20.4 27.4 21.5 29.1 – 25.8 2

22.9 20.4 23.3 – 23.0 – 27.4 21.5 28.9 – 24.8 21.0

% average difference in MU for open fields ((CC2PB)/PB)6100

% Difference in dose from DVH ((CC2PB)/PB)6100 PTV

Lung

Min.

Max.

Mean

+5.3

27.4

1.8

+1.4

+5.2

26.4

0.4

+4.1

1.3

Min.

Max.

Mean

0.0

22.8

22.2

+1.4

16.7

26.2

0

0.9

+0.4

23.1

28.6

+1.9

+16.5

14.7 10.4

+11.2

214.3

3.6

+10.7

+16.5

10.9

9.3

+11.0

222.2

6.1

+5.8

+9.6

2.6

1.5

+4.5

8.3

0.8

+1

DVH, dose–volume histogram.

Figure 6 shows that with increasing lung involvement in the plan, the mean dose to the PTV and lung increases. Patients 4 and 5 had maximum lung involvement of 2.4 cm and 2.8 cm, and for these the average percentage difference in MU was the highest, 16.5% and 16.6%, respectively. For patient 3, with similar lung involvement (2.5 cm), an analogous result was expected. But for this patient the normalization point was within 4.4 cm of breast tissue and the percentage difference in MU and the difference to the mean dose in the PTV between algorithms was the lowest observed. This would indicate that when there is sufficient breast tissue laterally, the influence of lung on the dose at the normalization point is reduced. With the inhomogeneity correction switched off, the differences in MU between PB and CC were smaller (Table 3). For patient 5, for example, the difference in MU was 2 (3%) when the lung was not accounted for, as opposed to 9 (16.6%) when accounted for in the calculation. This confirms that it is important to account Table 3. The difference in monitor units calculated between pencil beam (PB) and collapsed cone (CC) and their average percentage differences for the open fields with the inhomogeneity correction switched off Patient Beam Difference in MU modulation PB-CC

1

Open Wedged Open Wedged Open Wedged Open Wedged Open Wedged Open Wedged

2 3 4 5 6

834

Medial field

Lateral field

22.1 20.3 22 – 21.6 – 21 20.2 22 – 22 –

22.1 20.4 22 – 21.8 – 21 20.2 21.6 – 21.7 20.2

% average difference in MU for open fields ((CC2PB)/PB)6100

+3.9 +3.0 +2.7 +2.1 +3.0 +3.3

for the presence of lung in the dose calculation and when a lung heterogeneity is present, differences between algorithms increase with increasing involvement of lung. In Table 2 it is also seen that the minimum dose difference to the PTV between the algorithms was small in some cases and in particular for patient 3. As pointed out previously, the optimization of these plans was carried out based on the distribution calculated on the central slice of the study and this is why the results from DVHs are not optimal in terms of the minimum dose to the PTV. In addition to this, these PTVs were drawn close to the patient skin and close to the penumbral region of both fields that would contribute to small dose values.

Discussion and conclusions The dose at the normalization point in the case of isocentric breast irradiation using a 6 MV beam, with no lung involvement and in a region of lateral electronic equilibrium, is modelled to within +3% by the PB algorithm and within ¡1% by CC. For breast cancer patients who have undergone mastectomy and are also to be treated with external beam radiotherapy, the presence of lung tissue and often the small thickness of the chest wall present in the treatment volume (less than 2.5 cm), is a concern due to the known limitations of dose calculation algorithms in modelling dose in lung, and at lung-tissue interfaces, due to the missing tissue geometry and due to patient movement during treatment delivery. In terms of treatment planning, the CC algorithm calculates dose more accurately under most of these conditions (to within ¡1%), with the exception of positions very close to media interfaces. The PB algorithm can underestimate the required MU by up to 7% in some cases and in all cases generates a misleading dose distribution, overestimating the dose in lung and underestimating the dose to the PTV. The differences in MU between PB and CC depend on the thickness of chest wall and the position of the normalization point with respect to the The British Journal of Radiology, October 2006


Figure 6. Percentage differences in mean dose to the planning target volume (PTV) and lung between pencil beam (PB) and collapsed cone (CC) for patients included in this study plotted in terms of: (a) thickness of chest wall, (b) maximum thickness of lung on the central slice and (c) the distance from the normalization point to the skin. m, mean dose to PTV; & mean dose to lung.

lung. Our findings indicate that for a chest wall thickness equal or less than 2.5 cm and with 2 cm or more lung tissue involvement, the CC algorithm should be used instead of PB. For cases when the plan normalization point has to be placed in the lung and close to an interface with breast tissue, one needs to be aware of the limitations of the TPS algorithms in such regions. It is preferable to avoid placement of the plan normalization point close to interfaces, but as seen in some of the clinical cases examined here, the thin chest wall often leaves no other option. In our study, in some cases the breast tissue received higher than the prescribed dose (at least 5% higher) and the clinician should be advised to prescribe to the clinically significant isodose level covering the chest wall tissue. This work investigated the dosimetric accuracy of isocentric chest wall irradiation using 6 MV photon beams, which is the energy used for such irradiations at our centre. A previous comparison between the PB and CC algorithms with the plan normalization point (isocentre) in homogeneous water under a tangential irradiation geometry and for a range of beam energies (4 MV, 6 MV, 15 MV) has shown that the difference from measurement is greater at the lowest energy and in The British Journal of Radiology, October 2006

particular for the PB algorithm. At higher beam energies, for a tangential irradiation with the normalization point in low density medium, it is expected that the performance of both calculation algorithms against measurement could worsen, because at higher energies the range of travel of secondary particles is greater and modelling the deposition of their energy becomes more complex, especially in the lateral direction and close to media interfaces. Routine clinical implementation of the CC dose calculation algorithm on Helax-TMS is hindered because it is not possible to implement this interactively. However, in the special case of chest wall irradiation it is advised that an optimum treatment plan is produced using the PB algorithm and the same plan is also calculated with the CC algorithm for the clinician to make an informed decision on the prescription to the chest wall tissue. At NCCT the Helax-TMS system will soon be replaced with the Oncentra (OTP) TPS (Nucletron B.V.). Both PB and CC dose calculation engines are available on OTP, and CC could be used interactively. Therefore, CC would be the algorithm of choice for treatment planning of the chest wall, and other sites with significant involvement of non-unit density heterogeneity. 835


The isocentric breast planning technique at NCCT will develop further with the optimization and evaluation of plans based on distributions generated on all CT slices. New approaches have been investigated for plan optimization for the breast, such as electronic compensation using field-in-field (forward IMRT) [20, 21]. For these to be implemented clinically it is required that a PTV is defined on all slices of the CT study. This necessitates additional input by the clinician in treatment planning of the breast and/or the revision of guidelines for breast planning to be followed by experienced dosimetrists. New, complex treatment planning techniques for breast planning also necessitate a revision of the methods used for the independent check of monitor units generated by TPSs. Patient movement during treatment has not been addressed here. Future efforts in improving the treatment of chest wall patients would also need to account for chest wall movement [22, 23].

Acknowledgments The authors are thankful to Dr Otto Sauer (Department of Radiotherapy, Julius-Maximilians-University of Wu¨rzburg, Germany) for discussions on the derivation of chamber perturbation correction for measurements in lung. Mr Geoff Lambert (Head of Radiotherapy Physics, Regional Medical Physics Department, Newcastle upon Tyne) and Ms Gill Lawrence (Consultant Clinical Scientist in Radiotherapy Physics, Regional Medical Physics Department) are gratefully acknowledged for their support and helpful comments to the manuscript. Dr Ujjal Mallick (Consultant Clinical Oncologist, Northern Centre for Cancer Treatment) is acknowledged for his clinical input to this work.

References 1. Swerdlow A, Dos Santos Silva I, Doll R, editors. Cancer incidence and mortality in England and Wales: trends and risk factors. Oxford University Press, 2001:71–196. 2. Johnson S. Incidence and mortality in breast cancer. In: International handbook of breast cancer. Euromed Publications, 2002. 3. Ahnesjo¨ A. Collapsed cone convolution of radiant energy for photon dose calculation in heterogeneous media. Med Phys 1989;16:577–92. 4. Ahnesjo¨ A, Saxner M, Trepp A. A pencil beam model for photon dose calculation. Med Phys 1992;19:263–73. 5. Ahnesjo¨ A, Knoö¨s T, Montelius A. Application of the convolution method for calculation of output factors for therapy photon beams. Med Phys 1992;19:295–301. 6. Knoö¨s T, Ceberg C, Weber L, Nilsson P. The dosimetric verification of a pencil beam based treatment planning system. Phys Med Biol 1994;39:1609–28.

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7. Engelsman M, Damen EMF, Kiken PW, van’t Veld A, Ingen KM, Mijnheer BJ. Impact of simple tissue inhomogeneity correction algorithms on conformal radiotherapy of lung tumours. Radiother Oncol 2001;60:299–309. 8. Weber L, Nilsson P. Verification of dose calculations with a clinical treatment planning system based on a point kernel dose engine. J Applied Clin Med Phys 2002;3:73–87. 9. Aspradakis MM, Morrison RH, Richmond ND, Steele A. Experimental verification of convolution/superposition photon dose calculations for radiotherapy treatment planning. Phys Med Biol 2003;48:2873–93. 10. Wieslander E, Knoö¨s T. Dose pertubation in the presence of metallic implants: treatment planning system versus Monte Carlo simulations. Phys Med Biol 2003;48:3295–305. 11. Irvine C, Morgan A, Crellin A, Nisbet A, Beange I. The clinical implications of the collapsed cone planning algorithm. Clin Oncol 2004;16:148–54. 12. Constantinou C, Attix FH, Bhudatt RP. A solid water phantom material for radiotherapy x-ray and gamma-ray beam calibrations. Med Phys 1982;9:436–41. 13. White DR, Martin RJ, Darlison R. Epoxy resin based tissue substitutes. Br J Radiol 1977;50:814–21. 14. White DR, Constantinou C, Martin RJ. Technical Note: Foamed epoxy resin-based lung substitutes. Br J Radiol 1986;59:787–90. 15. Krieger T, Sauer O. Monte Carlo- versus pencil beam-/ collapsed -cone-dose calculation in a heterogeneous multilayer medium. Phys Med Biol 2005;50:859–68. 16. Siddon RL. Solution to treatment planning problems using coordinate transformations. Med Phys 1981;8:766–74. 17. Casebow MP. Matching of adjacent radiation beams for isocentric radiotherapy. Br J Radiol 1984;57:735–40. 18. Wilkinson J. Consensus management - a departmental study of the external beam therapy of carcinoma breast. Radiography Today 1992:17–20. 19. Carrasco P, Jornet N, Duch MA, Weber L, Ginjaume M, Eudaldo T, et al. Comparison of dose calculation algorithms in phantoms with lung equivalent heterogeneities under conditions of lateral electronic disequilibrium. Med Phys 2004;31:2899–911. 20. Richmond ND, Turner R, Dawes PDK, Lambert GD, Lawrence GL. Evaluation of the dosimetric consequences of adding a single asymmetric or MLC shaped field to a tangential breast radiotherapy technique. Radiother Oncol 2003;67:165–70. 21. Wilkinson M, Aspradakis MM. A study of electronic compensation methods in breast radiotherapy planning. In: IPEM Annual Scientific Meeting. York, UK: IPEM, 2004. 22. Pedersen AN, Korreman S, Nystrom H, Specht L. Breathing adapted radiotherapy of breast cancer: reduction of cardiac and pulmonary doses using voluntary inspiration breathhold. Radiother Oncol 2004;72:53–60. 23. Korreman SS, Pedersen AN, Nottrup TJ, Specht L, Nystrom H. Breathing adapted radiotherapy for breast cancer: comparison of free breathing gating with the breath-hold technique. Radiother Oncol 2005;76:311–8.



Patient and staff radiation doses from early radiological examinations (189921902) 1

C J KOTRE,

PhD

and 2B G LITTLE,

MSc

1

Regional Medical Physics Department, Newcastle General Hospital, Newcastle-upon-Tyne NE4 6BE and 2Radiotherapy Physics Department, North Glasgow University Hospitals, Dumbarton Road, Glasgow G11 6NT, UK ABSTRACT. A source of data on radiographic and fluoroscopic examinations, including radiographic technique factors, was used in conjunction with information about coldcathode X-ray apparatus to estimate patient and staff radiation doses for the years 1899 to 1902 at the Forth Banks Infirmary, Newcastle-upon-Tyne. Physical evidence from representative apparatus of the period was used with a beam spectral simulation program to characterize the X-ray beam, and information about the electrical supply waveform was produced by experimental operation of a contemporary induction coil. Results are given in terms of skin entrance dose, and these are compared with modern values. An estimate of the annual dose received by the radiographer known to have carried out all of the examinations within this period is also given.

A collection of historical artefacts associated with early radiology in the north-east of England is on display in the Radiology Department of the Royal Victoria Infirmary, Newcastle-upon-Tyne. This collection, assembled by former Newcastle radiologist Dr C K Warrick, includes a number of very early cold-cathode Xray tubes and an example of the type of induction coil high voltage apparatus which would have been used to power them. Also included is an original log-book listing the radiological examinations carried out at the Forth Banks Infirmary (later to become the Royal Victoria Infirmary) in the years 1899 to 1902. The log-book also contains entries on the exposure time in minutes and the distance used for the radiographic examinations (69% of the total), or denotes that the examination was carried out by fluoroscopy. Due to the coincidence of both original documentary and physical evidence relating to early radiographic practice in this collection, it was decided to use it to attempt an estimation of patient and staff radiation doses for comparison with modern values.

Received 3 January 2006 Revised 22 February 2006 Accepted 8 March 2006 DOI: 10.1259/bjr/16982267 ’ 2006 The British Institute of Radiology

taken as sodium glass as lead glass would have produced a blue glow [1]. The examination log quotes distances from the tube to the patient entrance surface for each exposure. For the calculations below, these are taken as distances from the outside of the tube envelope as some of the distances were very short and it is thought unlikely than the focal spot would be taken as the source of the radiation in routine work of the time. The distance from the centre of the target to the outside of the tube envelope was measured at 3 cm.

Induction coil The induction coil examined for this work (Figure 2) is also dated c. 1899 and was used in conjunction with coldcathode X-ray tubes at a pharmacist’s establishment in Sunderland. It was actually used with the X-ray tube detailed above, but not at the Forth Banks Infirmary. The

Evidence and assumptions X-ray tube A cold-cathode X-ray tube from around 1899 (Figure 1) was used to formulate the specification of the ‘‘typical’’ X-ray tube which would have been used at the Forth Banks Infirmary during the 1899–1902 period under study. The anode material was taken to be platinum, as tungsten anodes were not yet in use [1]. The target angle was measured at 35 ˚ and the thickness of the glass envelope estimated at 2 mm. From the evidence of Warrick [2], who quotes an early Newcastle radiologist remembering the ‘‘beautiful apple green fluorescence which always pleased me’’ given off by this type of equipment in operation, the glass type can be The British Journal of Radiology, October 2006

Figure 1. The cold-cathode X-ray tube of c. 1899 used to formulate the specification for the X-ray tube used in the simulations.

837

C J Kotre and B G Little

Figure 2. Induction coil of c. 1899 under test with a lowvoltage power supply.

coil is 15 cm in diameter and 30 cm in length, and stands on a hollow wooden plinth supporting the sprung contact-breaker arrangement and containing a capacitor composed of flat foil electrodes insulated with waxed paper. Despite being the size of a telephone book, this capacitor measured only 1 mF. The capacitor insulation was found to be intact and some evidence of re-wiring was found inside the plinth suggesting that restoration, possibly for teaching purposes, had taken place at some stage. The primary winding was found to be functional, but not the very fine secondary winding. Breakdown of the secondary winding insulation was also suspected. Various methods of deriving the probable tube potential supplied by the induction coil were investigated, but rejected due to the large number of unknowns related to the tube loading and current/voltage relationships of cold-cathode tubes at various states of evacuation. It was also suspected that the induction coil used at the Forth Banks Infirmary was rather larger than the Table 1. Classifications and uses of cold-cathode X-ray tubes from an early guide [3] together with assumed tube potentials for these examinations based on more modern experience [4] Tube type

Description

Applications

Assumed tube potential (kVp)

1

‘‘Soft’’

50

2

‘‘Medium’’

3

‘‘Medium–Hard’’

Tendons Limbs in children Epiphyses in very young subjects Less dense structures Thorax Renal calculi in thin patients Limbs in adults Fluoroscopy Thorax Renal calculi in stout patients Spine Pelvis Fluoroscopy

838

60

70

example to hand, so documentary evidence was relied upon. Warrick [2], quoting a letter from an early Newcastle radiologist, mentions that, ‘‘The coil was a large one of conventional pattern with a platinum make and break...I think the spark gap was ten inches’’. The potential corresponding to this 10-inch spark gap depends on the shape of the spark gap electrodes, being lowest for point electrodes and much higher for spherical electrodes. A 10-inch spark gap corresponds to approximately 120 kV for point electrodes, which seems rather high, although it is possible that the 10-inch recollection refers just to the size of the induction coil itself as the length of the coil was often quoted in advertising literature. In early radiography, part of the skill of the radiographer was selection of the best cold-cathode X-ray tube for the examination. The tubes had a short lifetime, and the hardness of the vacuum increased during use. Younger ‘‘soft’’ tubes were therefore preferred for examinations requiring less penetrating radiation and older ‘‘hard’’ tubes were selected for large body sections. An early guide to tube selection, quoted by Burrows [3], uses three classifications of tube and gives examples of their use. Table 1 gives these three classifications, together with an estimate of the required kVp based on more modern experience [4]. These latter values of kVp are assumed for the tube types used at the Forth Banks Infirmary in 1899–1902. The values lie at the low end of the modern diagnostic kVp range, but this is consistent with the finding reported in a number of the early X-ray examinations that the image was poor and underpenetrated for thicker body sections. The use of a mechanical contact-breaker with an induction coil produces an intermittent high voltage waveform, and some estimate of this waveform was needed. After some hesitation, the authors gave in to their curiosity and experimentally operated the 1899 induction coil from a high current capacity variable low voltage power supply. The supply voltage was increased until the contact-breaker operated with a regular, reasonably even frequency. Despite the discontinuity of the secondary winding and suspected breaks in its insulation, high voltages were produced across a chain of large resistors and a short spark-gap. Oscilloscope traces of the primary voltage waveform showed a very uneven pattern due to the arcing of the contact-breaker, but traces of the secondary high-voltage waveform showed a regular train of pulses (Figure 3) from which the beam-on time was estimated at 9.6% of the recorded exposure time. It was assumed that this figure would be typical for apparatus of the time. The final essential electrical parameter was the tube current. The unknown electrical characteristics of a typical cold-cathode X-ray tube made this also difficult to estimate from a physical basis. Burrows [3] gives values of 1–2 mA for early cold cathode tubes, noting that thin sheet platinum targets would melt at the higher values of current which later became available as heavier induction coils were developed. Reynolds [4] (who himself designed a modified cold cathode X-ray tube in 1900) gives values of 2–5 mA. For this early tube and induction coil, a compromise value of 2 mA was assumed for the Type I tube (50 kVp), with the current scaled down to a constant power output from the The British Journal of Radiology, October 2006

Radiation doses from early radiological examinations

Figure 3. Oscilloscope trace of the high-voltage secondary waveform obtained with the induction coil in operation.

induction coil for the harder tubes, giving 1.7 mA for the Type 2 tube (60 kVp) and 1.4 mA for the Type 3 tube (70 kVp). The specification of the X-ray generating apparatus arrived at from the above considerations is summarized in Table 2.

Calculation of patient entrance skin doses A useful quantity to estimate for comparison with modern day dose levels is the entrance surface dose (ESD). The first step in this calculation was to simulate the output spectrum from the tube used over the time period noted in the log-book. For this, a commercially available software package was used [6]. The input parameters required are peak tube potential, target angle, target material and filtration in the beam. The intensity of photons at a given energy is calculated for Bremsstrahlung and characteristic radiation via empirical relationships. For a description of the theory, the Table 2. Summary of quantities used to specify the X-ray tube and operating conditions Quantity

Value measured (M) or assumed from indirect evidence (A)

Anode material Anode angle Envelope material Envelope thickness Tube potential

Platinum (A) [1] 35 ˚ (M) Sodium glass (A) [1, 2] 2 mm (M) Type 1: 50 kVp (A) [3, 4] Type 2: 60 kVp (A) [3, 4] Type 3: 70 kVp (A) [3, 4] Type 1: 2.0 mA when beam on (A) [3, 5] Type 2: 1.7 mA when beam on (A) [3, 5] Type 3: 1.4 mA when beam on (A) [3, 5] 9.6% (M)

Tube current

Duty cycle (% of time beam on) Distance from focus to exterior of tube

3 cm (M)


reader is referred to Birch and Marshall [7]. The software limits the input parameters to those appropriate to modern day X-ray tubes, therefore some of the parameters from the 1899 X-ray tube could not be simulated directly. The parameters used were as follows: tungsten target, 22 ˚ target angle, 50 kVp, 60 kVp and 70 kVp for tube types 1, 2 and 3, respectively (Table 2), voltage ripple of 0% (for the actual beam on time) and beam filtration of 2 mm of glass. Where the parameters used were not that assumed from Table 2, various corrections were made and this is discussed later in more depth. The logged examination types were classified into groups thought to match up with tube types 1, 2 and 3 based on the applications given in Table 1. The simulation software produced values of air kerma per mAs at 75 cm from the target, which were then converted to ESDs for the historical examinations by multiplying them by the recorded time of exposure, tube current, backscatter factor and taking account of the actual distance from the tube. As the distances given are generally quite short (down to 3 cm) it is assumed that these are measured from the surface of the glass envelope rather than from what we would now call the focal spot. An extra 3 cm was therefore added to account for the distance from the focus to the outside of the tube envelope. A realistic backscatter factor of 1.16 was determined from tabulated data of backscatter factor for a glass window tube with no external filtration at a potential of 60 kV [8]. An output correction between the maximum target angle in the simulation package, 22 ˚, and that measured for the cold-cathode tube target, 35 ˚, was made using a simple model of the self attenuation of the anode [7] by fitting parameters to this model from the relationship between output and target angle produced by the software package at 50 kVp, 60 kVp and 70 kVp. Finally a correction was made for the difference in air kerma due to the difference in target materials. It is known that for Bremsstrahlung X-ray production the air kerma produced is proportional to the atomic number of the target material [9]. The atomic number of platinum is 78 as compared with 74 for tungsten and the ESD was therefore modified by the ratio of the atomic numbers. It was assumed that changes in the energy of the characteristic radiation made a negligible difference to the air kerma and no modification was made for this.

Calculation of annual staff dose The radiographer known to have carried out the examinations listed in the 1899–1902 record was Mr Thomas Dodd. When the Forth Banks Infirmary acquired its first X-ray equipment in 1899, Mr Dodd, the older brother of the head porter, was employed as technician to operate it and keep the log-book [3]. He worked at the old Infirmary and later the new Royal Victoria Infirmary for over 20 years. In the process, he suffered severe radiation injuries, losing three fingers (one of which is preserved in the historical collection) and ultimately dying as a result of his injuries in 1929 [10]. A later c. 1920 photograph of Mr Dodd in action [2, 3] shows him positioned between the patient and the high voltage generating equipment with the X-ray tube at his head height. It seems likely that this was his normal 839


working position. Although this later photograph shows what might be crude shielding around the tube, the early tubes were not shielded. The distance between his body centreline to the tube is estimated from the photograph to be 1.4 m. It is assumed that he was present throughout even the long exposures to monitor the induction coil operation and check for patient movement. In order to estimate Mr Dodd’s annual whole-body dose, the dose to a point at which a modern waist-level personal dosemeter would be worn was calculated. This point is taken as 1.6 m from the tube focus and 1.4 m from the beam entry point on the patient. For simplicity, the tube axis is taken to be parallel to the table with Mr Dodd alongside. No correction for the possible shielding effect of the anode was therefore applied. In order to estimate the annual staff dose, the radiographic exposure times and number of fluoroscopic examinations were totalled over the 2.8 years recorded in the log-book. The exposure time for the fluoroscopic exposures was estimated at 5 min per examination, to include the setting up of the tube and recording of the image by sketching. This latter practice was used in the 1920s when more technicians were employed in the department [3] and it is assumed here that a similar method was used in the early days when Mr Dodd worked alone. Using the model of the X-ray apparatus described above, the annual dose to a waist-level personal dosemeter, including backscatter, was estimated. Tube type 2 was assumed for all fluoroscopic examinations, and the same classification of tube type as used above was employed for the radiographic examinations. The scatter contribution was estimated using modern dose–area product to scatter dose conversion factors [11] and assuming (in the absence of any field limiting device) a scattering patient surface of 25 cm625 cm. For the fluoroscopic examinations, any attenuating effect of the patient on the primary radiation was neglected on the grounds that the patient section would only partially shield the radiographer from the uncollimated primary beam and unshielded tube.

Results and discussion Breakdown of examination types The frequency of examination types (as named in the record) is shown in Table 3 for the radiographic examinations and Table 4 for the fluoroscopic examinations. Over the 2.8 year time period in the log book there were 507 examinations noted. The top three examinations overall (for both fluoroscopy and radiography examinations) were investigation of fracture (or bone disorder), foreign bodies and renal calculi with proportions of 50%, 35% and 8%, respectively. Some patterns within this group are also of interest. Of the 174 foreign body examinations, the most frequently occurring (50%) were cases of needle and pin injuries including needle in throat, hands, hat-pin in stomach, etc. A further significant proportion of the foreign body examinations (18%) involved bullet injuries to subjects between the ages of 12 years and 26 years, with one-third of these being described as ‘‘bullet in hand’’. 840

Table 3. Breakdown of examination types for radiography Radiographic examination type

No. in record

Fracture Foreign body Renal calculus Pulmonary tuberculosis Sarcoma Aneurysm of aorta Ricketts Others Total radiographic

189 83 38 8 4 3 3 12 340

Table 4. Breakdown of examination types for fluoroscopy Fluoroscopic examination type

No. in record

Foreign body Fracture Aneurysm of aorta Sarcoma Others Total fluoroscopic

101 53 6 1 6 167

Patient ESDs compared with modern values Due to the high frequency and variety of peripheral and foreign body examinations in the historical data set, it is quite difficult to make meaningful comparisons with modern ESD data for the majority of the examinations. Three sub-groups were, however, extracted for comparison. These were adult abdomen/pelvis examinations (40 exposures), adult non-penetrated chest radiographs for diagnosis of tuberculosis and sarcoma of rib (10 exposures) and adult penetrated/lateral chest examinations associated with investigations of foreign bodies (14 exposures). The distributions of ESD for these examinations are shown in Figure 4 (abdomen/pelvis), Figure 5 (chest) and Figure 6 (penetrated/lateral chest). The median ESD values for these three sub-groups are 189 mGy for the abdomen/pelvis examinations, 68 mGy for the chest examinations and 354 mGy for the penetrated/lateral chest examinations. These can be compared with modern median ESD values [12] of 4.1 mGy for abdomen anteroposterior (AP), 3.2 mGy for pelvis AP, 0.14 mGy for chest AP views and 0.51 mGy for lateral chest. This gives a ratio approximately 52 times higher ESD for the abdomen/pelvis examinations, 486 times higher for the chest examinations and 694 times higher for the penetrated/lateral chest examinations. In order to cross-check the magnitude of these results and elucidate the dose implications of technique changes on the patient dose results, the simulation software [6] was further used to calculate the expected ratio of ESD between the historical and modern techniques [12] for these three examinations, normalized to a constant value of receptor dose. The differences in kVp, distances, filtration, grid use (abdomen/pelvis only), target angle and X-ray spectrum were considered. The effects of these technique differences result in an expected ratio of ESD historical/modern of 2.0 for the abdomen/pelvis, 5.2 for the AP chest examinations and 8.6 for the lateral/ penetrated chests. From the results above, this leaves a The British Journal of Radiology, October 2006

Radiation doses from early radiological examinations

ratio of 26 for the abdomen/pelvis, 93 for the (few) chest examinations and 80 for the lateral/penetrated chests to be accounted for by the improvements in radiographic speed offered by modern screen–film systems compared with photographic emulsion alone. These ratios are in the region of what would be expected from the intensification factor of modern screen–film systems (30–100 [13]). The highest radiographic ESD for a single exposure estimated for the whole log was 1.8 Gy for an exposure for ‘‘steel in neck’’ carried out at a focus–skin distance of 20 cm and requiring an exposure time of 20 min.

Annual staff dose

Figure 4. Histogram of entrance surface dose frequency for abdomen/pelvis examinations.

Figure 5. Histogram of entrance surface dose frequency for

The annual whole-body dose for Mr Dodd during the period of the record from both primary and scattered radiation is estimated to be 940 mSv year21 (97% from primary irradiation, 3% from scatter), a figure which is 47 times the annual dose limit for workers under current UK regulations. Mr Dodd was associated with radiography at the Forth Banks Infirmary, and later the Royal Victoria Infirmary which replaced it, for 20 years [3]. During this period, the need for radiation protection was becoming recognized, and the 1920s photograph of Dodd [2, 3] does appear to show protective plates either side of the X-ray tube, but equally it shows him working without any personal protective equipment. The workload of the Forth Banks and later Royal Victoria Infirmary Departments increased rapidly from the early level of approximately 200 examinations per year reported here, to almost 14 000 per year by 1918 [3]. Mr Dodd was the sole radiographer until around 1910 when an assistant radiographer, Cornelius McMeekin, was appointed [2, 3], and he had some assistance from two R.A.M.C privates during the war years [2], but with the documented workload expansion during his career it seems likely that Mr Dodd’s personal radiation exposure continued to be considerable. He died of cancer in December 1929 [10].

posteroanterior chest examinations.

Conclusions This work has attempted to use modern methods to estimate patient and staff radiation doses from X-ray examinations performed over a century ago. The actual values obtained show some inconsistencies and hang by a very long chain of stated assumptions, but physical evidence has been used where practicable to add weight to these assumptions. As would be expected, the results show patient doses of a magnitude unacceptable by modern standards, but not perhaps as high as might be guessed at merely by noting the very long exposure times in the original record. The dangerous level of staff irradiation for the earliest radiographers is quantitatively confirmed.

Acknowledgments Figure 6. Histogram of entrance surface dose frequency for lateral/penetrated chest examinations.


We would like to thank the Radiology Department, Royal Victoria Infirmary, Newcastle-upon-Tyne for 841


access to its historical collection, Dr C K Warrick for useful discussions on the origins of the material, and Mr Mike Feeney for discussions on early high-voltage generating equipment and assistance with the operation of the induction coil.

References 1. Mould RF. The early history of x-ray diagnosis with emphasis on the contributions of physics 189521915. Phys Med Biol 1995;40:1741–87. 2. Warrick CK. Notes on the history of the Department of Radiology of the Royal Victoria Infirmary, Newcastle upon Tyne. Radiography 1977;43:190–4. 3. Burrows EH. Pioneers and early years: a history of British radiology. Channel Islands: Colophon Ltd, 1986. 4. Goldman M, Cope D. A radiographic index, 4th edn. Bristol: John Wright and Sons, 1970. 5. Reynolds RJ. Sixty years of radiology. Br J Radiol 1956;29:238–45.

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6. IPEM. Catalogue of Diagnostic X-ray Spectra and other data, CDROM Report No. 78. York: Institute of Physics and Engineering in Medicine, 1997. 7. Birch R, Marshall M. Computation of Bremsstrahlung X-ray spectra and comparison with spectra measured with a Ge(Li) detector. Phys Med Biol 1979;24:505–17. 8. BJR Supplement 25. Central axis depth dose data for use in radiotherapy. London: British Institute of Radiology, 1996. 9. Dendy PP, Heaton B. Physics for diagnostic radiology, 2nd edn. Bristol and Philadelphia: IoP Publishing, 1999. 10. Obituary, Thomas Dodd. Br J Radiol 1930;3:95. 11. Sutton DG, Williams JR. Radiation shielding for diagnostic X-rays. London: British Institute of Radiology, 2000. 12. Hart D, Hillier MC, Wall BF. Doses to patients from medical X-ray examinations in the UK – 2000 Review. National Radiological Protection Board Report NRPB-W14. Didcot: Health Protection Agency, 2002. 13. Farr RF, Allisy-Roberts PJ. Physics for medical imaging. London: WB Saunders, 1998.



PICTORIAL REVIEW

Ultrasound spectrum in intraductal papillary neoplasms of breast S GANESAN,

MD,

G KARTHIK,

DNB,

M JOSHI,

MD, DNB

and V DAMODARAN,

MS, FRCS

Department of Radiology and Imaging, G.K.N.M Hospital and Research Centre, PN Palayam, Coimbatore – 641037, India ABSTRACT. Intraductal papillary neoplasms (IPNs) of breast form a wide spectrum of pathological changes with benign intraductal papilloma occupying one end of the spectrum and papillary carcinoma at the other end. Intraductal papillomas are known to occur anywhere within the ductal system and are broadly classified into central and peripheral types. Intraductal papillary carcinoma is an uncommon ductal malignancy forming papillary structures, and these lesions characteristically lack the myoepithelial layer present in benign papillary neoplasms. Three basic patterns of IPNs are recognized on ultrasound – intraductal mass with or without ductal dilatation, intracystic mass and a predominantly solid pattern with the intraductal mass totally filling the duct. Benign papillomas are known to exhibit calcifications which tend to be extremely dense and coarse. IPNs are highly vascular tumours and have a propensity to bleed spontaneously. A distinct vascular pedicle is identified within the central core of IPNs, with branching vessels arborising within the mass. In an older age group, presence of a large solid component and evidence of spontaneous intracystic bleed are more suggestive of papillary carcinomas than benign papillomas. We have serially studied 42 cases of intraductal papillary neoplasms with sonomammography and mammography from 2001 to 2004.

Intraductal papillomas are common neoplasms with a relative incidence of 2–3% [1]. In elderly patients, intraductal papillomas are often asymptomatic and are seen commonly as an incidental finding in biopsy specimens [2]. Even though intraductal papillomas are primarily benign, these lesions can pose problems in view of their similarity to intraductal papillary carcinoma clinically, on ultrasound and histologically [3]. Intraductal papillary carcinoma (IPC) is a rare ductal carcinoma forming papillary structures, with reported incidence of 1–4% of breast carcinomas [4]. These neoplasms have certain characteristic imaging features which help to differentiate these lesions from other focal breast abnormalities.

Pathological observations Papillomas are essentially benign proliferations of ductal epithelium. They may occur at any age between 30 years and 77 years [4], but are commonly seen between 30 years and 55 years [3]. They are known to occur anywhere within the ductal system and are broadly classified into central and peripheral types. Central types are usually solitary, while the peripherally located papillomas tend to be multiple within the terminal duct lobular unit. Lesions are often confined Address correspondence to: Dr Karthik Ganesan, Department of CT and MRI, Jaslok Hospital and Research Centre, 15, Dr G Deshmukh Marg, Mumbai – 400026, Maharashtra, India.


Received 24 November 2004 Revised 16 May 2005 Accepted 1 June 2005 DOI: 10.1259/bjr/69395941 ’ 2006 The British Institute of Radiology

to a single breast, while bilateral lesions are reported in up to 14% of cases [4]. Central papillomas are subareolar in location within a major duct. On macroscopic examination, a papilloma appears as a round to oval, small mass usually measuring a few millimetres in size within a dilated duct. Larger lesions dilate the duct more and extend along the long axis of the duct presenting a spheroidal shape. With ductal obstruction, the dilated duct with a papillary lesion may resemble a cyst with an intracystic solid component, this variant being termed as an intracystic papilloma. Histologically, papillomas show hyperplastic proliferation of ductal epithelium, having an arborescent growth pattern with branching fibromuscular core of myoepithelial and epithelial cells. Lesions may be pedunculated or broad based [3, 4]. Multiple peripheral papillomas are a rare entity in which the lesions are located in the peripheral duct system within the terminal ductal lobular unit. Several adjacent ducts are involved with segmental dilatation of the ducts, often resulting in a peripherally located mass. The incidence of nipple discharge is lower in these patients compared with the papillomas in larger ducts. There is an increased risk of carcinoma in peripheral papillomas which is directly related to the degree of cellular atypia. Peripheral papillomas are often associated with coexisting malignancy with a reported incidence of 10–30% [3–6]. Intraductal papillary carcinoma (IPC) is an uncommon ductal malignancy forming papillary structures. Histologically papillary carcinoma shows multilayered 843

S Ganesan, G Karthik, M Joshi and V Damodaran

papillary projections with microscopic frond formations extending from the vascularized stalks. These lesions characteristically lack the myoepithelial layer present in benign lesions. IPCs are reported in patients from 25 years to 89 years of age with a peak incidence between 40 years and 75 years [4]. IPCs have a wide spectrum of presentations varying from a focally invasive lesion with microscopic frond formation to a large mass located within a cystically dilated duct. Multiple lesions tend to occur within the same duct with papillary configurations.

Ultrasound features Ultrasound features of intraductal papillary neoplasms (IPN) primarily depend on the gross macroscopic appearance of the lesion. Three basic patterns of IPNs are recognized on ultrasound – intraductal mass with or without ductal dilatation, intracystic mass and a predominantly solid pattern with the intraductal mass totally filling the duct [3, 7]. If the tumour is small, a focally dilated duct may be the only observation. A solitary dilated duct, even in the absence of a demonstrable intraductal mass, is highly suggestive of an intraductal papilloma, especially, if the patient is presenting with a serosanguinous nipple discharge [3]. Dilated duct with an intraductal mass or a cyst with an intracystic solid mass is the hallmark of intraductal papillomas (Figure 1). The ductal component may vary in size from a minimally dilated duct to a large cystically dilated, obstructed duct. Similarly the intraductal soft tissue component may range in size from a very small lesion which may be impossible to image to a large mass completely filling the dilated duct or the cyst obscuring the ductal or cystic component simulating other solid masses (Figure 2) [3, 8]. Han et al analysed the

Figure 1. A focal mass arising from the ductal wall with relatively narrow base of attachment is present. Note the branching pattern and peripheral fronding typical of intraductal papilloma/papillary carcinoma (D-Duct).

844

Figure 2. Papillary carcinoma. A moderately large mass is seen to almost totally fill the entire dilated duct. Relatively hypoechoic debris is seen to fill the peripheral duct adjacent to the mass. A short segment of proximal duct is noted at 10– 11 o’clock position.

relationship between the mass and the duct on ultrasound and classified the masses into four categories: type I – intraluminal mass; type II – extraductal mass; type III – purely solid mass; type IV – mixed variety [9]. Benign papillomas are known to exhibit calcifications. These calcifications tend to be dense and coarse (Figures 3–5).

Figure 3. Benign intraductal papilloma with calcification. Focal dilatation of a solitary duct with intraluminal echogenic debris. Note small focal mass with dense, coarse calcifications in the proximal duct, with ductal obstruction (small arrows). The British Journal of Radiology, October 2006

Pictorial review: Intraductal papillary neoplasms of breast

(a)

(b)

Figure 4. Benign calcified intraductal papilloma with adjacent oil cyst. (a) Small focal mass with coarse, irregular and dense calcifications (small arrows) adjacent to a cystic mass (C). Echogenic floating debris within the cyst with floating fat-fluid level (long arrows). (b) Colour flow studies – focal increase in flow within the mass.

Small IPNs are often mammographically negative. A minimal to moderate duct dilatation may be observed on mammography as a progressively tapering band-like density extending from the nipple-subareolar region towards the breast parenchyma for a variable distance (Figure 6). Larger lesions in a dilated duct may resemble any other focal well-circumscribed dense mass on

mammography (Figure 7). Calcified IPNs exhibit dense, central, peripheral or combined form of coarse calcification similar to those seen in cases of calcified fibroadenomas (Figure 8). Boonjuwetat et al described the mammographic appearances of papillary neoplasms in a series of 15 cases. They reported that most lesions presented as solitary dense masses with no evidence of

(a)

(b)

Figure 5. Calcified giant intraductal papilloma. (a) Ultrasound and (b) mammography demonstrate a large, bilobed, densely calcified mass with distal shadowing (arrows). The British Journal of Radiology, October 2006

845


(a)

(b)

(c) Figure 6. Benign intraductal papilloma. (a) Mammography – oblique band like density along inferolateral quadrant of the left breast. (b) Ultrasound – focal dilatation of a solitary duct with an intraluminal mass arising from the ductal wall. (c) Doppler studies – distinct vascular pedicle within the central core with branching vessels arborising within the mass.

calcification in any of these lesions. A few lesions were mammographically negative either due to the size of the lesion or due to the dense parenchymal pattern [10]. IPNs are highly vascular and have a propensity to bleed spontaneously. Spontaneous haemorrhage into a dilated duct characteristically produces a fluid–debris level due to the denser cellular components settling down to the dependant position. The supernatant serum is anechoic while the dependant cellular debris is echogenic. Presence of fluid–debris level in a cyst, representing spontaneous 846

bleed into the cyst, is virtually suggestive of a mural proliferative lesion (Figures 9 and 10) [11]. IPNs have a characteristic flow pattern on colour flow studies. A distinct vascular pedicle is identified in IPNs within the central core with branching vessels arborising within the mass. Colour flow studies are sensitive in identifying even very small IPNs, in view of its characteristic vascularity (Figure 11). Intraductal papillomas and papillary carcinomas have considerable overlap in imaging features and it may not be possible to differentiate them on ultrasound. In The British Journal of Radiology, October 2006


(a)

(b)

Figure 7. (a) Mammography – focal well circumscribed dense mass along the retroareolar region of the left breast. (b) Ultrasound – large cystic mass with echogenic debris totally filling the cyst.

an older age group presence of a larger solid component and evidence of spontaneous intracystic bleed are more suggestive of papillary carcinomas than benign papillomas (Figure 12) [9]. The differential diagnosis of IPNs depends upon the basic imaging appearances. Presence of sectoral dilatation of ducts with no demonstrable intraductal mass has to be differentiated from mammary duct ectasia, which is a chronic inflammatory condition. Bleeding into a duct, inspissated material in mammary duct ectasia and ductal carcinoma in situ may produce dilated ducts with intraductal filling defects resembling IPNs. Mammary ductectasia is often bilateral and tends to affect multiple ducts. In intraductal carcinoma, ductal dilatation is unilateral, sectoral and irregular with ductal wall thickening. Colour flow studies reveal lack of flow in inspissated intraductal debris. Increased or variable periductal flow may be present in intraductal carcinomas while the IPNs reveal the characteristic arborescent The British Journal of Radiology, October 2006

vascularity. A cystically dilated duct may resemble a simple cyst when the intracystic component is very small. This has to be differentiated from other cystic masses like a simple cyst, complex cyst, haematoma, abscess and fat necrosis. A dilated duct with an intraductal solid component consisting of a central core and peripheral fronds, with characteristic flow on colour flow studies, is virtually diagnostic of IPNs. When the mass is large enough to fill the dilated duct or the cyst, it may not be possible to delineate the peripheral ductal or cystic component. These lesions have to be differentiated from other solid masses [3]. Fine needle aspiration cytology (FNAC) or core biopsy is required in all cases to arrive at a definitive diagnosis even though the imaging findings are suggestive of IPNs. FNAC from non-palpable small masses and from the solid component in large cystic lesions can be performed under ultrasound control. At our institution, small papillary lesions within a minimally dilated duct, observed as 847


Figure 10. Colour flow studies in intraductal papillary neoplasms (IPNs). Distinct vascular pedicle within the central core with branching vessels arborising within the mass.

Figure 8. Mammography. Well circumscribed peripherally calcified lesion within a progressively tapering band-like density extending from the nipple-subareolar region towards the breast parenchyma.

incidental findings in an asymptomatic patient on sonomammography are not subjected to FNAC or core biopsy. These patients are advised serial follow up with sonomammography. Larger lesions, lesions with atypical characteristics and lesions in symptomatic patients are subjected to FNAC and core biopsy. The criteria for calling

a core B3 are the following: papillary lesion, atypical intraductal epithelial proliferations, radial scar, lobular neoplasia and fibroepithelial lesions. Lee et al reported that the B3 core group is a more heterogeneous group and has a lower rate of malignancy on further biopsy. However, they concluded that the majority of lesions categorised as B3 required excision [12]. Agoff et al evaluated the need for surgical excision in intraductal papillary neoplasms and suggested that all such lesions with atypical ductal hyperplasia required excision owing to the high rate of associated neoplasia [13]. Although some lesions categorised as B3 lesions on core biopsy may be re-categorised

Figure 11. Benign intraductal papilloma in a cystically Figure 9. Intracystic papillary carcinoma. A moderately large cystic mass in the central breast region with large intracystic solid component is present. The mass is attached to the wall by a broad base and shows irregular branching pattern with peripheral fronding. (D – Cystically dilated duct).

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dilated duct. Cystically dilated duct with a small focal mass attached to the ductal wall with a narrow base between 10 and 11 o’clock position. A short segment of dilated proximal duct is identified with dependant echogenic debris with layering effect forming a fluid-debris level.



on excision biopsy as B2 lesions, all lesions categorised as B3 and above are subjected to excision biopsy at our institution.

8.

References

9.

1. Cilotti A, Bagnolesi P, Napoli V, et al. Solitary intraductal papilloma of breast. An echographic study of 12 cases. Radiol Med (Torino) 1991;82:617–20. 2. Kramer WM, Rush BF. Mammary duct proliferation in the elderly: a histopathologic study. Cancer 1973;31:130–7. 3. Tohno E, Cosgrove DO, Sloane JP. Benign processestumors. In: Tohno E, Cosgrove DO, Sloane JP, editors. Ultrasound diagnosis of breast diseases. Edinburgh: Churchill Livingstone, 1994:94–7. 4. Dahnert W. Breast disorders. In: Dahnert W. Radiology review manual, 4th edn. Philadelphia, PA: Williams and Wilkins, 1999:458–74. 5. Haegenson CD. Diseases of the breast, 3rd edn. Philadelphia, PA: W.B. Saunders, 1986:136–75. 6. Murad TM, Contesso G, Mouriesse H. Papillary tumors of large lactiferous ducts. Cancer 1981;48:122–33. 7. Kasumi F. Ultrasound of breast diseases. Shinohara – shuppan Co., Tokyo, 1983. Cited in Benign processes – Tumors, Intraductal Papilloma, In: Tohno E, Cosgrove DO,


10.

11.

12.

13.

Sloane JP, editors. Ultrasound diagnosis of breast diseases. Edinburgh: Churchill Livingstone, 1994:94. Yang WT, Suen M, Metrewell C. Sonographic features of benign papillary neoplasms of the breast: review of 22 patients. J Ultrasound Med 1997;16:161–8. Han BK, Choe YH, Ko YH, Yang JH, Nam SJ. Benign papillary lesions of the breast: sonographic-pathologic correlation. J Ultrasound Med 1999;18:217–23. Boonjunwetwat D, Prathombutr A. Imaging of benign papillary neoplasm of the breast: mammographic, galactographic and sonographic findings. J Med Assoc Thai 2000;83:832–8. Tohno E, Cosgrove DO, Sloane JP. Diagnostic features on ultrasound. In: Tohno E, Cosgrove DO, Sloane JP, editors. Ultrasound diagnosis of breast diseases. Edinburgh: Churchill Livingstone, 1994:58–9. Lee AH, Denley HE, Pinder SE, Ellis IO, Elston CW, Vujovic P, et al. Excision biopsy findings of patients with breast needle core biopsies reported as suspicious of malignancy (B4) or lesion of uncertain malignant potential (B3). Histopathology 2003;42:331–6. Agoff SN, Lawton TJ. Papillary lesions of the breast with and without atypical ductal hyperplasia: can we accurately predict benign behaviour from core needle biopsy? Am J Clin Pathol 2004;122:440–3.

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November 2006 Volume 79 Issue 947

November 2006, Volume 79, Issue 947 Government reform of the National Health Service: implications for radiologists and diagnostic services ●

Reporting overexposures and unintended exposures in diagnostic procedures ●

Are targeted contrast agents realistically going to reach the clinic? Recent regulatory experience with targeted MRI contrast agents ●

The development and optimization of high spatial resolution MRI for imaging the oesophagus using an external surface coil ●

Pancreatic adenocarcinoma: signs of vascular invasion determined by multi-detector row CT ●

Conversion factor for CT dosimetry to assess patient dose using a 256-slice CT scanner ●

Equivalent dose to organs and tissues in hysterosalpingography calculated with the FAX (Female Adult voXel) phantom ●

Comparison of radiation doses to patients undergoing standard radiographic examinations with conventional screen–film radiography, computed radiography and direct digital radiography ●

Dose escalation to combat hypoxia in prostate cancer: a radiobiological study on clinical data ●

Scatter from radiotherapy beams emerging from primary barriers: an aid to bunker design ●

The use of megavoltage cone-beam CT to complement CT for target definition in pelvic radiotherapy in the presence of hip replacement ●

●

CT of thoracic lymph nodes. Part I: anatomy and drainage

●

Book reviews

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Various holes and lesions

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Book reviews


COMMENTARY

Government reform of the National Health Service: implications for radiologists and diagnostic services G W L BOLAND,

MRCP, FRCR

Department of Radiology, Massachusetts General Hospital and Harvard Medical School, Massachusetts General Hospital, Boston, MA 02114, USA ABSTRACT. Demand for radiology services within the National Health Service (NHS) continues unabated and current NHS operations cannot keep up with demand. Therefore, to meet this demand, the government has decided to outsource a significant number of investigations to the independent sector and will actively promote patient referrals to the new government sponsored Treatment Centres as they become available. This presents opportunities to patients, but threatens existing public sector providers (including doctors) as competition for radiology services may result in both loss of patient referrals and revenue to these providers. This article is a personal opinion and will focus on the current challenges facing the provision of radiology services in the NHS. I will suggest the possible negative outcomes for providers (NHS hospitals and staff alike) and will offer strategies, tactics and tools that can be employed to counter the threat to their existing services.

The founders of the National Health Service (NHS) intended that it was the responsibility of government to provide free healthcare to all British citizens at the point of delivery. However, they could not have foreseen the dramatic medical innovations over the last 50 years, nor their cost. Nor could they have predicted the massive increase in demand for these services. Whether they would have advocated the same mandate with this knowledge in hindsight is open to debate, but the relentless and increasing demands on clinical services have undoubtedly demonstrated profound weaknesses in the structure of the NHS. While this is true for most clinical services, radiological services have been particularly exposed. Massive increases in demand have meant prolonged waiting lists for essential radiological services, with some patients having to wait months for basic diagnostic and therapeutic procedures [1]. The Audit Commission in 2002, for instance, demonstrated average waiting times for outpatient MRI at 20 weeks and for CT at over 6 weeks [1]. Although waiting lists for these services are coming down, most consider waiting lists for radiological services within the NHS to be unacceptably long. Worse, the demand for services continues unabated due to an ageing population and because patients and their doctors see an ever increasing value from these radiological services. Newer and better imaging techniques combined with newer imaging applications means that imaging is now increasingly Address correspondence to: Dr Giles Boland, Department of Radiology, White 270C, Massachusetts General Hospital, 55 Fruit Street, Boston, MA 02114, USA. E-mail: [email protected].

The British Journal of Radiology, November 2006


used as a primary diagnostic tool [2, 3]. This in turn has raised the stakeholder expectations from radiology for hospital administrations, referring physicians and patients alike. However, as the number of radiology investigations performed today cannot meet demand (and is likely to deteriorate further), an unsustainable tension has arisen between the NHS providers of radiology services (radiologists and radiographers) and their customers, the referring physicians and patients. The cause of this crisis within radiology, however, is not so clear. Many stakeholders, particularly the providers, advocate that the NHS has been significantly and chronically underfunded, leading to too few radiologists, radiographers and equipment (i.e. resources). Undoubtedly there is some truth to this despite the recent expansion in the number of trainee radiologists [4]. Britain has some of the fewest radiologists per head of population when compared with other Western nations, approximately half that of Germany and one quarter that of France, although it is interesting that Britain also has one of the fewest number of examinations performed per head of population in Western nations, suggesting it would need fewer radiologists (the UK performs 52 investigations per 100 head of population annually compared with a European annual average of approximately 100 investigations per 100 head of population) [5]. Furthermore, government itself has recognized that the NHS has been underfunded and is currently injecting massive increases in cash into the NHS in order to bring its expenditure on healthcare in line with other Western European nations [6, 7]. In 861

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England, the cash budget for the NHS has doubled since 1999 [6] with large increases to radiological services. However, the real and sometimes perceived lack of resources is too often used as the explanation for much of the NHS’ problems in general, and particularly for radiology, due to the high cost of its resources. Indeed, attempts by recent governments and others to seek better value out of existing resources have often been met with resistance [8, 9–11]. However, without precise data to benchmark the utilization of current resources, it is not realistically possible to accurately argue one way or the other. The Audit Commission of NHS radiological services in 2002 gave some insight as to what analysis of existing resource utilization might show – that a significant component of the shortfall in radiological capacity is due to the poor use of its resources. This report demonstrated major differences in the hours of operations of existing equipment and of radiologist productivity [1]. Most CT and half of MRI scanners were operational for less than 9 h day21 and examination throughput for the same scanners can be less than 50% of that seen in the USA [12]. Productivity rates for radiologists (the number of investigations they interpret) vary significantly, with some consultants being three times more productive than others [1]. It has also been demonstrated that other clinical services within the NHS also fail to use their resources effectively [8]. As more accurate data become available it will be increasingly evident to most informed healthcare professionals (and probably the public) that public sector NHS reform is desperately needed. Such data will likely continue to point to marked variations in resource utilization. It will therefore become increasingly hard to deny the widespread inefficient use of existing resources, an unacceptable situation considering the high cost of these resources and that they are ultimately paid for from public funding. What is controversial, however, is what reform is required, who should pay for it and what are the goals, incentives and achievable outcomes? Since the 1980s there has been a number of significant government policy agendas that have attempted to address this problem, but many have either run their course or been overturned by succeeding governments. However, the general concept of NHS reform and particularly market reform have stuck, and the current government has been actively seeking means to reduce waste and increase capacity through more efficient use of limited resources [6, 7, 13, 14]. This in turn has led to further tension amongst healthcare providers, particularly the clinical staff, as reform in healthcare is usually viewed with suspicion [8– 11]. Reform inevitably requires an evaluation and measurement of current practices to establish the strengths and weaknesses of those services. National Health Service physicians have been generally resistant to having their practices measured [8]. Not only have they seen themselves historically as autonomous professionals, some may fear that any measurement of their services may expose significant discrepancies (either personal or organizational). Indeed, some may argue that one of the primary reasons why such little data on resource utilization is available is because it might expose such discrepancies and therefore give 862

the doctor’s adversaries the data they need to force change. All this is, of course, understandable considering the organizational structure of the NHS. It is generally understood that it is not the role of doctors to measure resource utilization including their own productivity; rather it is left to the myriad of hospital managers that have sprung up over the last 20 years [7, 9]. Furthermore, if confronted with productivity data by their managers, physicians may feel that these managers do not fully understand the clinical issues that drive the healthcare process and therefore consider their data of little relevance. The result is that a situation of general mistrust and lack of morale pervades the system, whereby physicians feel over appraised and over inspected by professionals who either do not have their interest at heart, or more importantly, do not have the patient’s interest at heart [8]. A relative stalemate has therefore arisen over the last 10 years whereby the interests of physicians do not necessarily meet those of other stakeholders, particularly managers and government civil servants. Interestingly, both sides passionately feel that they are defending the interests of the patients. This lack of clear vision from NHS doctors to engage and address the variances in resource productivity and utilization in the public sector has forced the government to address the problems for them. The government has embarked upon an ambitious attempt to enforce better utilization of its resources, which in turn is hoped to increase the value of its healthcare system to its population. First, the government insisted on a new physician contract (albeit agreed to with significant difficulty), which was partly an attempt to align productivity with incentives [7, 8, 11]. Now it is attempting to reduce waiting lists by outsourcing large swathes of low risk outpatient surgery and diagnostics to the independent sector, either through development of specialized Treatment Centres currently under construction or directly to existing independent providers [15]. Indeed, the NHS is planning to purchase £1 billion of diagnostic services directly from the independent providers and hundreds of thousands of MRI services are already being outsourced to independent providers [6]. It is predicted that as much as 15% of the NHS funding may be shifted to the independent sector over the next 2– 3 years as the government attempts to meet its political objectives [7, 8]. Perversely, these events may initially be strongly welcomed and encouraged by NHS healthcare providers, both managers and physicians alike, as they attempt to manage the prolonged waiting lists and meet the government targets imposed to help reduce patient lists. Some physicians may actually feel that this scenario may be a vindication of their position after all – namely, the government is effectively agreeing to their long standing opinion that more resources are needed – and bolstered by the fact that the government has now provided the necessary funding, albeit partially to the independent sector. While these beliefs may bring satisfaction to some doctors in the short term, they may find that the government has the ‘‘last laugh’’. Unless physicians are willing to critically and objectively look at their The British Journal of Radiology, November 2006

Commentary: Government reform of the NHS

organizations and departments and make significant attempts to work collaboratively to improve performance and productivity, their influence and even income may be marginalized. Under the government’s proposals, the funds to pay for these clinical services will follow the patient to the institution that provides the services – the independent sector [7]. If the government’s intentions are played out, a large proportion of outpatient imaging (perhaps up to 15% or more) may be outsourced to the independent sector. While this may ease the current waiting list crisis within hospitals, it will inevitably deprive these hospitals of much needed income down the road. The effect of these government initiatives will equate to competition in the conventional business sense. A potential outcome for public sector hospitals that do not undergo significant reform and compete may be loss of valuable market share, which in turn will lead to lost revenue. In any traditional business, competition, which threatens its survival, is usually countered with restructuring and reform in order to become more productive and in turn provide similar services or products for less cost (i.e. they become more competitive). Ideally, in the process, restructuring may yield an increase in quality or value in the services, a very favourable circumstance. Within radiology, witness the impact that voice recognition technology can do to replace traditional transcriptional services. Within a few months of implementation, the financial costs are significantly less, but the value of the radiologist services (the ability to deliver a timely report) has greatly increased [16, 17]. The result is that customers (i.e. patients and referring physicians) have now gained a better product for less cost. It should come as no surprise to anyone within the NHS that this is exactly what the current government is trying to do. While many healthcare professionals may not like to think of themselves as conventional businesses, it is now increasingly evident that they need to think and act that way. While any changes may be good news for patients, they may not be good news for many doctors. Through an idiosyncratic historical contractual arrangement, doctors have been able to earn significant income outside their primary work place (the NHS) within the independent sector [18]. Indeed, this contractual arrangement may be seen by some as having presented a conflict of interest to senior doctors in the NHS, in which longer waiting lists in their NHS practice served to increase the likelihood that those patients who could afford to (an increasing proportion) would move to their more lucrative private practice. However, this potential opportunity is under threat. If the government’s reforms are successful, NHS waiting lists may start to shorten significantly. This could adversely affect many doctors who perform private practice. Doctors are likely to see a dwindling of their private practice revenues as more and more patients are either treated or evaluated in the government sponsored independent sector institutions or existing independent providers (paid for from government revenue). Any doctor thinking that they may benefit from the newly outsourced business to the independent sector is likely to be disappointed. The government has stipulated that most of the new opportunities in the independent The British Journal of Radiology, November 2006

sector may not be available to NHS employees or if they are, there will be significant limitations to the opportunities. What are radiologists and other doctors to do in this unpredictable environment? One option is to do little, assuming that this is yet another government attempt at healthcare market reform that will run its course. While this is possible, the evidence is against this outcome. The government has allocated vast new sums of money towards this agenda and is determined for it to succeed. Those doctors who fail to respond creatively and positively risk being left behind. Not only will the doctor’s influence be marginalized, it may be too late for them to save their private practice income should the government reforms be successful. Besides, even if the government initiatives fail, positive action by doctors to improve productivity and efficiency in the public sector should bring strong benefits to their organization and therefore the patients they serve. No one can argue that delivering timely efficient high quality healthcare to patients is not something worth striving for. A more sensible option therefore would be for doctors to rapidly understand what changes are happening, why they are happening and what they can do to respond to these changes. First and foremost, they need to recognize and accept that this is a serious effort by the government to increase the value of the healthcare system to its population. While doctors may not necessarily agree with these initiatives, they need to replace their ambivalence and inertia by recognizing that health systems in 2006 effectively have to act along efficient business lines. Those who remain entrenched, believing that clinical care and business do not mix, may have a lot to lose. Doctors need to rapidly gain an understanding of the business fundamentals required to efficiently operate their specialty. For radiologists, it is vital they develop an understanding of effective resource utilization and mechanisms to improve departmental productivity. After all, they are partly responsible for managing and operating some of the highest cost items of capital equipment within their organizations. Typically this will require analysis of their existing assets to determine current operational productivity. Once this is known, it should then be possible to benchmark their operation against the best practices in the industry. The analysis will inevitably require operational changes to increase productivity from their existing assets, i.e. greater patient throughput. While some additional resources may be needed to make this happen, the costs are minimal compared with the overall benefit to the organization. For example, in radiology, two radiographers can perform more than twice as many CT scans as a single radiographer (and possibly even three times) [19, 20]. The number of patients scanned each day per CT scanner could increase from approximately 25 per day to over 90 patients per day (current volume of the busiest scanners at the Massachusetts General Hospital, Boston, MA). The cost of an additional radiographer is minimal compared with the thousands of additional CT scans that can be performed annually. Furthermore, increasing operations into the weekend and evening hours will further increase capacity and help to significantly shorten 863

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waiting lists. Analyses of many other parameters should also be performed including length of waiting lists, time to perform the procedure, report turnaround time, radiologist’s productivity and costs per examination, amongst others [19]. Such an analysis of the operations will then permit radiologists to devise strategies and the tactics to counter the new threats and position their organization to compete with the independent sector. While radiologists may prefer to delegate analysis of their operations to their managers, there is unlikely to be broad acceptance by doctors for any changes that are recommended by other groups. Besides, because doctors are on the ‘‘front line’’ of the healthcare delivery cycle, they should be intimately involved in developing solutions to improve their operations. Doctors therefore need to become more business-like and understand their operations through careful data analysis and then cooperate with managers to maximize capacity and utilization of their fixed assets. Inevitably this will require physician leadership, which will in turn require greater knowledge, which for many will require further education and training. It should not, of course, be assumed that managers understand appropriate asset management either and further training may be needed by them to understand how they can help maximize their organization’s resources (staff and equipment). Doctors and managers will both need to learn to better cooperate with each other to improve the productivity and efficiency of their department, thereby creating a more valuable product for their organization. While this process is not easy (and typically change for some doctors has been difficult), it is essential that physician leaders respond to the challenge and convince their peers that it is in their best interest to change. This will require particular skill, particularly as many radiologists will immediately demand to know who is going to read the additional studies generated by any increase in productivity, particularly due to the relative shortage of radiologists. The answer to this conundrum is not at all clear given the current contractual arrangement within the NHS. There has structurally been little incentive for doctors in the NHS to work towards increasing productivity and services within the NHS. Some may, of course, argue that this is one of the NHS’ strengths; a system that provides financial incentives to doctors to perform more investigations on a fee for service basis is likely to be more costly [21]. However, it is extremely hard to significantly change organizational behaviour without offering some form of incentive. Unfortunately, attempts by some NHS physicians to help improve productivity and services have frequently been met with resistance from their managers who fear the increase in costs that may occur, even if marginal. However, Trusts are becoming more financially autonomous, so it is quite conceivable that radiologists could negotiate additional pay for the additional studies interpreted, perhaps on a modified fee for service basis. Besides, this financial arrangement has existed in the independent sector for many years. Furthermore, radiologists will be unlikely to participate in a service improvement program without some form of incentive, even if it is the right thing to do. It is therefore 864

strongly recommended that managers seriously consider incentive schemes for radiologists who add value to their organizations and help to shorten patient waiting lists. In summary, in order to bring greater value from the health service to its citizens, the British government is actively creating a publicly funded independent or partindependent sector that is likely to directly compete with the public sector NHS. A likely intention by the government, perhaps covertly, is to force public sector NHS organizations and particularly doctors to adapt to this new competitive threat. Those doctors who ignore or even resist these changes may have a lot to lose both in influence and financial reward. However, those doctors who respond creatively with knowledge, skill and leadership to maximize the productivity and value of radiology services within their organizations are likely to receive greater respect, recognition and probably financial reward. This in turn should increase the value that radiologists bring to their department, their organization, but most importantly, to the patients they serve.

References 1. Radiology. Review of National Findings. Audit Commission – United Kingdom July 2002 no. 7. 2. Wolbarst AB, William R, Hendee WR. Evolving and experimental technologies in medical imaging. Radiology 2006;238:16–39. 3. Thrall JH. Reinventing radiology in the digital age part III. Facilities, work processes, and job responsibilities. Radiology 2005;237:790–3. 4. Board of the Faculty of Clinical Radiology. The Royal College of Radiologists. Changing working lives. London: Royal College of Radiologists, 2005. 5. Board of the Faculty of Clinical Radiology. The Royal College of Radiologists Clinical Radiology: A workforce in crisis. London: Royal College of Radiologists, 2002. 6. High Anxiety. The NHS’s financial difficulties. Economist Magazine Dec 20th 2005. 7. Pollock AM. NHS plc: The privatisation of our health care. London: Verso, 2004. 8. Klein R. Britain’s National Health Service revisited. N Engl J Med 2004;350:937–42. 9. Dawson P. The British National Health Service. J Am Coll Radiol 2004;1:287–92. 10. Hargreaves S. A quarter of consultants ready to resign over new contract. BMJ 2003:326:569. 11. Doctors warn new contract will herald the death of the NHS. Guardian Newspaper, London. May 21, 2003. 12. Boland GWL, Palumbo D, Tabor-McEwen K, Harlem P, Pathak A. CT productivity differences between the UK and USA: opportunities for the National Health Services. RSNA Annual Scientific Meeting 2004;231(D):459. 13. Shapiro J. Can Labour take the NHS to market? BMJ 2005;331:359. 14. Steven S. Reform strategies for the English NHS. Health Affairs 2004;23:37–44. 15. Private treatment centres to expand. Guardian Newspaper, London. January 8, 2005. 16. Amit Mehta. Voice-Recognition Technology. The primary benefits of voice recognition have been decreased turnaround time and enhancement of the capabilities of PACS. In: Imagingeconomics.com [Accessed 13 September 2006].


Commentary: Government reform of the NHS 17. Mardini M, Mehta A. Voice recognition. In: PACS: A guide to the digital revolution. Springer Publications, 2005. 18. Bulstrode C. Embarrassing greed. BMJ 1995;310:198–9. 19. Boland GWL. Stakeholder expectations for radiologists: obstacles or opportunities. J Am Coll Radiol (In press JACR-D-05-00165).


20. Boland GWL, Meehan M, Doncaster R. Three technologist model for CT: cost implications and benefits. RSNA Annual Scientific Meeting 2004;231(D):365. 21. Maitino AJ, Levin DC, Parker L, Rao VM, Sunshine JH. Practice patterns of radiologists and nonradiologists in utilization of noninvasive diagnostic imaging among the Medicare population 1993–1999. Radiology 2003;228:795.

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COMMENTARY

Reporting overexposures and unintended exposures in diagnostic procedures P R CLEWER,

MPhil, MSRP, MIPEM

and P C JACKSON,

BSc, PhD, FIPEM

Department of Medical Physics and Bioengineering, Southampton General Hospital, Southampton, UK ABSTRACT. Legislation requires the reporting to regulatory authorities of incidents in which patients have been exposed to ionizing radiation to an extent ‘‘much greater than that intended’’. The authorities have published guidance on what is considered to meet this requirement. However, there is still some confusion regarding, particularly, the necessity to report some unintended doses. It is believed that there is a disproportionate amount of resource spent investigating some unintended exposures because all such exposures will have an effective overexposure factor of infinity, irrespective of the magnitude of the dose and the associated risk. This paper proposes changing the definitions of ‘‘overexposure’’ and ‘‘unintended exposure’’ and the adoption of a reporting process based upon risk assessment. All records and data would be collected and, if required, reported, but investigation of individual incidents would take place only for incidents carrying a greater risk than 1 in 10 000.

For many years now, the medical radiation protection community has lived with the legal requirement that when it is suspected that a patient has been exposed to ionizing radiation to an extent ‘‘much greater than that intended’’ due to equipment malfunction or defect, the employer must notify the Health and Safety Executive (HSE) forthwith. The requirement was stated in the Ionising Radiation Regulations 1985 (regulation 33(2)) and in the replacement Ionising Radiations Regulations 1999 (regulation 32(6)). The HSE published guidance [1] on what should be considered to be ‘‘much greater than intended’’. The HSE’s document included a table of ‘‘guideline multiplying factors’’ and stated that if the ratio of the suspected exposure to the intended exposure is greater than or equal to the appropriate multiplying factor, it should be presumed that HSE expects to be notified of the incident. The multiplying factors for diagnostic examinations were 3, 10 or 20, depending on the magnitude of the intended effective dose. For instance, an incident where a patient undergoing a chest X-ray receives an exposure of 20 times or more than intended, the HSE should be notified. For any CT examination, the ‘‘reporting threshold’’ was set at three times the intended exposure. The factors also applied to situations where equivalent multiples of patients received doses greater than intended since such situations might arise with equipment malfunctions. In 2000 the introduction of the Ionising Radiation (Medical Exposure) Regulations 2000, known as Address correspondence to: Mr Philip Clewer, Medical Physics and Bioengineering, Southampton General Hospital, Tremona Road, Southampton SO16 6YD, UK. E-mail: philip.clewer@suht. swest.nhs.uk.

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Received 24 January 2006 Revised 5 June 2006 Accepted 16 June 2006 DOI: 10.1259/bjr/68149575 ’ 2006 The British Institute of Radiology

IR(ME)R, brought a similar requirement for notification of incidents when a patient is exposed to ionizing radiation to an extent ‘‘much greater than intended’’, but due to any reason other than equipment malfunction or defect (regulation 4(5)). In this case, the notification is to the Department of Health (DoH). The guidance [2] issued with IR(ME)R stated that application of the HSE guidance on which doses are likely to be much greater than intended is appropriate. Therefore, for example, procedural failures that lead to a patient receiving 20 times the intended exposure from a chest X-ray should be reported to the DoH. Paragraph 6.8.2 of the same guidance stated that ‘‘patients who undergo a procedure that was not intended ... should be considered as having received an unintended dose of radiation’’. The introduction of the phrase ‘‘unintended dose of radiation’’ has caused some confusion because it appears to imply that unintended doses are separate to events described as ‘‘much greater than intended’’. This has raised some anomalies that lead to the question, ‘‘what is the purpose of reporting these incidents?’’ – a question to be addressed later in this paper. It appears that the DoH intended to publish guidelines that would be aimed specifically at incidents reportable under IR(ME)R rather than reportable under IRR99. These guidelines have not been forthcoming, but are still expected [private communication with DoH, 2005 and 2006]. For the purposes of this paper the authors will take the intended effective dose to be the ‘‘typical effective dose’’ published in 1999 by the National Radiological Protection Board (now the Radiation Protection Division of the Health Protection Agency) [3]. The British Journal of Radiology, November 2006

Commentary: Reporting overexposures and unintended exposures

Overexposures and unintended exposures If a patient is referred for a plain film chest X-ray, then the intended dose is 0.02 mSv. The guideline reporting threshold for an adverse event for such an examination is 20 times the intended dose. Therefore if the patient received 19 times the intended dose there is no requirement to report the incident, although the Employer may wish to do so. This means that the patient could receive an effective dose of 0.36 mSv over and above the intended 0.02 mSv, giving a risk of a fatal cancer of around 1 in 55 000, using the 5% per Sv risk factor model [4]. If a patient who is not intended to undergo any X-ray examination actually undergoes a chest X-ray for some reason, e.g. misidentification, the effective dose to that patient is 0.02 mSv with a risk of around 1 in 1 000 000 of a fatal cancer. Although the overexposure factor could be said to be infinity, the actual risk to the individual patient is much lower than the 19 times-the-intendeddose incident, but this incident is theoretically reportable because it was an ‘‘unintended exposure’’. From a patient perspective this may seem confusing in that the error with a lower risk of harm is reportable, whereas the error with a relatively higher risk of harm is not subject to that requirement. It is clear that an error is reportable if the overexposure factor is above the relevant threshold for reporting, e.g. someone has a pelvic radiograph (effective dose 0.7 mSv) when a chest radiograph (effective dose 0.02 mSv) was intended. In this example the overexposure is 0.72/0.02 (536) (assuming they go on to have the intended examination), i.e. greater than 20 and therefore reportable. However, if the sequence of events was the other way around the overexposure factor would be 0.72/0.7 (51.03). This would not be over the reporting threshold as far as an overexposure is concerned, but the chest radiograph would have been unintended and so the incident is reportable for this reason rather than overexposure. For instance, if a patient undergoes an incorrect examination, in practice this will lead to a ‘‘greater than intended’’ dose rather than an unintended dose since both the ‘‘incorrect’’ and ‘‘correct’’ examinations will have been performed. In this case, reporting to the DoH would be dependent on the magnitude of the overexposure. The DoH currently regards any case of wrong examination, such as ‘‘left ankle’’ instead of ‘‘right ankle’’, to be an ‘‘unintended exposure’’ and therefore reportable, regardless of magnitude. This means that hospitals have to report unintended doses even if they result in a dose and risk much lower than the dose and risk to patients who experience overexposures that do not exceed the reporting threshold. It appears that some hospitals are unclear as to the reporting requirements – some do report the ‘‘wrong ankle’’ examination whereas others do not.

The future In 2003, the HSE published proposals for a document, HS(G)226, to replace PM77. One of the significant The British Journal of Radiology, November 2006

proposals was that the guideline multiplying factor for all diagnostic procedures should be reduced to 1.5. Therefore all occasions when a repeat X-ray exposure is required due to equipment defect or malfunction would require reporting to HSE. This proposal attracted considerable comment and when the final document was published in March 2006 [5], the guideline multiplying factors for diagnostic radiology and nuclear medicine were divided into three bands as follows:

N N N

1.5 times intended dose – interventional radiology, radiographic and fluoroscopic procedures involving contrast agents, nuclear medicine with intended effective dose . 5 mSv and CT examinations 10 times intended dose – mammography, nuclear medicine with intended effective dose # 5 mSv but . 0.5 mSv, everything else not referred to elsewhere 20 times intended dose – radiography of extremities, skull, dentition, shoulder, chest, elbow, knee and nuclear medicine with intended effective dose # 0.5 mSv.

Hence the factors for middle and low dose examinations will remain unchanged, as do those for therapeutic procedures, but the higher dose diagnostic procedures have been given a guideline multiplying factor of 1.5. Therefore any repeat exposures in this category are reportable. The DoH has indicated that, certainly for the time being, they will not adopt the 1.5 times factor, but will continue with the 3 times factor for high dose diagnostic procedures [private communication with Department of Health, 2006]. The above arrangements are appropriate to situations where there has been equipment failure, but should the same approach be applied to the individual patient?

Why report? Perhaps it would be useful to return to the question ‘‘why do we report incidents?’’. IR(ME)R implement most of the requirements of the Medical Exposures Directive [6], but the Directive does not require member states to implement a reporting system. However, it does require in Article 11 that member states should ensure that ‘‘all reasonable steps to reduce the probability and the magnitude of accidental or unintended doses of patients from radiological practices are taken...’’. Perhaps this is in part a solution to the confusion; incident reporting can lead to prevention of similar occurrences by building up statistical evidence of where problems lie and where attention should be focused. The Directive also states that ‘‘the main emphasis in accident prevention should be on the equipment and procedures in radiotherapy, but some attention should be paid to accidents with diagnostic equipment’’. This implies a need to concentrate on preventing radiotherapy incidents because of their relatively high doses, and therefore risks, compared with diagnostic procedures. In practice there are two reasons for reporting incidents, be they overexposures or unintended exposures. 867

P R Clewer and P C Jackson

The first of these, as mentioned above, is to produce a national database of statistics on incidents, showing where the problems lie and which areas of practice require more attention to reduce the likelihood or magnitude of incidents in future. However, it is believed that this could be done more efficiently by employers being required to supply their own statistical information regularly, perhaps annually. The second reason for reporting incidents is that significant errors can be fully investigated by an independent body and appropriate actions considered in relation to employees and/or employers. These may lead to an investigation by the HSE or DoH, perhaps leading to an improvement notice or prosecution. Such incidents obviously require an individual notification to the appropriate regulatory body. The following is proposed as a way of combining ease of reporting for statistical purposes and the need for ‘‘enforcement’’ reporting:

N

N

N

N

N

Requirements for reporting should be redefined in terms of effective dose. Exposure is too vague a term for use in this context. Although there may be some controversy regarding the relationship between risk and effective dose, this would at least take into account the tissues and organs of the body that are irradiated. In fact, the PM77 does use effective dose in its table of guideline multiplying factors while the regulations refer to overexposures. ‘‘Overexposure’’ should be redefined as ‘‘overdose of radiation’’ and related to the occasion of a patient who it is intended should receive a dose receiving more than intended. By using the effective dose quantity, an overdose would include incidents where the patient initially received an examination of the wrong part of their body and then the correct examination. It would be intended that they receive an effective dose, but because they receive an additional dose they have been given a greater dose than intended. Therefore a patient who attends for a radiograph of the pelvis but receives a chest radiograph in error, followed by the intended pelvis radiograph, or vice versa, is the subject of an overdose of radiation, whichever way round the events occur – although the magnitude of the overexposure factor is very different. The system of published guideline multiplying factors for reporting of overdoses of radiation could be retained, but the authors suggest that an alternative system is adopted for reporting unintended doses as outlined below. ‘‘Unintended exposure’’ should be redefined as ‘‘unintended dose’’ and related to the occasion of a patient who it is intended should not receive a radiation dose receiving a dose. This would clearly identify those patients who should not have received any dose of radiation but who have for some reason, usually mistaken identity, done so. A threshold should be set, in terms of effective dose above which unintended doses should be reported to the regulatory authority individually and urgently. Below the threshold the incident should be logged by the employer and information sent to the regulatory authority at regular intervals as nationally agreed. Setting a threshold would enable appropriate effort to 868

be directed towards investigating and preventing reoccurrence of the more serious incidents in a timely manner. Of course, incidents involving doses below the threshold could still be reported individually if the employer believes there is merit in doing so. It is accepted that the public may view the setting of a threshold as condoning error and employers not being held to account to a higher authority. However, this situation might be addressed by the Healthcare Commission having access to the adverse event reporting of radiation events. If this philosophy is accepted, the next consideration is the setting of the reporting threshold for unintended doses and reviewing the threshold for reporting overdoses. Many values could be proposed, but before doing so it may be instructive to look at risks associated with other areas of life. In ‘‘Living with Radiation’’ [7], the National Radiological Protection Board published comparative risk data for the annual risk of death in the UK from some ‘‘common’’ causes. This included a 1 in 100 000 annual risk of being murdered, a risk of any 40-year-old individual dying from any cause as 1 in 700 and the annual average risk of death from cancer being 1 in 400. If the guideline multiplying factor of 20 for low dose examinations, which includes intended effective doses up to 0.5 mSv as described in PM77 [1], were to be retained then the threshold for reporting unintended doses could be set at 2060.5 mSv510 mSv. However, this does seem relatively rather a high dose for reporting individual incidents of this nature and is above the dose level of a CT chest examination. Another approach would be to set the threshold at 20 times a chest X-ray dose. This would set the threshold at 0.4 mSv and require individual reports of incidents involving many more types of diagnostic examination. This type of approach is subject to change in the effective dose of examinations over time and subjective assessment of the acceptability of a threshold. Alternatively, we could consider who is actually receiving this dose of radiation. A hospital patient receiving an unintended dose should not have undergone a medical exposure and is no different to any other member of the public. In 1993, the National Radiological Protection Board published guidance [8] relating to the latest international recommendations and included a recommendation for a public dose constraint for a single new source [of ionizing radiation] of 0.3 mSv per year. This dose, equivalent to a lateral lumber spine X-ray examination [4] and with a risk of a fatal cancer of 1 in 67 000 using the 5% per Sievert model, could be suggested as an appropriate reporting threshold for incidents where someone who should not have received a dose at all has been irradiated. This approach is perhaps less than acceptable in that it appears that a patient should not accept any greater risk of harm than a member of the public. This approach is not acceptable as, de facto, all forms of medical intervention have some risk. Apart from the unavoidable risk associated with the intended intervention, there is also a risk of an adverse event involving the patient but unrelated to the patient’s condition. The HSE’s approach to risk and benefit in the nuclear power industry may give some guidance. The HSE has The British Journal of Radiology, November 2006

Commentary: Reporting overexposures and unintended exposures

looked at the boundary between a tolerable risk (of death) and an unacceptable risk to both workers and the public from the use of a nuclear power station. For the public who have the risk imposed on them ‘‘in the wider interest of society’’, the boundary is said to be 1 in 10 000 per annum [9]. It would be difficult to maintain that the use of X-rays for medical diagnosis is not ‘‘in the wider interests of society’’ and so there will always be a risk of unintended exposures. Using the 5% per Sievert model, a risk of 1 in 10 000 is equivalent to an effective dose of 2 mSv, similar to that from a CT scan of the head and not very different from the average annual ‘‘background’’ dose to a citizen of the UK. This is still a reasonably high level and is likely to mean that no plain film unintended doses would require reporting. However, putting this into context, a lifetime risk of 1 in 10 000 of a fatal cancer is still very small in comparison with the natural risk of a fatal cancer of about 1 in 4 [10]. Table 1 summarizes the possible unintended dose reporting thresholds and reasons for their adoption. The authors propose a threshold of 2 mSv, above which employers should make a prompt individual report regarding unintended doses of radiation. Below this threshold the employer should retain statistical data for forwarding to the regulatory authority as required. By choosing this threshold, society is not viewing the risks associated with medical examinations involving ionizing radiations as being higher than those from the nuclear industry, rather that adopting such an approach has harmonized a consistent approach to radiation risks.

Conclusion The present system of reporting overexposures and unintended exposures can lead to confusion and the benefits are unclear. The current use of the linear no threshold dose–risk model means that all radiation doses should be as low as reasonably achievable. However, it does seem an inappropriate use of resources to give the same effort in investigating and reporting unintended doses to, say, the ankle as a CT scan of the pelvis. This paper proposes alternative definitions to ‘‘overexposure’’ and ‘‘unintended exposure’’ and the adoption of a reporting threshold for unintended doses. Unintended doses above this threshold would be the subject of an individual report to the relevant regulatory authority. Unintended doses below the threshold would be


Table 1. Possible unintended dose reporting thresholds Reporting threshold (mSv)

Argument for adoption

10

20 times the 0.5 mSv intended dose level from PM77 20 times the chest X-ray effective dose Public dose constraint Risk of 1 in 10000 of a fatal cancer

0.4 0.3 2

recorded by the employer and a statistical report made to the regulatory authority at appropriate intervals.

References 1. Health and Safety Executive. Fitness of Equipment Used for Medical Exposure to Ionising Radiation, HSE PM77. London, UK: HSE, first edition 1992, second edition 1998. 2. Department of Health. The Ionising Radiation (Medical Exposure) Regulations (2000 together with notes on good practice). London, UK: available in electronic format only at www.dh.gov.uk/assetRoot/04/05/78/38/04057838.pdf [Accessed 3 July 2006]. 3. National Radiological Protection Board. Guidelines on Patient Dose to Promote the Optimisation of Protection for Diagnostic Medical Exposures, Documents of the NRPB, vol. 10. no. 1. Didcot, UK: NRPB, 1999. 4. International Commission on Radiological Protection. 1990 Recommendations of the International Commission on Radiological Protection, ICRP Publication 60. Oxford, UK: Pergamon Press, 1991. 5. Health and Safety Executive. Equipment used in Connection with Medical Exposure. Guidance Note PM77 (third edition) March 2006 on HSE website: www.hse. gov.uk/pubns/guidance/pm77.pdf [Accessed 3 July 2006]. 6. Council of the European Union. Council Directive 97/43/ Euratom of 30 June 1997 on health protection of individuals against the dangers of ionising radiation in relation to medical exposure. The Council of the European Union, 1997. 7. National Radiological Protection Board. Living with Radiation. National Radiological Protection Board. Didcot, UK: NRPB, 1998. 8. National Radiological Protection Board. Board Statement on the 1990 Recommendations of ICRP, Documents of the NRPB, vol. 4, no. 1. Didcot, UK: NRPB, 1993. 9. Health and Safety Executive. Reducing Risks, Protecting People. Norwich, UK: HMSO, 2001. 10. Cancer Research UK website: http://info.cancerresearchuk. org/cancerstats/mortality/ [Accessed 3 July 2006].

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COMMENTARY

Are targeted contrast agents realistically going to reach the clinic? Recent regulatory experience with targeted MRI contrast agents A DZIK-JURASZ,

PhD, FRCS, FRCR

Cambridge, Massachusetts, USA

Received 6 April 2006 Revised 25 July 2006 Accepted 31 July 2006 DOI: 10.1259/bjr/68807002 ’ 2006 The British Institute of Radiology

Developments in the visualization of molecular events in vivo are progressing at an unprecedented rate, a characteristic of rapidly developing or emergent technologies. The need to target the molecular mechanisms of disease has been the driving force behind this rapid growth in knowledge whilst the bridging of disciplines, such as imaging science, chemistry and biology, in some cases with a good deal of entrepreneurship, has made such technological advances possible. At the vanguard is the field of targeted molecular imaging agents. Targeted molecular imaging agents are now used in biological, medical and pharmaceutical research where they are proving to be powerful tools in the investigation of experimental in vivo systems. It is almost taken for granted that these technologies will translate readily into clinical practice, providing unique molecular information on disease. But will targeted molecular contrast agents reach the clinic and deliver the benefits expected of them? Journals, conferences and societies are now dedicated to molecular imaging and targeted contrast agents. But to make an agent available to widespread clinical use requires regulatory approval from either the Food and Drug Administration (FDA) in the USA or the European Agency for the Evaluation of Medicinal Products (EMEA) in Europe. The main purpose of these agencies is to ensure that safe and efficacious agents reach the market while denying access to agents that do not meet these standards. In defining targeted molecular imaging agents I am specifically referring to agents that will be used widely in the clinic and be prescribed because of their role in guiding clinical management. This is quite different from how such agents are currently used in, for example, probing cell biology or validating drug targets. Positron emission tomography (PET) tracers are making an impact in drug development and are likely to increase in use, but the widespread translation of these 870

techniques into the clinic is unlikely to occur. Also excluded are the niche applications such agents might have in the manner of in vitro biomarkers [1]. Nevertheless, the medical community has high expectations of targeted molecular imaging. Witness the investment by national, private and business institutions across Europe and the USA. But if targeted contrast agents are increasingly viewed as potentially important clinical biomarkers, the role and value of using these agents remains to be established, and in this commentary I discuss in brief whether there is a realistic prospect of this occurring. Clinically, molecular imaging has been successfully applied for several decades via radionuclide imaging and, more recently, via modalities such as magnetic resonance [2, 3] and ultrasound. To address whether targeted contrast agents will reach the clinic requires an understanding of the process behind the development of a contrast agent. A contrast agent, despite rarely being administered more than once, is a drug and is treated as such by regulatory authorities. Prior to approval for clinical use, a contrast agent would have been guided through the discovery and development phases by a biotechnology or large pharmaceutical company, often in collaboration with academia. The processes require the coordination and cooperation of multidisciplinary teams. The financial outlay in bringing a contrast agent from the laboratory to the clinic is in the order of $100–150 million. This is a small sum compared with a therapeutic drug whose development costs, depending on the manner in which the data are interpreted, are reported to be on the order of $800 million [4]. However, the time required to develop such a contrast agent is no less than a therapeutic; it takes at least 8–10 years to bring any drug to market. More importantly, diagnostic imaging agents are very unlikely to return the type of profits seen with therapeutic drugs [5]. It is difficult to imagine how The British Journal of Radiology, November 2006

Commentary: Targeted contrast agents in the clinic?

molecularly targeted agents are going to bring an equal or greater return than the non-specific imaging agents currently available in the clinic. The cost of targeted imaging agents may therefore be higher than expected, and this effect could be compounded by increasing price controls being imposed on drugs by governments. Any drug must demonstrate safety and efficacy in order to gain regulatory approval, and much of the drug development process is geared toward achieving this goal. The most recent regulatory review of a MRI agent was that of EPIX Pharmaceuticals, Inc. (Cambridge, MA). An NDA (new drug application) had been filed in December 2003 with the FDA, supported by four phase 3 international multicentre trials. The drug (VasovistTM) is a blood pool agent targeted to human serum albumin, and is intended for use as an MR angiography agent. The FDA’s fundamental concern was with the efficacy data. My personal view (the author was the head of imaging at EPIX following the original approvable letter) is that the conflict between the FDA and the company can be distilled to two issues, namely (1) to what level is it reasonable and necessary to deconstruct a clinical experiment (in the manner of a laboratory bench top experiment)? And (2) what is a reasonable level of experimental standardization? These two issues are intimately related and affect how widely applicable the results are in clinical practice. The company unsuccessfully argued that clinical practice at individual sites or the equipment manufacturer’s sequence recommendations are a de facto basis of standardization (and showed that the variation in acquisition techniques was in fact very small and did not significantly affect the results). The FDA declined to accept this argument on the grounds that the efficacy results might still be biased in favour of the drug and therefore did not approve the compound (the actual term was ‘‘approvable’’, but the result is that the compound is not approved for use in man in the USA). Further details can be found in the press release section of EPIX’s website at www. epixpharma.com. The role of the FDA is to protect the health of the USA by approving, on a scientific basis, safe and effective drugs. But it also has the clear duty of not withholding beneficial drugs from the public. More often than not there is a fine balance between these two alternatives, especially at a time of heightened public and political scrutiny of the drug development business. The European authorities approved the drug in September 2005 after a 15-month review utilizing the same clinical data submitted to the FDA. This is not to say that there is a fundamental difference in the goals of EMEA and the FDA. Although organizationally different, EMEA differs little from the FDA with respect to the challenge of evaluating new targeted contrast agents. One fundamental difference is that EMEA’s decisions tend to be more reliant on the review of impartial outside experts for the technical assessment of an application. EMEA, for example, has a document analogous to the FDA’s guidance document on imaging agents ‘‘Points to consider on the evaluation of diagnostic agents’’. Of note is the European position that ‘‘…there are no generally accepted principles for the evaluation of efficacy of diagnostic agents’’, indicating it The British Journal of Radiology, November 2006

is for the sponsor to provide convincing clinical scientific evidence of efficacy. I remain troubled that the American public has been denied a valuable diagnostic agent. No one will deny the importance of experimental standardization, and standardization is straightforward when dealing with the physical parameters that define radionuclide or X-ray based imaging. Standardization is subtler when dealing with MRA where there is no meaningful standard comparator, and there are several different, but equally representative ways to approach imaging inflowing blood. In the EPIX trials, VasovistTM was compared against blood flow techniques such as time-of-flight and phase contrast angiography. EPIX’s argument was that the drug was effective in clinical practice, which is a legitimate means by which to seek approval. Indeed, adopting arbitrary standardized parameters, in the case of flow-sensitive techniques when there is no proven superiority among several approaches, might itself result in unexpected bias. Those in the field of MRA and particularly those who are involved in drug-related studies should oppose where appropriate the unilateral imposition of views that do not take into account the complexities of the modality being studied. Whatever the merits of this case, this series of events provides insight into the regulatory hurdles to the delivery of a drug/imaging agent to market. Such setbacks are ultimately resolvable, but remain a concern for industry strategists and decision makers who finance investments in research and development (R&D). The recent review of another MRI contrast agent (Combidex; Advanced Magnetics, Inc., Cambridge, MA; www.advancedmagnetics.com) highlights additional regulatory challenges faced by an imaging agent. Combidex, an iron oxide nanoparticulate, was issued an ‘‘approvable’’ letter by the FDA following review by an advisory committee, but for very different reasons from EPIX. An advisory committee consists of a panel of experts called by the FDA to make a recommendation, usually on one or a small handful of questions that guides the Agency’s decision making. Details of the FDA’s advisory committee (which sat in April 2005) on the subject can be found at http://www.fda. gov/ohrms/dockets/ac/05/transcripts/2005-4095T1.pdf. Bearing in mind that submission for regulatory approval encompasses a substantial body of work including chemistry, manufacturing, toxicity, pre-clinical and clinical data that are typically the sum of 8–10 years of work, Combidex remains unapproved because both the FDA and the independent advisory committee criticised aspects of its safety profile, requesting too broad an indication and in providing limited efficacy data. These two cases as indices of the regulatory climate for imaging agents in man have left a sense of considerable unease in the world of diagnostic imaging agent development. Efficacy, safety and regulatory issues not withstanding, additional technical issues for consideration in the development of a targeted contrast agent include identifying a target of interest, recognizing its biological significance and assessing the likelihood of discovering and synthesising an agent that will bind specifically to and be at sufficient concentration at the target site to generate a recordable signal. Any of these issues could prove to be as problematic as any regulatory, efficacy or 871

A Dzik-Jurasz

safety issue. But in part, the role of the translational radiologist in collaboration with scientific colleagues will be to navigate and guide those imaging studies that will define the clinical efficacy and relevance of a targeted imaging agent to satisfy both regulators and consumers. These studies will influence whether the commercial argument can be made that will stimulate investment in developing targeted contrast agents. Ultimately, the sales of an agent will be an index of its clinical value. A successful agent will be rapidly assimilated into clinical practice where further refinements in dosage, indication and safety profile will be determined. Can targeted agents be brought into the clinic, or will they simply remain laboratory tools divorced from clinical application? The answer is, of course, that any efficacious and safe agent can be brought to the clinic, but at what price? More important is the question of whether such agents will make an impact on clinical practice, and this is a more difficult answer to predict. Regulators are not in agreement as to the requirements for demonstrating likely clinical impact before approval. The need for development both in clinical trial designs and in appropriate criteria for regulatory requirements may override any other limiter on the adoption of molecular imaging in the clinic. Indeed it is surprising that the imaging community, despite its resolve to promote molecular imaging, has been almost mute in its response to the current regulatory climate. I strongly encourage the radiological community to give it some thought. As end-users, radiologists and medical-imaging scientists have influence over contrast agent development. Their technical expertise is sought to guide the practical application of imaging strategies, but the radiological community has limited exposure of the drug development world in terms of the requirements, needs and execution of imaging trials. Imaging experts could become more engaged in the development of new imaging agents, particularly of targeted contrast agents by:

N

Understanding the needs and requirements of the regulatory process behind drug approval

872

N N N

Promoting an improved understanding of clinical trial design, execution and requirements including the standardization of experimental design in multicentre trials Fostering an initiative to advance and promote new outcome measures, particularly in respect of targeted contrast agents Developing realistic strategies for the development of targeted contrast agents based on clinical need

The opportunity is there and need only be taken up by the radiological community. It would be sad if the only legacy of targeted imaging agents were a wealth of good science, but shameful if patients were not to benefit from its findings.

Acknowledgments I would like to express my gratitude to Dr J A Newmyer, Dr E Parsons and Dr A Uprichard for their critical review of the manuscript.

Disclaimer The opinions expressed in this manuscript are those of the author and do not necessarily reflect those of other parties or institutions. No financial conflict is declared.

References 1. Ludwig JA, Weinstein JN. Biomarkers in cancer staging, prognosis and treatment selection. Nat Rev Cancer 2005;5:845–56. 2. Padhani AR, Leach MO. Antivascular cancer treatments: functional assessments by dynamic contrast-enhanced magnetic resonance imaging. Abdom Imaging 2005;30:324–41. 3. Jaffer FA, Weissleder R. Molecular imaging in the clinical arena. JAMA 2005;293:855–62. 4. DiMasi JA, Hansen RW, Grabowski HG. The price of innovation: new estimates of drug development costs. J Health Econ 2003;22:151–85. 5. Nunn AD. The cost of developing imaging agents for routine clinical use. Invest Radiol 2006;41:206–12.



The development and optimization of high spatial resolution MRI for imaging the oesophagus using an external surface coil A M RIDDELL,

BSc, FRCS, FRCR,

C RICHARDSON,

DCRr,

E SCURR,

BSc, DCRr

and G BROWN,

MD, MRCP, FRCR

Department of Radiology, Royal Marsden Hospital, Fulham Road, London SW3 6JJ, UK ABSTRACT. This paper describes the development and optimization of an innovative technique using an external surface coil to obtain high resolution, thin section MR images of the oesophagus using volunteers. T2 weighted fast spin echo sequences were performed with and without cardiac gating. The field of view (FOV), matrix size, slice thickness, number of signal averages (NSA), and repetition time (TR)/echo time (TE) were altered to optimize signal to noise ratio (SNR) whilst maintaining spatial resolution. The effect of cardiac gating was also investigated. Workstation images were evaluated on the ability to visualize: individual oesophageal wall layers; perioesophageal fat; the azygos vein and wall of the descending aorta, giving qualitative assessment of image clarity. The optimum sequence enabled the layers of the oesophageal wall and perioesophageal tissues to be demonstrated in an acceptable scan time of 7.07 min. A FOV of less than 250 mm degraded image quality so that individual oesophageal wall layers could not be depicted and noise within the image impaired visualization of posterior mediastinal structures. The results indicate that high resolution imaging of the oesophagus using an external surface coil can depict anatomic structures clearly and that the use of cardiac gating improves image clarity. The technique offers an alternative, non-invasive method of detailed imaging of the oesophagus.

Endoscopic ultrasound is currently used for local staging of oesophageal cancer, but has technical limitations, with a failure rate in up to 17–24% of patients due to stenotic tumours preventing the passage of the endoscope [1, 2]. It is also operator dependent, with a recognized learning curve [3]. The high frequency probe (12 MHz) used to achieve high spatial resolution, has a limited sonographic range and, as a consequence, the relationship of the oesophagus to the surrounding perioesophageal tissues can be difficult to visualize with clarity. There have been few studies that have evaluated MRI of the oesophagus. Early in vivo studies used an integrated body coil and a low field strength magnet (0.35–0.5 T). As a consequence, the images generated had a low in-plane spatial resolution [4, 5]. More recent studies have concentrated on the use of an endoluminal surface coil attached to the tip of a modified endoscope. Preliminary results have been encouraging, with confirmation of the ability to depict the layers of the oesophageal wall [6–8]. However, the technique has recognized limitations, namely: the inability of the coil to traverse strictures, limiting the use for the evaluation of oesophageal cancers; a short radius for receiving signal (3–4 cm), which necessitates repositioning of the endoscope for evaluation of long tumours therefore increasing the overall scan time; and the action of peristalsis, resulting in motion artefact within the images acquired and possible coil migration. The only study to date using a 1.5 T magnet and an external surface coil examined oesophageal specimens in vitro [9]. The study concluded that individual layers of The British Journal of Radiology, November 2006


the oesophageal wall were more clearly depicted on T2 weighted rather than T1 weighted images. The aim of this study was to use healthy volunteers to develop the technique of high resolution thin slice MRI, using an external surface coil in vivo, as a non-invasive method of imaging the oesophagus and perioesophageal tissues; and to both describe the methods used to optimize the signal to noise ratio (SNR) and to assess the effect of cardiac gating.

Methods and materials Four volunteers were imaged using a 1.5 T magnet (Philips Intera, software version 9.5.2, Andover, MA) and an external 5 channel surface coil (Philips Sense CardiacTM). The study evaluated images of the lower oesophagus, as clinically the majority of patients at our institution who are considered for surgical resection are those with adenocarcinoma of the lower oesophagus and gastro-oesophageal junction. The study objectives were to: 1. Maximize spatial resolution by alterations to the feld of view (FOV), matrix size and slice thickness, whilst maintaining acceptable image quality. 2. Maximize SNR by alterations to the echotime (TE) and number of signal averages (NSA), whilst maintaining acceptable image quality and scan time. 3. Minimize the effects of cardiac motion, by the use of a saturation band placed over the heart and also the use of cardiac gating. 873

A M Riddell, C Richardson, E Scurr and G Brown

4. Maintain an acceptable scan time. The maximum limit for an individual sequence was set at 7 min 30 s for 22 axial images. T2 weighted sequences were used as it had previously been established that these are optimum for delineating the layers of the oesophageal wall [9, 10]. A sagittal sequence was performed initially to localize the oesophagus and to ensure optimum coil placement. Axial images were acquired perpendicular to the long axis of the oesophagus. A standard fast spin echo T2 weighted sequence used for pelvic imaging [11] was modified and used as a baseline sequence for the studies without cardiac gating, and the baseline parameters used when cardiac gating was employed were modified from a standard cardiac T2 weighted sequence, with a reduction in the FOV. In both instances, the baseline sequence was given a relative SNR of 1. Changes in the SNR due to alterations in sequence parameters were expressed as a ratio from this baseline.

Assessment of image quality For each of the study objectives, the images acquired using specific sequence parameters were viewed on a workstation (eFilm workstationTM Version 1.9.3) and were scored objectively by one radiologist (AMR) as good, moderate or poor, on the basis of the ability to visualize individual layers of the oesophageal wall, the perioesophageal fat, the azygos vein and wall of the descending aorta. This provided a qualitative evaluation of image clarity. Using the same parameters, the effect of cardiac gating on image quality was assessed by comparing equivalent images with and without cardiac gating.

Figure 1. The margins of acceptable signal to noise, confined within the white lines. A saturation band is placed over the heart (X) to reduce artefact from cardiac motion. The oesophagus is clearly seen anterior to the vertebral column (white arrows).

achieved with an acceptable SNR was obtained using a 256 mm6256 mm matrix, with 240 mm FOV, 3 mm slice thickness and 6 NSA. These parameters gave an in-plane resolution of 0.94 mm60.94 mm and a voxel size of

Results The sagittal T2 weighted sequence provided valuable information to optimize coil placement. Figure 1 illustrates the boundary of acceptable signal and marked signal drop off outside this region. The individual layers of the oesophageal wall could be depicted clearly on the axial high resolution T2 weighted images. The mucosa returned intermediate to low signal. This was surrounded by high signal intensity submucosa and the outer low signal intensity muscularis propria, as shown in Figure 2. The perioesophageal fat returned high signal. The volume of fat varied between subjects and for those subjects with less fat, the clarity of the oesophageal wall layers and perioesophageal fat was graded as moderate, compared with a grading of good for subjects with a greater volume of perioesophageal fat.

Maximizing spatial resolution For the T2 weighted sequence without cardiac gating a relative SNR, when compared with the initial baseline sequence parameters, of less than 0.6 resulted in unacceptable image quality. The highest spatial resolution 874

Figure 2. The main oesophageal wall layers: low signal mucosa (white arrow), surrounded by high signal submucosa (black arrow) and low signal muscularis propria (arrowheads). The descending thoracic aorta (A) and vertebral body (V) are marked.


MRI of the oesophagus Table 1. The optimum sequences for the T2 weighted sequences without and with cardiac gating Sequence type

TR (ms)

TE (ms)

Matrix (mm)

FOV (mm)

NSA

Slice thickness (mm)

SNR

Voxel size (mm3)

No cardiac gating No cardiac gating Cardiac gating

3328 5027 2250

120 80 90

2566256 2566256 3016512

240 225 250

6 6 8

3 3 4

1 1.4 0.44*

2.63 2.32 1.62

*This value for SNR is relative to the baseline for the sequence with cardiac gating. SNR, signal to noise ratio; TR, repetition time; TE, echo time; FOV, field of view; NSA, number of signal averages.

2.63 mm3. A matrix size of 512 mm6256 mm produced higher resolution images, but unacceptable noise within the image. Image quality was also significantly impaired by reducing the FOV to less than 240 mm. The optimum sequence parameters are shown in Table 1. Using cardiac gating, the highest achievable spatial resolution with acceptable signal to noise was achieved by employing a 250 mm FOV, a 301 mm6512 mm matrix and a slice thickness of 4 mm, giving an in-plane spatial resolution of a 0.5 mm60.8 mm and a voxel size of 1.62 mm3, as shown in Table 1.

increased to 9.28 min for 22 slices with an NSA of 8, exceeding the maximum limit. The optimum sequence parameters are shown in Table 1.

Minimizing the effects of cardiac motion

The optimized sequence parameters, without cardiac gating, achieved from the first study were used as the baseline for technique refinement. A decrease in TE resulted in an increase in the relative signal to noise (Figure 3), whist maintaining the T2 weighting. As a consequence, by utilizing a shorter TE, a smaller field of view could be used to improve spatial resolution. Increasing the NSA improved the SNR (Figure 4); an NSA of 6 was considered optimum as the scan time was

The images acquired with a saturation band alone remained subject to motion artefact, primarily due to cardiac pulsation. Image quality was superior for the sequence with cardiac gating, when directly compared with the equivalent image without cardiac gating (Figure 5). Specifically, the clarity of the perioesophageal tissues and wall of the descending thoracic aorta was improved with the use of cardiac gating, with the majority of images with cardiac gating being scored as good quality. Depiction of the individual oesophageal wall layers was also superior with cardiac gating, although the improvement was not as marked. Initially, the RF pulse was set to trigger on every third R wave (3 beats), giving an effective TR of 2250 ms. This gave adequate T2 weighting to the image. Increasing the effective TR to 4000 ms (6 beats) resulted in a reduction in image quality, in particular the clarity of the perioesophageal tissues (Figure 6).

(a)

(b)

Maximizing SNR whilst maintaining spatial resolution

Figure 3. Images at the same level of the oesophagus in one volunteer. The echo time (TE) is reduced from (a) 120 ms to (b) 80 ms. All other parameters are maintained (matrix 256 mm6256 mm, repetition time (TR) 5446 ms, field of view (FOV) 225 mm, number of signal averages (NSA) 6, turbo spin echo (TSE) 16). The signal is increased, improving the conspicuity of the oesophagus (white arrow), the wall of the aorta (arrowhead) and the perioesophageal fat (double arrow heads) with the shorter TE.


875


(a)

(b)

Figure 4. Images at the same level of the oesophagus in one volunteer. The number of signal averages (NSA) is increased from (a) 4 to (b) 8. All other parameters are maintained (matrix 256 mm6256 mm, repetition time (TR) 5446 ms, field of view (FOV) 225 mm, echo time (TE) 80 ms, turbo spin echo (TSE) 16). The higher NSA improves image quality by improving conspicuity of the oesophageal wall layers: the high signal submucosa (white arrow) and lower signal muscularis propria (arrowhead); the perioesophageal tissues (black arrow) and wall of the aorta.

When the T2 weighted sequence without cardiac gating was modified to utilize the sequence parameters optimized with cardiac gating, there was a 37% drop in relative signal to noise and the scan time increased from 7.07 min to 10.0 min for 22 slices, exceeding the acceptable scan time limit.

(a)

Scan duration Using the optimized sequence parameters without cardiac gating, an acceptable scan time of 7.07 min for 22 slices was achieved. The use of cardiac gating did not significantly increase the scan duration. The scan time

(b)

Figure 5. Images taken at the same level of the oesophagus in one volunteer, using the optimized high resolution T2 weighted sequence (a) without and (b) with cardiac gating. The sequence with cardiac gating provides improved image quality. This is illustrated by the clarity of the oesophageal wall submucosa (black arrow), the aortic wall (white arrow) and right pleural reflection (fine black arrow). 876


MRI of the oesophagus

(a)

(b)

Figure 6. Images taken at the same level of the oesophagus in one volunteer. Increasing the repetition time (TR) from (a) 3 beats (effective TR 2250 ms) to (b) 6 beats (effective TR 4000 ms) caused blurring within the image. The oesophageal wall submucosa (black arrowhead), muscularis propria (black arrow), and right pleural reflection (white arrow) are more clearly seen with an effective TR of 2250 ms (3 beats).

using cardiac gating varied slightly between volunteers due to differences in resting heart rate, but for the optimized sequence parameters did not exceed the scan time limit.

Discussion High resolution thin slice MRI of the oesophagus using an external surface coil is potentially challenging due to the relatively low signal to noise returned from the structure due to its small size and location within the thorax. This study shows that the individual layers of the oesophageal wall and surrounding anatomical structures can be delineated using the technique. The signal returned from the individual wall layers in this study is in accordance with the description given by Yamada et al [9, 10]. Endoscopic ultrasound can also delineate the layers of the oesophageal wall. However, the ability of the technique to delineate the perioesophageal tissues is limited by the relatively short ultrasound range of the probe (in the order of 3–4 cm). The axial images acquired of the posterior mediastinum obtained using high resolution MRI clearly delineate the relationship of the oesophagus to all its surrounding anatomical structures, providing improved detail regarding potential resection margins for those patients with oesophageal carcinoma who are being considered for surgery. The variation in the amount of perioesophageal fat between subjects resulted in a reduction in image clarity in subjects with a small amount of perioesophageal fat, which could be considered a limitation of the technique. The British Journal of Radiology, November 2006

The non-invasive nature of the technique will allow optimum visualization of stenotic tumours, not amenable to endoscopic evaluation. In many instances these bulky tumours are not considered surgically resectable. However, a technique which allows for detailed initial evaluation of the local extent of tumour will provide the most accurate method of local staging of the tumour. It would also provide a basis for alternative treatment planning and a baseline for future imaging to fully assess treatment response. The perioesophageal tissues are visualized with CT, but recent studies have shown that the ability of the technique to determine resectability is relatively low. Markland et al reported a sensitivity of between 0% and 66.7% [12] and a further study of 51 patients showed that 59% had more advanced disease than was appreciated on pre-operative CT, reinforcing the need to refine the techniques for the local staging of oesophageal carcinoma to improve patient selection for radical therapy [13]. The superior contrast resolution achieved with MRI would suggest that the technique is likely to be useful in differentiating tumour abutting surrounding structures (stage T3) from direct invasion of tumour into the surrounding tissues (stage T4).

Spatial resolution Previous attempts to achieve high resolution, thin slice imaging of the oesophagus have been limited to in vitro studies of resected oesophageal specimens from patients with known oesophageal carcinoma [9, 10]. The studies 877


used a 4 cm loop surface coil and a 1.5 T or 4.5 T magnet, with FOV of between 40 mm and 60 mm, and as such were able to achieve a voxel size of 0.08–0.16 mm3. Given this high resolution, 6–8 layers of the oesophageal wall were visualized and the technique was able to identify tumour which was confined to the mucosa. Prior in vivo studies using surface coils used low field strength magnets (0.35–0.5 T) and acquired images with a slice thickness of 8–10 mm, to achieve adequate signal to noise. As a consequence, these studies concluded that MRI was no better than CT at providing detailed analysis of the oesophageal wall and surrounding anatomical structures [4, 5]. Investigations using an endoluminal coil have achieved an in plane resolution of 0.469 mm 60.625 mm (voxel size 1.17 mm3), enabling visualization of the three main mural layers (mucosa, submucosa and muscularis) [14]. Our study shows that using fast spin echo sequences, high resolution images can be acquired with an external surface coil, with a voxel size of 1.62 mm3. At this resolution the main component layers of the oesophageal wall were readily delineated, indicating that it would be possible to use the technique for the local staging of oesophageal tumours.

Scan time Our experience is that, in general, patients are able to lie still for an individual sequence which lasts for between 7 min and 7K min. Our study used 7K min as the maximum acceptable scan time for 22 axial images. Although image quality was improved by increasing the NSA from 6 to 8, for the T2 weighted sequence without cardiac gating, this exceeded our upper time limit. These scan times compare favourably with endoluminal MRI studies, which have documented scan times of between 7 min and 10 min for 8–12 images. No previous study has specifically assessed the affect of cardiac gating on scan time; these results show that its use does not significantly increase the overall scan duration. We conclude that our study indicates that the technique of high resolution, thin slice MRI using and external surface coil is a reproducible, non-invasive method of imaging the oesophagus, which could potentially provide an alternative method for the local staging of oesophageal carcinoma.

References Signal to noise The benefit of using an endoluminal surface coil is that the coil is closer to the region of interest, thus improving signal to noise. As a consequence, the spatial resolution can be maximized. There is, however, rapid signal decay beyond a 3–4 cm radius of the endoluminal coil and therefore the technique has limited ability to assess the perioesophageal tissues and determine tumour resectability. Previous MRI studies with an external coil used a body coil integrated into the bore of the magnet. The introduction of multi-channel external surface phased array coils, as used in our study, has improved the achievable SNR, by the use of parallel imaging techniques. The technology allows for faster sequences enabling higher resolution imaging to be acquired with increased signal to noise.

Cardiac gating Several of the previous studies using endoluminal and external surface coils have used cardiac gating, but no previous study has directly compared the effect of using the technique on image quality. The acquisition of imaging data when using cardiac gating is synchronized with and limited to specific times or phases of the cardiac cycle. The R wave on the electrocardiogram (ECG) is used to trigger each radiofrequency (RF) pulse. The RF signal is therefore generated at the same point of subsequent cardiac cycles. Since the R to R interval is controlled by the patient’s heart rate – the TR (and as a consequence the image weighting and number of slices) depends on the heart rate. To enable variation in the TR value, the system can be adapted to trigger an RF pulse at every second or third R wave; in this way the ‘‘effective TR’’ is lengthened, allowing for T2 weighted imaging. Our study would indicate that the use of cardiac gating does improve image quality. 878

1. Wakelin SJ, Deans C, Crofts TJ, Allan PL, Plevris JN, Paterson-Brown S. A comparison of computerised tomography, laparoscopic ultrasound and endoscopic ultrasound in the preoperative staging of oesophago-gastric carcinoma. Eur J Radiol 2002;41:161–7. 2. Preston SR, Clark GW, Martin IG, Ling HM, Harris KM. Effect of endoscopic ultrasonography on the management of 100 consecutive patients with oesophageal and junctional carcinoma. Br J Surg 2003;90:1220–4. 3. Schlick T, Heintz A, Junginger T. The examiner’s learning effect and its influence on the quality of endoscopic ultrasonography in carcinoma of the esophagus and gastric cardia. Surg Endosc 1999;13:894–8. 4. Quint LE, Glazer GM, Orringer MB. Esophageal imaging by MR and CT: study of normal anatomy and neoplasms. Radiology 1985;156:727–31. 5. Lehr L, Rupp N, Siewert JR. Assessment of resectability of esophageal cancer by computed tomography and magnetic resonance imaging. Surgery 1988;103:344–50. 6. Inui K, Nakazawa S, Yoshino J, Yamao K, Yamachika H, Wakabayashi T, et al. Endoscopic MRI: preliminary results of a new technique for visualization and staging of gastrointestinal tumors. Endoscopy 1995;27:480–5. 7. Kulling D, Feldman DR, Kay CL, Hoffman BJ, Reed CE, Young JW, et al. Local staging of esophageal cancer using endoscopic magnetic resonance imaging: prospective comparison with endoscopic ultrasound. Endoscopy 1998;30: 745–9. 8. Dave UR, Williams AD, Wilson JA, Amin Z, Gilderdale DJ, Larkman DJ, et al. Esophageal cancer staging with endoscopic MR imaging: pilot study. Radiology 2004;230: 281–6. 9. Yamada I, Izumi Y, Kawano T, Yoshino N, Tetsumura A, Ohashi K, et al. Superficial esophageal carcinoma: an in vitro study of high-resolution MR imaging at 1.5T. J Magn Reson Imaging 2001;13:225–31. 10. Yamada I, Murata Y, Izumi Y, Kawano T, Endo M, Kuroiwa T, et al. Staging of esophageal carcinoma in vitro with 4.7-T MR imaging. Radiology 1997;204:521–6. 11. Brown G, Daniels IR, Richardson C, Revell P, Peppercorn D, Bourne M. Techniques and trouble-shooting in high spatial


MRI of the oesophagus resolution thin slice MRI for rectal cancer. Br J Radiol 2005;78:245–51. 12. Markland CG, Manhire A, Davies P, Beggs D, Morgan WE, Salama FD. The role of computed tomography in assessing the operability of oesophageal carcinoma. Eur J Cardiothorac Surg 1989;3:33–6.


13. Sariego J, Mosher S, Byrd M, Matsumoto T, Kerstein M. Prediction of outcome in ‘‘respectable’’ esophageal carcinoma. J Surg Oncol 1993;54:223–5. 14. Feldman DR, Kulling DP, Hawes RH, Kay CL, Muckenfuss VR, Cotton PB, et al. MR endoscopy: preliminary experience in human trials. Radiology 1997;202:868–70.

879


Pancreatic adenocarcinoma: signs of vascular invasion determined by multi-detector row CT 1

H LI,

MD,

2

M S ZENG,

MD, Pro,

2

K R ZHOU,

MD, Pro,

3

D Y JIN,

MD, Pro

and 3W H LOU,

MD, Pro

1

Department of Radiology, The Affiliated First People’s Hospital, Shanghai Jiao Tong University, 85 Wujin Road, 200080, Shanghai, Departments of 2Radiology and 3General Surgery, The Affiliated Zhongshan Hospital, Fudan University, 180 Fenglin Road, 200032, Shanghai, P R China ABSTRACT. The purpose of this study was to analyse multi-detector row CT (MDCT) signs of peripancreatic arterial and venous invasion in pancreatic carcinoma. Among 101 patients with pancreatic carcinoma examined by MDCT, 54 candidates for surgery were pre-operatively evaluated for vascular invasion based on MDCT signs. The peripancreatic major vessels (including superior mesenteric artery, coeliac artery, common hepatic artery, superior mesenteric vein and portal vein) were examined carefully by surgeons during the operation. At surgical exploration, 78 of 224 vessels were invaded by tumour. The invaded peripancreatic major arteries (n529) and veins (n549) presented different MDCT signs: 43% of invaded veins (18/42, except for 7 occluded veins) were surrounded by tumour less than 50% of the vessel circumference compared with 97% (28/29) of the invaded arteries, which were surrounded by tumour more than 50% of the vessel circumference or were embedded in tumour (p,0.001). 69% (34/49) of the invaded veins had vascular stenosis or obliteration, compared with 41% (12/29) of the invaded arteries (p,0.05). Irregularity of the vein wall, 74% (31/42, except for 7 occluded veins); occurred more often than that of the artery wall, 45% (13/29) (p,0.05). In conclusion, the MDCT signs of peripancreatic arterial and venous invasion have different characteristics, which should be considered in pre-operative evaluation.

One decisive factor for the irresectability of pancreatic carcinoma is the major vascular invasion [1]. Multidetector row CT (MDCT) scanning of the pancreas enables multiphasic thin collimation scanning, with excellent spatial resolution, especially in the z-plane [2]. These volume data sets can be easily manipulated with three-dimensional imaging, potentially providing additional information to conventional axial display [3–6]. The purpose of our study was to assess the MDCT signs of arterial and venous invasion in pancreatic carcinoma.

Received 28 July 2005 Revised 17 November 2005 Accepted 2 May 2006 DOI: 10.1259/bjr/19684199 ’ 2006 The British Institute of Radiology

MDCT images of the 101 patients were all prospectively evaluated for resectability: including vascular invasion and the presence of metastatic disease. Of the 54 candidates for surgery (37 males and 17 females, 40– 79 years old, average 61.2 years), 17 patients underwent pancreaticoduodenectomy, and the others were found to be irresectable during surgical exploration. The time interval between MDCT examination and surgery was #2 weeks.

Examination techniques

Materials and methods Patients Between December 2001 and February 2004, 255 consecutive patients with presumed pancreatic carcinoma underwent pancreas examinations by triphasic MDCT. 101 patients with pancreatic carcinoma were confirmed by surgical-pathology (n554) or clinical follow-up (n547). 54 patients were pre-operatively judged to be resectable or requiring surgical bypass due to jaundice. Pancreatic adenocarcinoma was pathologically confirmed by biopsy or fine-needle aspiration. The remaining 47 patients deemed irresectable received chemotherapy or radiotherapy, and all of them survived less than 2 years. 880

All MDCT examinations of the pancreas were performed on Mx8000 (Philips, Best, The Netherlands; fourslice spiral CT scanners). 600–800 ml water was routinely administered 10–20 min before the examination to distend the stomach, duodenum and proximal jejunum. Each patient received 120 ml of non-ionic contrast material (iopromide 350 mg ml21; Ultravist; Schering, Berlin, Germany) via intravenous injection at the rate of 5 ml s21. Unenhanced and triphasic (arterial phase, pancreatic phase and hepatic phase) enhanced scans were performed. Unenhanced scan images were obtained from the top of the diaphragm to the caudal level of the uncinate process. The scanning parameters were: 120 kVp, 250 mA, a 0.875 pitch and 5 mm collimation. After a 20 s delay from the start of the intravenous The British Journal of Radiology, November 2006

Signs of vascular invasion

infusion, the arterial phase was obtained from the level of the hepatic hilum through the entire pancreas. The collimation was altered to 2.5 mm, and the other parameters were the same as those of the unenhanced scan. The delay time of the pancreatic phase and the hepatic phase was 45 s and 80 s, respectively. The scanning range and parameters were the same as those of the arterial phase and the unenhanced scans, respectively. After scanning, the data were transferred to the image server linked to an Mxview workstation with software version 3.5. This workstation was used for threedimensional analysis of the local anatomy utilizing volume rendering (VR) combined with maximum intensity projection (MIP) and multiplanar reconstruction (MPR).

Image analysis and surgical correlation CT signs of peripancreatic major vessels (including coeliac artery (CA); common hepatic artery (CHA); superior mesenteric artery (SMA); portal vein (PV); superior mesenteric vein (SMV)) were determined preoperatively at consensus reading of the axial and 3D images by two observers (MSZ and HL, 10 years and 5 years of experience reading pancreatic imaging, respectively):

224 vessels were examined carefully at surgery by two experienced surgeons (DYJ, WHL, 12 years and 8 years of experience operating pancreatic carcinoma, respectively) together. The surgical criterion of vascular ingrowth was that the vessel could be observed, or found by palpation, to be infiltrated or occluded at surgery. It was not considered vascular invasion if tumour was adherent to but could be separated from the vessel due to inflammatory or fibrotic reaction.

Statistical analysis Chi-square tests were performed on the CT signs: (a) tumour surrounding more than 50% of the vessel circumference; (b) vessel stenosis or occlusion; (c) vessel wall irregularity, to check for any significant difference between the invaded arteries and veins. p , 0.05 was required to show statistical significance.

Results

1. Contiguity of tumour with the adjacent vessel was graded A,D: (grade A, fat plane or normal pancreatic tissue visible between tumour and vessel; grade B and grade C, tumour surrounding of less than and more than 50% of the vessel circumference; grade D, arterial embedment in tumour or venous occlusion.) 2. Detailed vascular anatomic deformation: vessel stenosis presented a semi-circular or concentric smaller contour of the vessel. But a straight contour on one side was regarded as flattened, not as stenosed. Vessel wall infiltration presented an irregular and indented shape at the vascular margin abutting tumour.

43 of the 54 tumours at surgery were located in the pancreatic head or uncinate process, 11 in the body or tail. 17 tumours were surgically resectable, of which 16 were pre-operatively correctly diagnosed by MDCT. One was incorrectly judged irresectable because MDCT had shown tumour ingrown in the SMA (Figure 1). Of 37 patients who received by-pass palliative surgery or surgical exploration, 6 tumours were incorrectly diagnosed resectable due to underdiagnosed invaded vessels (n54), missed hepatic metastases (n51) and peritoneal metastases (n51). Of the 224 observed vessels, 146 vessels were found not to be invaded at surgery and 78 invaded. Except for the above-mentioned SMA, which MDCT incorrectly judged to be locally infiltrated by tumour on MDCT, the calibre of the other 145 not invaded vessels was unchanged and the vessel wall was regular. On MDCT, the correlation between the surgically confirmed vessels that were not invaded and tumour was:

(a)

(b)

Figure 1. Pancreatic body carcinoma. (a,b) Consecutive axial images showed that the superior mesenteric artery (SMA) was surrounded to more than 50% of the vessel circumference by tumour and the vessel wall appeared infiltrated. Pre-operatively, this case was judged irresectable. At surgical exploration, the SMA was found not to be invaded (only fibrotic infiltration) by tumour and was successfully resected.


881

H Li, M S Zeng, K R Zhou et al

(a)

(b)

Figure 2. Tumour in the uncinate process of the head of the pancreas. (a,b) Consecutive axial images showed the superior mesenteric artery (SMA) to be surrounded more than 50% of the vessel circumference by tumour. It was found not to be invaded by tumour at exploration and tumour was successfully resected.

1. Grade A of contiguity of tumour with the adjacent vessel (fat plane or normal pancreatic tissue preserved) (n5129, 22 SMAs, 29 CAs, 30 CHAs, 18 SMVs and 30 PVs) 2. Grade B of contiguity of tumour with the adjacent vessel (,180 ˚) (n59, 3 SMAs, 1 CA, 3 SMVs, 2 PVs) 3. Grade C of contiguity of tumour with the adjacent vessels (.180 ˚) (n58, 5 SMAs, 2 CAs, 1 CHA) (Figure 2)

The MDCT appearances of surgically confirmed invaded arteries (n529) and veins (n549) are shown in Tables 1 and 2, respectively. Statistical analysis related to invaded arteries and veins is shown in Table 3.

Circumferential involvement 43% of invaded veins (18/42, except for 7 occluded veins) were surrounded by less than 50% of the vessel

Table 1. Multi-detector CT (MDCT) appearances of surgically confirmed invaded arteries

Grade Grade Grade Grade Grade Grade Grade Grade Grade

D & irregular wall & stenosed vessel D & irregular wall & unchanged calibre D & regular wall & stenosed vessel D & regular wall & unchanged calibre C & irregular wall & stenosed vessel C & irregular wall & unchanged calibre C & regular wall & stenosed vessel C & regular wall & unchanged calibre B & regular wall & unchanged calibre

SMA

CA

CHA

4 2 2 2 2 1 – 4 1

1 – 1 2 – 1 – – –

1 1 1 2 – – – 1 –

SMV

PV

4 11 2 1 – 4 2 1 3 2 –

3 6 3 1 – 2 1 1 – 1 1

Grade A,D: the grade of contiguity of tumour with the adjacent vessel. SMA, superior mesenteric artery; CA, coeliac artery; CHA, common hepatic artery.

Table 2. Multi-detector CT (MDCT) appearances of surgically confirmed invaded veins

Grade Grade Grade Grade Grade Grade Grade Grade Grade Grade Grade

D C & irregular wall & stenosed vessel C & irregular wall & unchanged calibre C & regular wall & stenosed vessel C & regular wall & unchanged calibre B & irregular wall & stenosed vessel B & irregular wall & unchanged calibre B & regular wall & stenosed vessel B & regular wall & calibre was ‘‘teardrop’’ B & regular wall & unchanged calibre A & regular wall & unchanged calibre

Grade A,D: the grade of contiguity of tumour with the adjacent vessel. SMV, superior mesenteric vein; PV, portal vein.

882


Signs of vascular invasion Table 3. Statistical analysis between invaded arteries and veins CT signs

1* 2 3*

Artery

Vein

CA

CHA

SMA

Total

SMV

PV

Total

5/5 2/5 2/5

6/6 2/6 2/6

17/18 8/18 9/18

28/29 12/29 13/29

14/26 21/30 19/26

10/16 13/19 12/16

24/42 34/49 31/42

Chi2

p-value

13.6 5.9 6.1

,0.001 ,0.05 ,0.05

1: Tumour surrounding more than 50% of the vessel circumference. 2: Vessel stenosed or occlusive. 3: Vessel wall irregularity. *Seven veins (4 SMVs, 3 PVs) occluded by tumours were not included. CA, coeliac artery; CHA, common hepatic artery; SMA, superior mesenteric artery; SMV, superior mesenteric vein; PV, portal vein.

circumference by tumour. 97% (28/29) of invaded arteries were surrounded by more than 50% of the vessel circumference by tumour or were embedded in tumour. Statistical analysis showed a significant difference between invaded arteries and veins (except for seven occluded veins) utilizing circumferential involvement of more than 50% of the vessel circumference by tumour (p,0.001).

Stenosis or occlusion 69% (34/49) of invaded veins presented stenosed or occluded (Figure 3). Eight veins (5 SMVs, 3 PVs), which were surrounded less than 50% of the vessel circumference by tumour, also appeared stenosed (Figure 4). 41% (12/29) of invaded arteries did not appear stenosed, and the difference was statistically significant

(a)

(b)

(c)

(d)

Figure 3. Tumour in the uncinate process of the head of the pancreas. (a,b) Axial images and (c,d) 3D images showed that the tumour in the uncinate process of the head of the pancreas eroded a side of the superior mesenteric vein (SMV) vessel wall and penetrated it to form tumour thrombus (arrow). Axial images also showed that the superior mesenteric artery (SMA) was embedded. Surgery confirmed tumour invasion of the SMV and the SMA.


883


(a)

(b)

(c)

884

Figure 4. Pancreatic head carcinoma. (a,b) Consecutive axial images showed the portal vein (PV) vessel calibre (arrow) was changed, although it was surrounded less than 50% of the vessel circumference by tumour. (c) Volume rendering (VR) 3D image showed that a segment of PV was stenosed. The axial images also showed that the common hepatic artery (CHA) was embedded in tumour and the coeliac artery (CA) was surrounded by more than 50% of the vessel circumference. PV, CHA and CA were proven to be infiltrated at surgical exploration.



(a) Figure 6. Pancreatic head carcinoma. The confluence of the portal vein (PV) and the superior mesenteric vein (SMV) was shown to be surrounded by less than 50% of the vessel circumference by tumour, with unchanged vessel calibre. But a focal area of vessel wall (arrow) was irregular and infiltrated. The confluence of the PV and the SMV was proven to be infiltrated at surgical exploration.

Infiltration

(b)

74% of invaded veins (31/42, except for 7 occluded veins) appeared infiltrated with an irregular wall. Three veins (2 SMVs, 1 PV) were found to be invaded only from their irregular wall, were surrounded by less than 50% of the vessel circumference and were not stenosed (Figure 6). 45% (13/29) of invaded arteries appeared infiltrated less often than invaded veins (except for 7 occluded veins) (p,0.05).

Other MDCT appearances of arterial and venous invasion 24% (7/29) of invaded arteries appeared deformed and seemed to be stretched or stiffened in threedimension reconstruction images, such as VR (Figure 7). 10% (3/30) of invaded SMVs were diagnosed correctly because they appeared as ‘‘teardrop’’ in axial images (Figure 8). When invaded veins were highly stenosed or occluded, 20% (10/49), there were multiple venous collaterals. (c) Figure 5. Tumour in the uncinate process of the head of the pancreas. (a,b) Axial images showed coeliac artery (CA) (short arrow) and common hepatic artery (CHA) (long arrow) embedded in tumour with regular vessel walls, although (c) volume rendering (VR) 3D image showed that the vessel calibre of CA and CHA was unchanged. Surgery confirmed tumour invasion of CA and CHA.

(p,0.05). Nine arteries (2 CAs, 3 CHAs, 4 SMAs), embedded in tumour, did not appear stenosed (Figure 5). The British Journal of Radiology, November 2006

Discussion According to previous reports, using single-detector row CT the diagnostic accuracy of pancreatic irresectability was 95% and of pancreatic resectability was 70,80% [7–10]. The most important reason for inaccurate assessment of resectability is underestimation of the vascular invasion. To improve the accuracy of estimating invaded vessels, it is necessary to evaluate the MDCT signs of arterial and venous invasion separately because we have found clinically that the features of peripancreatic arterial and venous invasion on CT are different. 885


(a) Figure 8. Tumour (T) in the uncinate process of the head of the pancreas tethered the superior mesenteric vein (SMV) (arrow) into a teardrop shape, although the SMV was surrounded by less than 50% of the vessel circumference and vessel wall remained regular. The superior mesenteric artery (SMA) (arrowhead) was also shown to be nearly embedded in tumour. The SMV and the SMA were proven to be infiltrated at surgical exploration.

(b)

(c) Figure 7. Pancreatic head carcinoma. (a,b) Consecutive axial images showed that the superior mesenteric artery (SMA) was embedded in tumour. (c) Volume rendering (VR) 3D image showed a segment of the SMA to be stiffened (arrow) where the vessel calibre was low-grade stenosed. Surgery confirmed tumour invasion.

In our opinion, the major reason for the different CT signs of arterial and venous invasion is that the vein wall is thinner and weaker than the artery wall. When veins are surrounded and infiltrated by tumour, the wall tends to be irregular and the calibre becomes narrowed. At the same time, the flow rate in veins is slow, and tumour often penetrates the vein wall to form tumour thrombus, causing vascular occlusion (Figure 3). As the artery wall is thicker and more flexible than the vein wall, and the artery calibre is smaller, even when the arteries are 886

embedded in tumour, the calibre is not easily changed and the wall remains regular. In our study, some arteries were found not to be invaded at surgical exploration, although they were surrounded by more than 50% of the vessel circumference, and their calibre remained unchanged and the wall was regular. Hough et al [11] found that tumours in the head of the pancreas could cause a tethered, teardrop appearance of the SMV in axial images. They believed that the teardrop SMV sign was a reliable indicator of irresectability. In their retrospective study, teardrop SMV was the only sign of irresectability in 13 of 17 patients. This sign presumably results from either direct tumour infiltration or peritumoural fibrosis adherent to the vessel that retracts or tethers the vessel, changing its normal round shape. In our series, there were three invaded SMVs confirmed by surgical exploration exhibiting the teardrop sign on axial images in which the vessels were surrounded by less than 50% of the vessel circumference, and the wall of which was regular. Since pancreatic carcinoma may be accompanied by focal tissue fibrosis, the invaded arteries may appear stretched. In our study, seven stretched arteries were appreciated on three-dimensional reconstruction MDCT angiography (MDCTA) images [3]. Lepanto et al [12] found that CTA significantly increased the ability to identify venous invasion but did not improve the detection of arterial invasion. There are few MDCTA cases reported. As our experience has shown, the image quality of MDCTA exceeds that of CTA. MDCTA in particular allows the three-dimensional display of the relationship between tumour and vessel, but conventional axial images should also be reviewed as axial images appear better able to demonstrate the contiguity of tumour to vessel and change in calibre of the vessel wall. A major limitation of our study is our gold standard of surgical palpation, because irresectable tumours The British Journal of Radiology, November 2006


could not be histologically assessed, although the surgical margins of resected tumours were confirmed by pathology. In conclusion, invaded peripancreatic arterial and venous MDCT signs are different. It is important to pay attention to these differences in order to improve the accuracy of diagnosing vascular invasion and pancreatic resectability.

References 1. Hommeyer SC, Freeny PC, Crabo LG. Carcinoma of the head of the pancreas: evaluation of the pancreaticoduodenal veins with dynamic CT-potential for improved accuracy in staging. Radiology 1995;196:233–8. 2. McNulty NJ, Francis IR, Platt JF, Cohan RH, Korobkin M, Gebremariam A. Multi-detector row helical CT of the pancreas: effect of contrast-enhanced multiphasic imaging on enhancement of the pancreas, peripancreatic vasculature, and pancreatic adenocarcinoma. Radiology 2001;220:97–102. 3. Horton KM, Fishman EK. Multidetector CT angiography of pancreatic carcinoma: part 1, evaluation of arterial involvement. AJR Am J Roentgenol 2002;178:827–31. 4. Horton KM, Fishman EK. Multidetector CT angiography of pancreatic carcinoma: part 2, evaluation of venous involvement. AJR Am J Roentgenol 2002;178:833–6. 5. Horton KM, Fishman EK. 3D CT angiography of the celiac and superior mesenteric arteries with multidetector CT data


6.

7.

8.

9.

10.

11.

12.

sets: preliminary observations. Abdom Imaging 2000;25:523–5. Nino-Murcia M, Jeffrey RB, Beaulieu CF, Li KCP, Rubin GD. Multidetector CT of the pancreas and bile duct system: value of curved planar reformations. AJR Am J Roentgenol 2001;176:689–93. Hommeyer SC, Freeny PC, Crabo LG. Carcinoma of the head of the pancreas: evaluation of the pancreaticoduodenal veins with dynamic CT-potential for improved accuracy in staging. Radiology 1995;196:233–8. Raptopoulos V, Steer ML, Sheiman RG, Vrachliotis TG, Gougoutas CA, Movson JS. The use of helical CT and CT angiography to predict vascular involvement from pancreatic cancer: correlation with findings at surgery. AJR Am J Roentgenol 1997;168:971–7. Diehl SJ, Lehmann KJ, Sadick M, Lachmann R, Georgi M. Pancreatic cancer: value of dual-phase helical CT in assessing resectability. Radiology 1998;206:373–8. Kaneko K, Honda H, Hayashi T, Fukuya T, Irie H, Masuda K. Helical CT evaluation of arterial invasion in pancreatic tumors: comparison with angiography. Abdom Imaging 1997;22:204–7. Hough TJ, Raptopoulos V, Siewert B, Matthews JB. Teardrop superior mesenteric vein: CT sign for unresectable carcinoma of the pancreas. AJR Am J Roentgenol 1999;173:1509–12. Lepanto L, Arzoumanian Y, Gianfelice D, Perreault P, Dagenais M, Lapointe R, et al. Helical CT with CT angiography in assessing periampullary neoplasms: identification of vascular invasion. Radiology 2002;222:347–52.

887


Conversion factor for CT dosimetry to assess patient dose using a 256-slice CT scanner S MORI,

PhD, MPH, RT,

K NISHIZAWA,

PhD, MPH,

M OHNO,

MSc, MPH

and M ENDO,

PhD, MPH

Department of Medical Physics, National Institute of Radiological Sciences, Chiba 263-8555, Japan ABSTRACT. Recent rapid progress in CT technology has yielded equipment with large numbers of detector rows and standard computed tomography dose index (CTDI) is therefore no longer an adequate integration range. An integration range of 300 mm is necessary to accurately measure dose under a nominal beam width of 128 mm due to scattered radiation. However, such a long phantom is inconvenient to use routinely in cone-beam CT patient dose checking. To assess patient dose accurately with standard dosimetry methods, we determined a conversion factor (CF) which was calculated from the weighted dose profile integral (DPIw) for the 300 mm integration range with a 300 mm long CTDI phantom using a 300 mm long ionization chamber divided by that for the 100 mm integration range with a standard CTDI phantom (140 mm long) with a 100 mm long chamber. CF values increase with increasing nominal beam width and effective energy in the range from 1.5 to 2.0. CF values can also be adapted for use with other CT systems as their dose profiles are thought to be analogous to those for the 300 mm phantom and are useful in any hospital situation to assess accurate patient doses using standard dosimetry methods.

State-of-the-art commercially based CT represents a marked improvement over conventional multislice CT (MSCT), especially in cardiac imaging. Its craniocaudal coverage without gantry movement, however, is typically only 20–40 mm and this has limited the width of coverage for cine imaging in the craniocaudal direction. To overcome this disadvantage, we developed a 256-slice CT scanner. Volumetric cine imaging is realised by scanning continuously at the same position (without table movement); this provides a large amount of diagnostic information and solves some of the limitations of present helical CT methods [1–5]. However, since the maximum nominal beam width is 128 mm, which is three or four times larger than the latest MSCT in common use, the dose is increased in proportion to the scan time. Therefore, it is very important to assess the dose for volumetric cine imaging. For conventional CT dose measurement, a standard computed tomography dose index (CTDI) phantom (140 mm long) [6] and 100 mm long ionization chamber have generally been used. However, the conventional CT dose measurement is not sufficient in 256-slice CT because its beam width is larger than the 100 mm long ionization chamber. We therefore extended the FDArecommended CTDI phantom [6] to a length of 300 mm, on the basis of our group’s previous report that phantom length and integration range for dosimetry needed to be at least 300 mm to represent more than 90% of the line integral dose with a beam width between 20 mm and 138 mm [7]. Since a 300 mm CTDI phantom is quite inconvenient, determining a conversion factor (CF) is useful for all hospitals to assess patient dose accurately using standard dosimetry methods. 888

Received 14 November 2005 Revised 14 March 2006 Accepted 14 March 2006 DOI: 10.1259/bjr/66519303 ’ 2006 The British Institute of Radiology

We propose CF values obtained by calculating ratios of the dose profile integrals (DPIs) with the standard CTDI and 300 mm long phantoms.

Materials and methods Second model of the 256-slice CT-scanner The second model of our 256-slice CT scanner was based on the design of the first model [8, 9] which used a wide-area cylindrical 2D detector incorporating current CT technology. The second model was mounted on the gantry frame of a 16-slice CT [10] (Aquilion; Toshiba Medical Systems, Otawara, Tochigi). The 256-slice CT has 912 (transverse) 6 256 (craniocaudal) elements, each approximately 0.5 mm60.5 mm at the centre of rotation. The 128 mm total beam width allows the continuous use of several collimation sets (e.g. 25660.5 mm, 12861.0 mm, 6462.0 mm). Large and small filters are shaped to compensate for the variable path length of each patient across the scan field of view (FOV). The small filter is used for an object under 240 mm field of view (FOV), and the large filter is used for over 240 mmFOV (e.g. chest and abdomen). The second model of the 256-slice CT scanner incorporates several improvements over the first model, including the acceleration of rotation time from 1.0 s to 0.5 s per rotation, and elongation of the detector dynamic range from 16 to 18 bits [11]. The detector element consists of a Gd2O2S ceramic scintillator and single-crystal silicon photodiode, as used in conventional multislice CT. Maximum X-ray exposure time per scan is 60 s and exposure can be The British Journal of Radiology, November 2006

Conversion factor using 256-slice CT

repeated in a series of scans according to a predetermined program. The x–y coordinate plane is parallel to the transverse direction, and the z-coordinate axis is parallel to the craniocaudal direction.

Phantoms The length of the FDA-recommended CTDI phantom [6] is at least 140 mm. This conventional phantom contains holes just large enough to accept a pencilshaped ionization chamber. For cone-beam CT dose measurement, the phantom length should be longer, because the nominal beam width of 128 mm is longer than that of the conventional CTDI dosimetry method. Therefore, a CTDI phantom of 300 mm length designed for cone-beam CT (Kyoto-kagaku, Kyoto, Japan) [7] was used. The phantoms, 140 mm and 300 mm long, were made of PMMA (polymethylmethacrylate) with diameters of 160 mm for head examinations and 320 mm for body examinations. Holes of 10 mm diameter for the pencil-shaped ionization chamber were located parallel to the rotation axis, and the centres of the holes were located at the cylinder centre and also 10 mm below the cylinder surface at 90 ˚ intervals.

and body phantoms, respectively. Average effective energy (Eave) was estimated using the following equation: 1 2 Eave ~ Ec z Ep 3 3

where Ec and Ep were defined as the effective energy at the centre and periphery, respectively.

Dose profile integral (DPI) The DPIs were measured with the 300 mm long chamber for the 300 mm long phantoms (extended dosimetry) and with the 100 mm long chamber for the 140 mm long phantoms. The DPI was given as the output of the ionization chamber. The weighted DPI (DPIw) for x, y coordinates was given by: 1 2 DPIL,w ðvÞ~ DPIL,c ðvÞz DPIL,p ðvÞ mGy:mm 3 3

Effective energy Effective energy was calculated from the attenuation curve of X-ray intensity. The 300 mm-length ionization chamber was positioned at the horizontal plane that passed through the rotational axis centre plane. The Xray tube was fixed under the ionization chamber, and Xray irradiation was initiated. The X-ray intensity was measured by setting aluminium attenuators of various thicknesses between the X-ray tube and the ionization chamber [12]. The effective energy was derived from an attenuation length that gave half of the X-ray intensity produced without aluminium attenuators (half-value layer (HVL)), calculated from the attenuation curve of the beam intensity [13]. Effective energy was measured at 0 mm, 50 mm, 100 mm and 150 mm along the transverse direction for 80 kV, 100 kV, 120 kV and 135 kV with the nominal beam width of 128 mm. The effective energy at the periphery was averaged from x50 mm to x5100 mm for the small filter and from x50 mm to x5150 mm for the large filters. Those measurement ranges almost completely covered the head The British Journal of Radiology, November 2006

ð2Þ

where DPIc and DPIp denote DPIs in the measurement range L at the centre and periphery of the phantom. v denotes the beam width. Then we calculated the DPIw ratio as the CF value:

Dosemeter probe A pencil-shaped ionization chamber with an active length of 100 mm (CT-10; Applied Engineering Inc., Tokyo, Japan) or 300 mm (CT-30; Applied Engineering Inc., Tokyo, Japan) was connected to a dosemeter (AE132; Applied Engineering Inc., Tokyo, Japan). The 300 mm long chamber was an extended form of the 100 mm long pencil-shaped ionization chamber. The dosemeter was calibrated at NMIJ (National Metrology Institute of Japan) for the appropriate radiation qualities by comparison with the secondary radiation standards.

ð1Þ

CF~

DPI300,w DPI100,w

ð3Þ

Again, we omitted the variable v from the notation of DPI. Using the above equations we calculated the CF for various beam sizes. The phantom was placed on the patient table and its centre was aligned at the isocentre. An ionization chamber was inserted into either the central or one of the peripheral cavities of the phantom (other cavities were filled with PMMA rods). The exposure (expressed as C kg21) was obtained with the ionization chamber dosemeter and converted to the values of the absorbed dose for PMMA. The exposure length-integral (expressed as C kg21 cm) was obtained with the ionization chamber dosemeter and converted to the values of DPI for PMMA with the f-factors (0.898 cGy (C kg21)21) [14]. Absorbed doses at the centre and periphery of the phantom were calculated using the effective energy at the isocentre (x50 mm) and the periphery, respectively. All scans for the dose measurement were made in the axial scan (non-helical) mode using the 256-slice CT. Scan conditions were tube voltage of 80 kV, 100 kV, 120 kV, or 135 kV; 100 mAs; and nominal beam width of 8 mm, 32 mm, 64 mm, 96 mm, or 128 mm. The results were averaged for 8 repeated DPI measurements.

Results The effective energies for 80 kV, 100 kV, 120 kV and 135 kV, which were used in converting the exposure to the values of absorbed dose, are summarized in Table 1. The effective energy for the large filter is slightly higher 889

S Mori, K Nishizawa, M Ohno and M Endo Table 1. Effective energies for CT dose index (CTDI) phantoms Phantom

Voltage (kV)

Head 135 (small filter) 120 100 80 Body 135 (large filter) 120 100 80

Effective energy (keV) Centre

Periphery

Average

47.9

54.9

52.5

40.6 36.5 33.3 48.7

49.9 44.0 40.8 61.1

46.8 41.5 38.3 56.9

46.2 43.3 39.6

56.0 50.3 43.8

52.8 48.0 42.4

than that for the small filter due to the different filter shapes. Figure 1 shows DPI300,w and DPI100,w which are normalized against 100 mAs for the body phantom. The relationship between DPI300,w and beam width shows good linearity and the linear correlation coefficient is 0.9996

for the 135 kV tube voltage. Other tube voltages have similar values of linear correlation coefficient. DPI100,w shows reasonably good linearity between the beam widths of 8 mm and 128 mm (e.g. the linear correlation coefficient is 0.9910 for 135 kV). Since the 140 mm long phantom could not cover primary dose and scattered dose in the wider beam width such as 128 mm, second term polynomial fitting was better for DPI100,w (e.g. the fitting correlation is 0.9991 for 135 kV). DPI100,w and DPI300,w were represented as follows: DPI100,w ðvÞ~av2 zbvzc

ð4Þ

DPI300,w ðvÞ~dvze

These coefficients are summarized in Table 2. The DPIws also increase with increasing tube voltage; in particular, DPI300,w for 135 kV for the body phantom is 4.5 times higher than at 80 kV. DPIs for the head phantom (results not shown) are similar to those for the body phantom.

(a)

(b)

Figure 1. The relationships between weighted dose profile integral (DPIw) at the centre of the body phantom and the beam width. (a) DPI300,w. (b) DPI100,w.

Table 2. Coefficients for the fitting equations for DPI100,w, and DPI300,w Phantom

Head

Body

Tube voltage (kV)

Averaged effective energy (keV)

Coefficient a

b

c

d

e

135 120 100 80 135 120 100 80

52.5 46.8 41.5 38.3 56.9 52.8 48.0 42.4

20.0355 20.0263 20.0168 20.0085 20.0086 20.0084 20.0052 20.0026

18.4 13.9 9.1 5.0 8.4 6.5 4.1 2.1

81.2 56.1 32.6 17.3 63.2 32.0 16.2 6.8

25.4 18.9 12.2 7.0 13.8 10.1 6.2 3.1

210.0 141.7 83.7 34.6 120.5 70.4 36.5 16.1

DPI, dose profile integral.

890


Conversion factor using 256-slice CT

(a)

(b)

Figure 2. The relationships between the conversion factor and the beam width for (a) the body and (b) the head phantoms.

Figure 2 shows the relationships between CF value and the nominal beam width for each tube voltage. CF value increases gradually up to the nominal beam width of 128 mm with increasing nominal beam width as well as tube voltage. CF values range from 1.5 to 2.0 for the beam widths and tube voltages. The relationships between CF value and the effective energy are shown in Figure 3. The CF increases with increasing effective energy, and it has the same tendency as in Figure 2.

(a)

Discussion DPIws were measured using the standard CTDI and the 300 mm long phantoms for nominal beam widths of 9–128 mm. We believe that the standard 100 mm ionization chamber and 140 mm long phantom are not reliable, even for the 8 mm beam width [7, 15]. Although the standard phantoms measured DPIws for the nominal beam width less than 100 mm, CF was not 1.0. In particular, CF value ranged from 1.5 to 1.8 for the

(b)

Figure 3. The relationships between the conversion factor and the effective energy for (a) the body and (b) the head phantoms. The British Journal of Radiology, November 2006

891

S Mori, K Nishizawa, M Ohno and M Endo

nominal beam widths of 8 mm and 32 mm, which can be selected by the MSCT. These results can be explained by scattered radiation which depends on the phantom length and which affects DPI strongly [7]. Therefore, CF value is important to both MSCT and 256-slice CT in order to assess doses accurately. CF should be adaptable to MSCT systems because the scatter tails of their dose profiles are thought to be analogous to those for the 256-slice CT, if the effective energies are similar [16].

Conclusion Use of the standard 100 mm ionization chamber and 140 mm phantoms instead of the inconvenient 300 mm long alternatives is made possible by use of the CF values determined in our paper. It was found that the standard system was unreliable even at 8 mm beam width. The relationship demonstrated between the CF and effective energy will allow this method to be adapted to other CT systems.

References 1. Mori S, Endo M, Obata T, Murase K, Fujiwara H, Susumu K, et al. Clinical potentials of the prototype 256-detector row CT-scanner. Acad Radiol 2005;12:148–54. 2. Mori S, Kondo C, Suzuki N, Yamashita H, Hattori A, Kusakabe M, et al. Volumetric cine imaging for cardiovascular circulation using prototype 256–detector row computed tomography scanner (4-dimensional computed tomography) a preliminary study with a porcine model. J Compt Assist Tomogr 2005;29:26–30. 3. Mori S, Obata T, Kishimoto R, Kato H, Murase K, Fujiwara H, et al. Clinical potentials for dynamic contrast-enhanced hepatic volumetric cine imaging with the prototype 256MDCT scanner. AJR Am J Roentgenol 2005;185:253–6. 4. Mori S, Obata T, Nakajima N, Ichihara N, Endo M. Volumetric perfusion CT using prototype 256-detector row computed tomography scanner: preliminary study with healthy porcine model. Am J Neuroradiol 2005;26: 2536–41.

892

5. Kondo C, Mori S, Endo M, Kusakabe K, Suzuki N, Hattori A, et al. Real-time volumetric imaging of human heart without ECG gating by 256-detector row computed tomography: initial experience. J Compt Assist Tomogr 2005;29:694–8. 6. US FDA Code of Federal Regulations. Diagnostic x-ray systems and their major components. Govt. Printing Office, August 1984;21CFR 11020.23. 7. Mori S, Endo M, Nishizawa K, Tsunoo T, Aoyama T, Fujiwara H, et al. Enlarged longitudinal dose profiles in cone-beam CT and the need for modified dosimetry. Med Phys 2005;32:1061–9. 8. Endo M, Mori S, Tsunoo T, Kandatsu S, Tanada S, Aradate H, et al. Development and performance evaluation of the first model of 4D CT-scanner. IEEE Trans Nucl Sci 2003;50:1667–71. 9. Mori S, Endo M, Tsunoo T, Kandatsu S, Tanada S, Aradatte H, et al. Physical performance evaluation of a 256-slice CTscanner for four-dimensional imaging. Med Phys 2004;31:1348–56. 10. Saito Y, Aradate H, Igarashi K, Ide H. Large area 2dimensional detector for real-time 3–dimensional CT (4D CT). Proc SPIE 2001;4320:775–82. 11. Endo M, Mori S, Kandatsu S, Tanada S, Sugihara N, Saito Y, et al. Development of real 4D CT with real-time reconstruction and display. IEEE NSS/MIC Conf Record 2006;M11–252. 12. Edyvean S, Lewis M, Britten A. CT scanner dose survey: measurement protocol Ver. 5. http://www.impactscan. org/1997 [Accessed 5 May 2006]. 13. Hubbell J, Selzer S. Tables of x-ray mass attenuation coefficients and mass energy-absorption coefficients. National Institute of Standard and Technology (NIST) physical reference data 1996. 14. International Commission on Radiation Units and Measurements. Radiation dosimetry: X rays generated at potentials of 5 to 150 kV: ICRU. Publication. 17, Washington D.C.:ICRU Publications, 1970. 15. Nakaonechny K, Fallone B, Rathee S. Novel methods of measuring single scan dose profiles and cumulative dose in CT. Med Phys 2005;32:98–109. 16. Felmlee JP, Gray JE, Leetzow ML, Price JC. Estimated fetal radiation dose from multislice CT studies. AJR Am J Roentgenol 1990;154:185–90.



Equivalent dose to organs and tissues in hysterosalpingography calculated with the FAX (Female Adult voXel) phantom 1

R KRAMER,

PhD,

1

H J KHOURY,

PhD,

1

C LOPES,

PhD

and 2J W VIEIRA,

PhD

1

Departamento de Energia Nuclear, Universidade Federal de Pernambuco, Avenida Prof. Luiz Freire, 100, Cidade Universita´ria, CEP 50740-540, Recife, PE and 2Escola Politećnica, UPE, Recife, PE, Brazil ABSTRACT. Hysterosalpingography (HSG) is a radiological examination indicated for investigating infertility or uterine and tubal pathologies. Women who undergo HSG are relatively young, typically between 20 years and 40 years, and equivalent doses to the ovaries are usually reported to be around 4 mSv per examination. A review of studies on patient dosimetry in HSG revealed that almost all absorbed doses to organs and tissues had been calculated with conversion coefficients (CCs) based on hermaphrodite versions of MIRD5-type phantoms. The CCs applied had been taken from data sets for abdominal or pelvic examinations because CCs for HSG examination were not available. This study uses the FAX (Female Adult voXel) phantom in order to calculate equivalent doses to radiosensitive organs and tissues especially for exposure conditions used in HSG. The calculations were also performed for the MIRD5-type EVA phantom to demonstrate the influence of anatomical differences on organ equivalent dose. The results show organ and tissue equivalent doses as a function of the variations of the exposure conditions. At 4.56 mSv the ovarian equivalent dose calculated for the FAX phantom is about 21% greater than the average ovarian equivalent dose reported in the literature, which reflects the anatomical differences between the FAX and the MIRD5-type phantoms.

Hysterosalpingography (HSG), a radiological examination which delivers relatively high equivalent doses to the ovaries and the uterus, is used to examine the uterine cavity and the patency of the Fallopian tubes. Common indications for HSG are primary and secondary infertility, assessment of tubal patency following reversal of sterilization, of tubal blockage following a difficult sterilization and of the uterine cavity following division of an intrauterine septum. Equivalent doses to the ovaries from HSG examinations are usually around 4 mSv, which triggered a series of investigations on patient dosimetry and on possibilities for the reduction of patient exposure. Table 1 summarizes typical values for the entrance surface air kerma (ESAK), including the contribution from backscattered radiation, and the equivalent dose to the ovaries reported in recent publications [1–5], in which the equivalent dose to the ovaries was calculated by multiplying the ESAK by a conversion coefficient (CC) determined by Monte Carlo methods for hermaphrodite MIRD5-type phantoms, i.e. male bodies with female organs such as ovaries, uterus and breasts [6–8]. However, during the last decade in radiological protection the mathematical MIRD5-type phantoms have increasingly been replaced by tomographic or This work was funded by the Conselho Nacional de Desenvolvimento Cientı´fico e Tecnolo´gico (CNPq) and the Fundaçaõ de Amparo a` Cieˆncia do Estado de Pernambuco (FACEPE).


Received 26 December 2005 Revised 5 April 2006 Accepted 12 April 2006 DOI: 10.1259/bjr/61874578 ’ 2006 The British Institute of Radiology

voxel-based phantoms, which are a true-to-nature representation of the human body. The recently developed FAX (Female Adult voXel) phantom [9], together with the traditional MIRD5-type EVA phantom [10], have been selected to simulate HSG examinations in order to see the influence of the different anatomies on organ and tissue equivalent doses, and also to demonstrate the effect of varying exposure parameters, tube voltage, filtration, field size, field orientation, focus-tofilm distance and projection on these doses.

Materials and methods The phantoms The recently developed FAX (Female Adult voXel) phantom has been segmented from CT images of patients [9]. Organ and tissue masses correspond to the anatomical specifications recommended by the International Commission on Radiological Protection (ICRP) in its Publication 89 for the female reference adult [11], while tissue compositions and densities are based on data published by the International Commission on Radiation Units (ICRU) in its report No. 44 [12]. The EVA phantom has been developed from the first hermaphrodite MIRD5 phantom, which already had ovaries and a uterus, by scaling down the male body to the height of the female reference adult from ICRP Publication 23 and by introducing female breasts [10]. 893

R Kramer, H J Khoury, C Lopes and J W Vieira Table 1. Entrance surface air kerma and ovarian equivalent doses reported in references [1–5] Reference

ESAK (mGy)

Ovarian dose (mSv)

OD/ESAK (Sv /Gy)

Average voltage (kVcp)

Field size (cm6cm)

Fife et al [1] Fernandez et al [2] Gregan et al [3] Calicchia et [4] Khoury et al [5] Average

13.3 23.4 13.1 24.6 19.2 18.7

2.8 4.6 3.1 4.6 2.9 3.6

0.21 0.20 0.24 0.19 0.15 0.20

82 80 78 72 70

24630 24630 24630 24630 30624

Tissue compositions and densities of the EVA phantom have been taken from an early MIRD5 publication [13].

or the filtration reduces the mAs necessary for constant exposure to the detector system.

The EGS4 Monte Carlo code

Conversion coefficients

The EGS4 Monte Carlo code [14] simulates coupled electron-photon transport through arbitrary media. The default version of EGS4 applies an analogous Monte Carlo method, which was used for the calculations of this investigation. Rayleigh scattering has been taken into account, but secondary electrons have not been transported. With respect to the simulation of radiological examinations, a special user code has been developed that outputs absorbed dose to radiosensitive organs and tissues normalized to the ESAK. The X-ray spectra have been taken from the IPEM spectra catalogue [15].

Exposure conditions Based on a review of the exposure conditions reported in references [1–5], the following representative irradiation parameters have been identified for the simulation of the HSG examination: X-ray generator: constant potential Target: tungsten, 17 ˚ Voltage: 70–120 kV at a constant potential (kVcp) Filtration: 2.0–4.0 mm Al Projection: anterior–posterior (AP), posterior–anterior (PA) Field size: 18 cm624 cm, 24 cm630 cm in the detector plane Field position: Centred on uterus Focus to skin distance (FSD): 70 cm, 80 cm, 90 cm Focus to film distance (FFD): 100 cm, 110 cm, 120 cm Figures 1 and 2 show silhouettes of the FAX and the EVA phantoms with X-ray beams, field sizes and FFDs, the uterus in grey and the ovaries in black, respectively.

Results The results will be presented as CCs between organ equivalent doses and the ESAK as a function of the tube voltage and the filtration. ESAK refers to the centre of the radiation field at the surface of the patient’s body. As tube voltage and filtration increase, so do the CCs. However, the absolute value of the equivalent dose to organs will decrease, because increasing the tube voltage 894

Tube voltage Figure 3 shows equivalent dose to the ovaries of the FAX and the EVA phantoms normalized to the ESAK as a function of the tube voltage for field sizes of 18 cm624 cm, and 24 cm630 cm, respectively. For the whole range of tube voltages, the FAX ovarian equivalent dose is ca. 25% greater than the EVA ovarian equivalent dose for the smaller field size, while for the large field this number is about 18%. The reason for the differences is the different depth at which the ovaries are located in the two phantoms – beginning at 6.5 cm depth in the FAX phantom and beginning at 8.3 cm depth in the EVA phantom. The average CC between ovarian equivalent dose and ESAK in Table 1 for references [1–3] is 0.22, and for references [4, 5] it is 0.17. This should be compared with the CCs for the EVA phantom in Figure 3 for a 24 cm630 cm field, for 80 kVcp and 70 kVcp, respectively. For the EVA phantom one finds CCs of 0.24 for 80 kVcp, and of 0.20 for 70 kVcp. These EVA CCs agree reasonably well with the CCs from Table 1, if one takes into account the differences between the real HSG examination and its simulation that is not known here, like the position of the ovaries, patient thickness, X-ray generator, filtration, etc. When the tube voltage increases from 70 kVcp to 120 kVcp, the CCs in Figure 3 increase by ca. 65% over this range of tube voltage. Figure 4 presents CCs for the uterus as function of the tube voltage for both phantoms and the two field sizes already mentioned. The uterus equivalent dose of the EVA phantom is ca. 36% greater than the equivalent dose to the uterus of the FAX phantom over the whole range of photon energies. Again the explanation comes from different depths at which the uteri are located in the two phantoms. When the tube voltage increases from 70 kVcp to 120 kVcp, the CCs in Figure 4 increase by ca. 55% over this range of tube voltage.

Filtration Figure 5 shows CCs for the ovaries for the two phantoms as function of the filtration for the two field The British Journal of Radiology, November 2006

Hysterosalpingography equivalent dose to the FAX phantom

Figure 1. The FAX phantom: hysterosalpingography (HSG) exposure set-up for field524 cm630 cm and FFD 5100 cm.

sizes. When the filtration increases from 2.0 mm Al to 4.0 mm Al, the CCs in Figure 5 increase by ca. 20% over this range of filtration.

Ovaries – For the FAX phantom an increase of the ovarian equivalent dose has been found to be between 5% and 8%, and between 2% and 4% for changing the field orientation from A2 to A1, and from B2 to B1, respectively.

Field orientation Taking the dimensions of the field given as width6 height, then one can see from Table 1 that HSG examinations in Recife/Brazil are usually made with the larger field side representing the width, while the opposite is the case for the other references cited. In order to get an idea of the effect of the field orientation on the organ equivalent doses, the calculations shown so far have been repeated with the larger field side representing the field width. For the radiation fields A1524 cm618 cm, A2518 cm624 cm, B1530 cm624 cm and B2524 cm630 cm, the following results have been found for the range of tube voltages between 70 kVcp and 120 kVcp: Uterus – No change of the uterus equivalent dose has been found when changing the field orientation from A2 to A1, or from B2 to B1. The British Journal of Radiology, November 2006

For the EVA phantom an increase of the ovarian equivalent dose has been found between 13% and 16%, and between 4% and 6% for changing the field orientation from A2 to A1, and from B2 to B1, respectively.

Focus-to-skin distance (FSD) Variations of the FSD between 100 cm and 120 cm did not significantly change the equivalent dose to the ovaries and the uterus.

Equivalent doses per radiograph Although the CCs shown in Figures 3 and 4 increase with the tube voltage, the absolute equivalent dose per 895

R Kramer, H J Khoury, C Lopes and J W Vieira

Figure 2. The EVA phantom: hysterosalpingography (HSG) exposure set-up for field524 cm630 cm and FFD5100 cm.

radiograph usually decreases with increasing tube voltage, because for constant exposure to the detector system the ESAK decreases with increasing tube voltage. Figure 6 shows the ESAK measured with a PMMA phantom as a function of the tube voltage. The reduction

is about 50% when the tube voltage increases from 70 kVcp to 80 kVcp, while for voltages above 80 kVcp the reduction is about 10–15% per 10 kVcp increase. Application of the ESAK data from Figure 6 to the CCs from Figures 3 and 4 provides the organ equivalent

Figure 3. Conversion coefficient between equivalent dose

Figure 4. Conversion coefficient between equivalent dose

to the ovaries and entrance surface air kerma as function of the tube voltage for field sizes of 18 cm624 cm, and 24 cm630 cm.

to the uterus and entrance surface air kerma as function of the tube voltage for field sizes of 18 cm624 cm, and 24 cm630 cm.

896


Hysterosalpingography equivalent dose to the FAX phantom

Figure 5. Conversion coefficient between equivalent dose to the ovaries and entrance surface air kerma as function of the filtration for field sizes of 18 cm624 cm, and 24 cm630 cm.

doses per radiograph shown in Figure 7, which confirm the reduction of exposure to the patient with increasing tube voltage by up to 50%.

The influence of the projection on organ equivalent dose In the HSG studies cited in Table 1, anterior–posterior (AP) was the most frequently used projection. However, in some countries undercouch X-ray tubes are more prevalent than overcouch ones, i.e. posterior–anterior (PA) projections dominate the HSG examinations. Therefore some of the calculations reported above for the FAX phantom have also been done for PA projection. Figure 8 presents ratios between equivalent doses for PA and AP projection for the ovaries and the uterus. The ratios show that using the PA projection can lead to significant equivalent dose reductions compared with the application of the AP projection. The main reason is the presence of the pelvis in the radiation field, which shields internal abdominal organs like the ovaries and the uterus from the incident photon radiation. The position of the ovaries in the frontal part of the FAX abdomen causes an additional reduction of equivalent dose because for PA incidence the ovaries are located at a greater depth. For both field sizes used in this study, the

Figure 6. Entrance surface air kerma per radiograph measured with a homogeneous PMMA phantom 19 cm thick as function of the tube voltage.


Figure 7. Ovarian and uterine equivalent dose per radiograph in hysterosalpingography (HSG) as function of the tube voltage.

effect of equivalent dose reduction is about 60–75% for the ovaries, and about 30–43% for the uterus between 70 kVcp and 120 kVcp tube voltage.

Equivalent dose per HSG examination with AP projection The results of the previous sections have shown that the CCs change with the variations of the exposure parameters. However, a tube voltage of 80 kVcp, a filtration of 3.0 mm Al, a field size of 24 cm630 cm, a FSD of 100 cm and AP projection can be considered typical for a HSG examination. For these settings, one finds the following CCs for the FAX phantom in Figures 3 and 4: Ovaries: 0.244 Sv Gy21 Uterus: 0.205 Sv Gy21 From Table 1, one can find an average ESAK of 18.7 mGy. In absolute terms, this means that for a typical HSG examination of the FAX phantom one gets: Equivalent dose to the ovaries: 4.56 mSv Equivalent dose to the uterus: 3.83 mSv The average ovarian dose for the MIRD5-type phantom from Table 1 is 3.6 mSv, i.e. 21% less than the value for the FAX phantom, which is in agreement with the findings of Figure 3.

Figure 8. Ratios between organ equivalent doses for posteroanterior (PA) and anteroposterior (AP) projection as a function of the tube voltage. 897

R Kramer, H J Khoury, C Lopes and J W Vieira

Conclusion Equivalent dose to the ovaries and to the uterus was calculated with the FAX/EGS4 exposure model for HSG examinations as a function of tube voltage, filtration and FSD. The results have been compared with similar data for the MIRD5-type EVA phantom, which was often used in recent studies. The data have shown that ovarian equivalent doses are 18–25% greater in the FAX phantom compared with the corresponding values for the EVA phantom due to a 2 cm difference between the location of this organ below the surface in the two phantoms. For similar reasons, the uterus equivalent dose of the EVA phantoms was found to be 36% greater than the corresponding value for the FAX phantom. As for the variation with exposure parameters, the calculations revealed an increase of 65% of the ovarian CC, and of 55% of the uterus CC for an increase of tube voltage from 70 kVcp to 120 kVcp. But as demonstrated, the absolute values of ovarian and uterus equivalent dose decrease by up to 50% when the tube voltage increases. Therefore, increasing the tube voltage is usually recommended as step to reduce the exposure to the patient. This applies also to the increase of filtration, although in absolute terms this effect was not shown here, and especially also to the choice of the PA projection. Finally it has to be pointed out that the orientation of the field with regard to width and height can cause differences of between 2% and 15% for the organ equivalent doses discussed in this presentation. The CCs presented can serve as a tool for patient dosimetry. If the ESAK, tube voltage, filtration and field size are known, one can multiply the ESAK by the appropriate CC from Figures 3–5 to obtain an estimate of the equivalent dose to the ovaries or to the uterus.

Acknowledgments The authors would like to thank the Conselho Nacional de Desenvolvimento Cientı´fico e Tecnolo´gico - CNPq and the Fundaçaõ de Amparo a` Cieˆncia do Estado de Pernambuco - FACEPE for the financial support.

References 1. Fife IAJ, Wilson DJ, Lewis CA. Entrance surface and ovarian doses in hysterosalpingography. Br J Radiol 1994;67:860–3.

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2. Fernandez JM, Vanõ´ E, Guibelalde E. Patient doses in hysterosalpingography. Br J Radiol 1996;69:751–4. 3. Gregan ACM, Peach D, McHugo JM. Patient dosimetry in hysterosalpingography: a comparative study. Br J Radiol 1998;71:1058–61. 4. Calicchia A, Chiacchiarelli L, de Felice C, Gigliotti T, Indovina PL, Mazzei F, et al. Evaluation of effective dose in histerosalpingography. Radiat Prot Dosim 1998;80:159–61. 5. Khoury HJ, Maia A, Oliveira M, Drexler G, Kramer R. Patient dosimetry in hysterosalpingography. IAEA International Conference on Radiological Protection of Patients in Diagnostic and Interventional Radiology, Nuclear Medicine and Radiotherapy. Malaga, Spain, 2001. 6. Peterson LE, Rosenstein M. Computer program for tissue doses in diagnostic radiology. Food and Drug Administration, Centre for Devices and Radiological Health, Rockville, Maryland, USA, 1989. 7. Jones DG, Wall BF. Organ doses from medical X-ray examinations calculated using Monte Carlo techniques. NRPB-R186. London: HMSQ, 1985. 8. Hart D, Jones DG, Wall BF. Estimation of effective dose in diagnostic radiology from entrance surface dose and dosearea product measurements. NRPB-R262. London: HNSQ, 1994. 9. Kramer R, Khoury HJ, Vieira JW, Loureiro ECM, Lima VJM, Lima FRA, et al. All about FAX: a female adult voxel phantom for monte carlo calculation in radiation protection dosimetry. Phys Med Biol 2004;49:5203–16. 10. Kramer R, Zankl M, Williams G, Drexler G. ‘‘The calculation of dose from external photon exposures using reference human phantoms and Monte Carlo methods. Part I: The male (Adam) and female (Eva) adult mathematical phantoms’’. GSF-Report S-885. Institut fu¨r Strahlenschutz, GSFForschungszentrum fu¨r Umwelt und Gesundheit, Neuherberg-Mu¨nchen, 1982. 11. ICRP 89. ‘‘Basic Anatomical and Physiological Data for Use in Radiological Protection: Reference Values’’. ICRP Publication 89, International Commission on Radiological Protection. Oxford: Pergamon Press, 2003. 12. ICRU 44. ‘‘Tissue substitutes in radiation dosimetry and measurement’’. ICRU Report 44. International Commission on Radiation Units and Measurements, Bethesda, MD, 1989. 13. Snyder WS, Ford MR, Warner GG, Fisher HL. Estimates of absorbed fractions for monoenergetic photon sources uniformly distributed in various organs of a heterogeneous phantom. Medical Internal Radiation Dose Committee (MIRD) Pamphlet No. 5. J Nucl Med 1969;10(Suppl.):3. 14. Nelson WR, Hirayama H, Rogers DWO. ‘‘The EGS4 Code System’’. SLAC-265 Stanford Linear Accelerator Center, Stanford University, Stanford, California, 1985. 15. Cranley K, Gilmore BJ, Fogarty GWA, Desponds L. Catalogue of diagnostic X-ray spectra and other data. The Institute of Physics and Engineering in Medicine (IPEM), Report No. 78, Electronic version prepared by D Sutton, September 1997.



Comparison of radiation doses to patients undergoing standard radiographic examinations with conventional screen–film radiography, computed radiography and direct digital radiography 1

G COMPAGNONE, PhD, 2M CASADIO BALENI, and 1C BERGAMINI, PhD

MD,

1

L PAGAN,

MSc,

2

F L CALZOLAIO, 2L BAROZZI,

MD

1

Medical Physics Department and 2Accident and Emergency Department, S. Orsola-Malpighi Hospital, Via Massarenti 9, 40138 Bologna, Italy

ABSTRACT. New flat-panel direct digital radiography equipment has recently been installed in our Accident and Emergency Department; its characteristics and versatility are well suited to the work undertaken in this environment. The aim of this study was to compare radiation doses to patients undergoing standard radiographic examinations using conventional screen–film radiography, computed radiography and direct digital radiography; entrance surface dose and effective dose were calculated for six standard examinations (a total of 10 projections) using standard patient exposure parameters for the three imaging modalities. It was found that doses for computed radiography (all examinations) were higher than the doses for the other two modalities; effective doses for direct digital radiography were ,29% and ,43% lower than those for screen–film radiography and computed radiography, respectively. The image quality met the criteria in the European guidelines for all modalities.

In Accident and Emergency (A&E) departments, patients undergo radiological examinations to establish their clinical condition. The importance of having a radiologist to report the images is documented in the literature; discrepancies of up to 58% have been reported between primary care physicians and radiologists [1]. In our hospital, there are eight departments where Xray equipment is used. One of these is the A&E department, comprising three X-ray rooms where approximately 50 000 standard radiographic examinations out of a total of 400 000 are performed annually. In addition to outpatient casualties presenting at A&E, the department also performs radiological examinations for inpatients at night and at the weekend, when the other radiology departments in the hospital are closed. It is for these reasons that new technologies may be installed in A&E departments first. With digital radiography, namely computed radiography (CR) and direct digital radiography (DDR), becoming a viable technology for acquiring X-ray images, departments are looking to replace conventional screen–film radiography (SFR). Potential advantages of digital systems over conventional radiography are well known [2], i.e. that they have a greater dynamic range, wider exposure latitude, postprocessing facilities available, and that there is improved access to images by clinicians and decreased film costs. New flat-panel DDR equipment has recently been installed in the A&E department in our hospital because its characteristics and versatility are well suited to the work undertaken in this environment, e.g. being able to manage taking X-rays of critically ill patients with relative ease. The British Journal of Radiology, November 2006

Received 6 May 2005 Revised 13 March 2006 Accepted 12 April 2006 DOI: 10.1259/bjr/57138583 ’ 2006 The British Institute of Radiology

The aim of this study was to compare radiation doses to patients undergoing standard radiographic examinations using SFR, CR and DDR; entrance skin dose (ESD) and effective dose (E) were calculated for six standard examinations (a total of 10 projections) using standard patient exposure parameters for the three imaging modalities. ESD and E are considered to be efficient and powerful parameters in the protection of patients [3, 4].

Materials and methods In 2003, the SFR system in the A&E department was replaced with a CR system and in May 2004, the X-ray equipment in X-ray Room 2 was replaced with a Siemens Axiom Aristos FX radiography system with a flat panel detector (Siemens, Erlangen, Germany). The Axiom Aristos is a multifunctional system which enables virtually the entire range of radiographic applications to be performed in a single room. A brief description of the X-ray equipment used before and after May 2004 is given in Table 1. The SFR system comprised Kodak TMAT G/RA film (Kodak, Rochester, NY) and Trimax Regular screen (nominal speed class 400). The CR system comprises two Kodak CR readers (CR-850 and CR-900) and three sizes of general purpose image plates, i.e. 18 cm624 cm, 24 cm630 cm and 35 cm643 cm. The image plates contain a europium-activated barium fluoro-halide phosphor with 179262392, 204862500 and 204862500 matrix sizes, and 0.100 mm, 0.120 mm and 0.172 mm pixel pitches, respectively. 899

G Compagnone, M Casadio Baleni, L Pagan et al Table 1. Description of the X-ray equipment used in the Accident and Emergency Department (before and after May 2004) Room

Before May 2004

1

IAE OVERCOUCH TUBE; Tube output @ 80 kV: 77 mGy/mAs @ 75 cm; Focal spot sizes: 0.6 mm, 1.2 mm; Antiscatter grid ratio 12:1, 44 strips/cm COMET OVERCOUCH TUBE; Tube output @ 80 kV: 65 mGy/mAs @ 75 cm; Focal spot sizes: 0.6 mm, 1.2 mm; Antiscatter grid ratio 12:1, 36 strips/cm COMET OVERCOUCH TUBE; Tube output @ 80 kV: 69 mGy/mAs @ 75 cm; Focal spot sizes: 0.6 mm, 1.3 mm; Antiscatter grid ratio 12:1, 36 strips/cm

2

3

After May 2004

PLANMECA ORTHOPANTOMOGRAPH; Tube output @ 76 kV: 167 mGy/mAs @ 35 cm; Focal spot size: 1.2 mm

Unchanged

COMET CEILING-MOUNTED TUBE; Tube output @ 80 kV: 78 mGy/mAs @ 75 cm; Focal spot sizes: 0.6 mm, 1.2 mm; Antiscatter grid ratio 10:1, 36 strips/cm

SIEMENS AXIOM ARISTOS FX; Tube output @ 81 kV: 118 mGy/mAs @ 75 cm; Focal spot sizes: 0.6 mm, 1.0 mm; Antiscatter grids ratio 12:1, 40 strips/cm, fgd5115 cm and 180 cm Unchanged

VARIAN CEILING-MOUNTED TUBE; Tube output @ 80 kV: 55 mGy/mAs @ 75 cm; Focal spot sizes: 0.6 mm, 1.2 mm; Antiscatter grid ratio 12:1, 36 strips/cm

The DDR system comprises a 0.5 mm thick thallium doped caesium iodide phosphor (CsI:Tl). The readout array, to which the phosphor is coupled, consists of a 43 cm643 cm amorphous silicon (a-Si) photodiode and a thin film transistor (TFT) array (300063000 pixels and 0.143 mm pixel pitch). The X-rays interact with the phosphor and release light photons. The light from the phosphor promotes charge in the photodiode array with the amount of charge being proportional to the intensity of the incoming X-ray photons. Each TFT element is then sequentially addressed so that the charge in the photodiodes is read out and digitized. The detector resolution is limited by pixel pitch; this may be reduced by the effect of light spread in the phosphor layer. The images were finally reported by a consultant radiologist on a Kodak Autorad workstation. ESD were calculated for six standard radiographic examinations (a total of 10 projections): anteroposterior (AP) Abdomen, posteroanterior (PA) Chest, lateral (LAT) Chest, AP Lumbar Spine, LAT Lumbar Spine, LAT Lumbo-Sacral Joint, AP Pelvis, AP Skull, LAT Skull, AP Urinary Tract. To calculate E, first we measured the X-ray tube output using a technique previously described [5]. For each X-ray tube, 30 air kerma measurements were made with an ionization chamber (model 90X6-6, connected to a Radiation Monitor Controller model 9010; Radcal Corporation, Monrovia, USA) held in a scatter-free support on the central axis of the X-ray beam (Table 1). Instruments are calibrated annually, with the calibration traceable to an SIT (National Calibration Service in Italy) centre. The Harpen mathematical model [6] was then used to obtain the X-ray tube output for every tube voltage (kVp), milliampere-seconds product (mAs) and focus-toskin distance (FSD) used in each of the clinical protocols (provided by the radiologist). A backscatter factor of 1.35 was used to calculate the ESD from air kerma, as suggested in the European Guidelines [7]. Finally, E has been calculated from ESD using the NRPB conversion coefficients [8]. In order to correctly apply these factors, it is necessary to know the tube potential used in clinical protocols and 900

the total X-ray tube filtration. The total tube filtration is measured annually as part of a quality assurance programme; acceptance, status and constancy tests are performed by the Medical Physics Department (the department is certified to UNI EN ISO 9001-2000). Assessment of image quality was undertaken by Consultant Radiologists within A&E; images from the three modalities, i.e. SFR, CR and DDR, were qualitatively evaluated to ensure that the quality criteria for diagnostic radiographic images of the European Guidelines [7] were met.

Results The parameters used for standard radiographic examinations on adult patients are reported in Table 2. It should be noted that for some of the examinations performed with the Siemens Axiom Aristos equipment, the FSDs are different to those used with SFR and CR as the system is self-positioning and self-centring; this is a feature of the equipment. Manufacturers of digital systems often recommend the use of low tube potentials, but this will lead to less penetrating beams and hence possibly higher doses. It is important, therefore, to ensure that systems are optimized [9]. For Kodak CR systems, the exposure index (EI) is the average raw data pixel value within the anatomical area exposed [10]: an EI ,2000 is indicative of optimum exposure level. EI is directly related to patient dose by the equation EI510006log10(exposure in mR)+2000 [10]. Although the use of low tube potentials is often recommended, clinically diagnostic images can be achieved using high tube potentials. An example of this is the PA chest radiograph obtained at 125 kVp (Figure 1); an EI very close to the value recommended by the manufacturer was obtained, i.e. EI51970, and the quality criteria of the European Guidelines [7] were satisfied. With the digital systems we generally used a higher tube potential and lower mAs to reduce the patient dose (Table 2). Careful adjustment of these parameters will allow patient doses to be optimized without sacrificing image quality by taking full advantage of the The British Journal of Radiology, November 2006

Comparison of SFR, CR and DDR radiation doses in standard examinations Table 2. Parameters used for standard radiographic examinations using conventional screen–film radiography (SFR), computed radiography (CR) and direct digital radiography (DDR) Examination and projection

AP Abdomen PA Chest LAT Chest AP Lumbar Spine LAT Lumbar Spine LAT Lumbo-Sacral Joint AP Pelvis AP Skull LAT Skull AP Urinary tract

kVp

FSD (cm)

Total filtration (mm Al)

SFR

CR

DDR

SFR

CR

DDR

SFR

CR

DDR

85 120 130 75 85 85 75 75 70 80

86 120 135 80 90 90 85 72 72 86

81 125 125 81 90 90 81 73 73 N/A

92 151 140 91 86 86 90 92 98 90

92 151 140 91 86 86 90 92 98 90

79 155 144 78 73 73 77 79 85 N/A

4.0 4.3 4.3 4.3 4.3 4.3 4.3 4.3 4.3 4.0

4.0 4.4 4.4 4.4 4.4 4.4 4.4 4.4 4.4 4.0

3.3 3.3 3.3 3.3 3.3 3.3 3.3 3.3 3.3 N/A

FSD, focus-to-skin distance; AP, anteroposterior; PA, posteroanterior; LAT, lateral.

post-processing tools, because higher tube potential can reduce contrast [11]. In this study, doses to patients were assessed by calculating ESD for six standard radiographic examinations (a total of 10 projections), and E for complete examinations (Tables 3 and 4). 5 months after the installation of the DDR equipment, a survey was undertaken on 175 A&E patients to establish the preferred imaging modality. Approximately 30% of examinations were necessarily performed using the CR imaging modality, e.g. patients being X-rayed using mobile radiographic equipment. Of the remaining patients, approximately 86% were X-rayed using the DDR equipment.

Discussion Dose comparisons between CR, SFR and DDR have been reported for chest radiography [12, 13]; the results of this study confirm the findings of the two reported studies, i.e. that CR generally results in higher ESDs than those in SFR and DDR. Comparison between DDR and SFR doses shows that, in general, DDR results in lower ESDs than those in SFR; the AP Skull projection is the only one where the ESD for DDR is higher than that for SFR. A possible reason for this is that AP Skull is not performed by high tube potential technique with DDR. Table 2 shows that AP Skull and Abdomen are the only examinations where both tube potential and FSD for DDR are lower than tube potential and FSD for SFR, respectively: the effect of these two factors is an overall decrease in the mAs value for DDR equipment compared with SFR system. This reduction in mAs is greater for Abdomen than for AP Skull: as a consequence, the AP Skull is the only projection where ESD for DDR is higher than ESD for SFR. However, E for complete Skull examination for DDR is still lower than that for SFR. Table 4 shows that for complete examinations, the values of E for DDR are approximately 29% and 43% lower than those for SFR and CR, respectively. E has been calculated from ESD using the NRPB conversion coefficients [8]; the coefficients increase with applied tube potential and to a lesser extent with tube filtration. The magnitude of any dose saving that can be achieved with different imaging modalities is controversial [13]: some authors found a dose reduction with CR The British Journal of Radiology, November 2006

systems of approximately 50% compared with a conventional 200-speed SFR system [14], whilst others have reported the same values of ESD and E for the LAT Lumbar Spine radiograph when a CR system replaced a 300-speed SFR system [15]. Other authors have reported a dose increase in CR systems of 33–58% compared with a 400-speed SFR system [16]. The findings of this study are therefore neither completely unexpected nor in contradiction with those of other trials. It would seem reasonable to state that doses with the use of CR systems are approximately the same as those for a 300-speed SFR system, but it has been also reported that the speed class quoted by manufacturers does not predict the actual relative speeds of the SFR systems [17] and there are wide differences in image quality for similar speed systems [18]: therefore the importance of optimization, whichever system is being used, must be stressed. The CR system used in this study has been optimized together with the manufacturer engineers: nevertheless, a full optimization may take a long time because it is a dynamic process where radiologists, physicists,

Figure 1. Posteroanterior (PA) chest radiograph obtained with computed radiography: male, 45 years old, 125 kVp, focus to skin distance (FSD) 5151 cm, Exposure Index51970 (the Exposure Index recommended by the manufacturer is 2000). The image satisfies the quality criteria for PA chest radiographs given in the European Guidelines [7]. 901

G Compagnone, M Casadio Baleni, L Pagan et al Table 3. Entrance surface dose (mGy) for standard radiographic examinations using conventional screen–film radiography (SFR), computed radiography (CR) and direct digital radiography (DDR), and percentage differences between the systems Examination and projection

SFR

CR

DDR

CR vs SFR

DDR vs SFR

DDR vs CR

AP Abdomen PA Chest LAT Chest AP Lumbar Spine LAT Lumbar Spine LAT Lumbo-Sacral Joint AP Pelvis AP Skull LAT Skull AP Urinary tract

1.88 0.07 0.15 1.77 4.27 5.12 1.81 1.44 1.10 2.18

2.40 0.11 0.20 2.54 5.39 5.39 1.83 1.61 1.11 2.51

1.64 0.06 0.13 1.16 1.72 1.72 1.02 1.58 0.89 N/A

+28% +57% +33% +44% +26% +5% +1% +12% +1% +15%

213% 214% 213% 234% 260% 266% 244% +10% 219% N/A

232% 245% 235% 254% 268% 268% 244% 22% 220% N/A


Table 4. Effective dose (mSv) for standard radiographic examinations using conventional screen–film radiography (SFR), computed radiography (CR) and direct digital radiography (DDR), and percentage differences between the systems Examination

SFR

CR

DDR

CR vs SFR

DDR vs SFR

DDR vs CR

Abdomen (AP) Chest (PA & LAT) Lumbar Spine (AP & LAT) Pelvis (AP) Skull (AP & LAT) Urinary tract (AP)

0.280 0.029 0.309 0.295 0.027 0.168

0.358 0.041 0.476 0.326 0.029 0.193

0.223 0.023 0.179 0.168 0.022 N/A

+28% +41% +54% +11% +7% +15%

220% 221% 242% 243% 219% N/A

238% 244% 262% 248% 224% N/A


(a)

(b)

Figure 2. Posteroanterior (PA) chest radiographs obtained with computed radiography and direct digital radiography. (a) Direct digital radiography: female, 47 years old, 125 kVp, focus to skin distance (FSD) 5155 cm. (b) Computed radiography: male, 28 years old, 125 kVp, FSD5151 cm. Both images satisfy the quality criteria for PA chest radiographs given in the European Guidelines [7]. 902


Comparison of SFR, CR and DDR radiation doses in standard examinations

radiographers and manufacturers are involved day after day in the achievement of continuous improvement. On the whole, CR increases doses for a single exposure, but the larger dynamic range reduces the number of examinations that have to be repeated [19, 20], e.g. for stretcher patients or for patients in intensive care where the FSD often cannot be standardized. However, the increase in CR patient doses found in this study should be put into perspective, e.g. chest examinations. In this example, the additional E per complete examination (1 PA projection + 1 Lateral projection) using CR was approximately 0.01 mSv and 0.02 mSv compared with SFR and DDR, respectively. The risk of developing a fatal or other cancer, or other serious defect (including hereditary effects), during the course of life for the population undergoing chest examinations in the A&E Department (about 20 000 examinations/year) is the same as for the whole population, i.e. 7.3% per Sievert [3]. Thus, nearly 1 million patients would each have to undergo complete chest examinations with the use of a CR system in the A&E Department to produce 1 and 2 additional health defects in this population compared with SFR and DDR systems, respectively. Moreover, the probability of a fatal cancer being induced in an individual patient is dependent on the age of the patient: e.g. for patients exposed after the age of 70 years, the risk can be assumed to be reduced to one-third of the average risk as the patient is likely to die before the cancer develops [21]. Similar calculations could be made for all other examinations. Potential advantages of digital systems over conventional radiography are well known [2], i.e. that they have a greater dynamic range, wider exposure latitude, postprocessing facilities available, and that there is improved access to images by clinicians and decreased film costs. The post-processing tool is a very important and useful characteristic provided that the individual performing this task is adequately trained [22]. The DDR equipment has been shown to have the following advantages: (a) the radiographer can stay close to the patient throughout the examination; (b) the equipment is self-positioning and self-centring around the patient, and this characteristic is very useful for stretcher patients; (c) the image acquisition times are much shorter than in SFR or CR; (d) an increased patient throughput (a further survey showed that out of a total of 175 patients, 60% of examinations are now performed using DDR); (e) the rapid availability of images for timely clinical decision-making, which may be vitally important for severely injured patients. When the workload is heavy, the patient throughput can be maximized by using the DDR equipment for chest, abdomen, pelvis, spine and skull examinations, and by using CR for extremities. The image quality characteristics of flat-panel technology are theoretically better than those of SFR and CR because at a comparable resolution the DDR has a higher detective quantum efficiency (DQE) [23]; this advantage can be utilized in reducing the patient dose whilst maintaining image quality and without changing the The British Journal of Radiology, November 2006

signal-to-noise ratio. The appropriateness of this approach would, however, require verification. All images met the criteria in the European Guidelines for all modalities and were used for reporting by the Consultant Radiologists. Figure 2 shows chest images obtained using CR and DDR. It shows that both systems satisfied the quality criteria for diagnostic radiographic images [7]: full inspiration; symmetrical reproduction of the thorax; scapulae outside the lung field; reproduction of the whole rib cage above the diaphragm; visualization of the retrocardiac lung and the mediastinum; sharp reproduction of the trachea and proximal bronchi, the borders of the heart and aorta; visualization of the spine through the heart shadow. Finally, the radiologists prefer the appearance of the DDR images.

Conclusions This study compared radiation doses to patients undergoing standard radiographic examinations using SFR, CR and DDR; ESD and E were calculated using standard exposure parameters. Effective doses for DDR were found to be ,29% and ,43% lower than those for SFR and CR, respectively. The image quality met the criteria in the European Guidelines for all modalities. Even though in the A&E department most examinations are now performed with the DDR system, we are aware that the optimization effort has to also continue for the CR equipment, because digital systems need to be optimized very carefully by a dynamic process that may take a long time. Conventional technologies are soon to be replaced by digital ones and although DDR equipment has been shown to have many advantages, further work is required to specifically look at image quality and dose for real patients.

Acknowledgments We express our thanks to the Reviewers for their valuable advice regarding this study.

References 1. Gatt ME, Spectre G, Paltiel O, Hiller N, Stalnikowicz R. Chest radiographs in the emergency department: is the radiologist really necessary? Postgrad Med J 2003;79:214–7. 2. Volk M, Strotzer M, Gmeinwieser J, Alexander J, Frund R, Seitz J, et al. Flat-panel X-ray detector using amorphous silicon technology. Invest Radiol 1997;32:373–7. 3. International Commission on Radiological Protection. 1990 Recommendations of the ICRP. Annals of the ICRP Vol. 21 Publication 60. Oxford, UK: Pergamon Press, 1991. 4. McCollough CH, Schueler BA. Calculation of effective dose. Med Phys 2000;27:828–37. 5. Compagnone G, Pagan L, Bergamini C. Local diagnostic reference levels in standard X-ray examinations. Radiat Prot Dosim 2005;113:54–63. 6. Harpen MD. A mathematical spread sheet application for production of entrance skin exposure nomograms. Med Phys 1996;23:241–2. 7. European Commission. European guidelines on quality criteria for diagnostic radiographic images. EUR 16260EN. Office for Official Publ. of the Europ Commun. 1996.

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G Compagnone, M Casadio Baleni, L Pagan et al 8. Hart D, Jones DG, Wall BF. Estimation of effective dose in diagnostic radiology from entrance surface dose and dosearea product measurements. NRPB-R262. Chilton, Didcot: National Radiological Protection Board, 1994. 9. Peters SE, Brennan PC. Digital radiography: are the manufacturers’ settings too high? Optimisation of the Kodak digital radiography system with aid of the computed radiography dose index. Eur Radiol 2002;12:2381–7. 10. Bogucki TM, Trauernicht DP, Kocher TE. Characteristics of a storage phosphor system for medical imaging. Technical and Scientific Monograph 6. Rochester, USA: Kodak Health Sciences, 1995. 11. Livingstone RS, Koshy CG, Ray DV. Evaluation of work practices and radiation dose during adult micturating cystourethrography examinations performed using a digital imaging system. Br J Radiol 2004;77:927–30. 12. Weatherburn GC, Bryan S, Davies JG. Comparison of doses for bedside examinations of the chest with conventional screen-film and computed radiography: results of a randomized controlled trial. Radiology 2000;217:707–12. 13. Herrmann KA, Bone´l H, Stabler A, Kulinna C, Glaser C, Holzknecht N, et al. Chest imaging with flat-panel detector at low and standard doses: comparison with storage phosphor technology in normal patients. Eur Radiol 2002;2:385–90. 14. Sandmayr H, Wallentin D. Computer integrated radiology system: analogue goes digital. Eur Radiol 1997;7(Suppl. 3):S90–3. 15. Weatherburn GC, Bryan S. The effect of a picture archiving and communication systems (PACS) on patient radiation doses for examination of the lateral lumbar spine. Br J Radiol 1999;72:534–45.

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16. Bragg DG, Murray KA, Tripp D. Experiences with computed radiography: can we afford the cost? AJR Am J Roentgenol 1997;169:935–41. 17. Brennan PC, Devereux SA. An assessment of the usefulness of screen-film speed classifications. Eur Radiol 2002;12: 1577–83. 18. Guibelalde E, Fernandez JM, Vano E, Llorca A, Ruiz MJ. Image quality and patient dose for different screen-film combinations. Br J Radiol 1994;67:166–73. 19. Strotzer M, Gmeinwieser J, Spahn M, Volk M, Frund R, Seitz J, et al. Amorphous silicon, flat-panel, X-ray detector versus screen-film radiography. Invest Radiol 1998;33: 33–8. 20. Weatherburn GC, Bryan S, West M. A comparison of image reject rates when using film, hard copy computed radiography and soft copy images on picture archiving and communication systems (PACS) workstations. Br J Radiol 1999;72:653–60. 21. Cox R, McGibbon BH. Diagnostic medical exposures: advice on exposure to ionising radiation during pregnancy. Estimates of late radiation risks to the UK population. Documents of the National Radiological Protection Board, Vol.4,No.4, 1993. 22. Tsapaki V, Kottou S, Kollaros N, Dafnomili P, Koutelou M, Vano E, et al. Comparison of a conventional and a flat-panel digital system in interventional cardiology procedures. Br J Radiol 2004;77:562–7. 23. Samei E, Hill JG, Frey GD, Southgate WM, Mah E, Delong D. Evaluation of a flat panel digital radiographic system for low-dose portable imaging of neonates. Med Phys 2003;30:601–7.



Dose escalation to combat hypoxia in prostate cancer: a radiobiological study on clinical data 1

J Z WANG,

PhD,

2

X A LI,

PhD

and 1N A MAYR,

MD

1

Department of Radiation Medicine, The Ohio State University, Columbus, OH 43210 and Department of Radiation Oncology, Medical College of Wisconsin, Milwaukee, WI 53226, USA

2

ABSTRACT. Earlier studies have demonstrated that hypoxic regions exist in human prostate cancer and the degree of hypoxia correlates with the treatment outcome of radiotherapy. Using the concept of the clinical oxygen enhancement ratio (COER), the linear-quadratic (LQ) model was extended to account for the effect of tumour hypoxia. The clinical data collected at the Fox Chase Cancer Center for prostate cancer were analysed based on the LQ model as well as the tumour control probability (TCP) model. The LQ and TCP parameters (a50.15 Gy 21, a/b53.1 Gy and the number of clonogens K5106,107 cells) determined in earlier studies were used to derive the COER for prostate cancer: COER51.4 with a standard confidence interval (CI) of (1.2, 1.8). The result is consistent with the in vitro OER measurements of human tumour cell lines under chronic hypoxia conditions. This implies that a higher dose is needed to overcome tumour hypoxia. For prostate tumours, the prescription dose required to overcome tumour hypoxia is 165 Gy (CI: 153,186 Gy) for permanent 125I implants and 88 Gy (CI: 74,118 Gy) in 2 Gy fractions for external-beam radiotherapy. The impact of LQ parameters on the calculations of COER and dose escalation was discussed. This study provides a preliminary estimate of the dose escalation needed to overcome tumour hypoxia based on clinical data. More clinical data with better statistics and longer follow-up time are required to further tune the radiobiological modelling of hypoxia for prostate cancer.

Hypoxia has been identified in many human cancers, including head and neck, lung, breast, cervix, bladder and prostate cancer, and the adverse impact of hypoxia on clinical tumour control has been well recognized [1–10]. Recently, Movsas et al [1–3] and Parker et al [4] have demonstrated that hypoxic regions do exist in human prostate carcinoma and the treatment outcome of radiation therapy correlates with the degree of hypoxia. Oxygenation status (oxygen pressure, PO2) may thus be another underlying biological parameter beyond the classic prognostic factors (clinical stage, Gleason score and prostate specific antigen (PSA)) predicting a critical component of treatment failure in prostate cancer. Movsas et al [1] reported a significant difference in the biochemical control rate at 2 years (31% versus 92%) between two patient groups stratified by the prostate/ muscle (P/M) PO2 ratio (,0.05 versus >0.05). Many other investigators have also found similar correlations between hypoxia and tumour progression [5–10]. It is generally believed that high-risk tumours tend to exhibit more severe hypoxia [3, 10]. Although it is well established that hypoxic cells/regions are less sensitive to radiation and that tumour hypoxia can result in clinical radiation therapy failure, it remains unclear how Address correspondence to: Jian Z Wang, Department of Radiation Medicine, James Cancer Hospital and Solove Research Institute, The Ohio State University, 300 W 10th Ave, Rm 094, Columbus, OH 43210, USA. E-mail: [email protected].


Received 4 January 2006 Revised 2 June 2006 Accepted 5 June 2006 DOI: 10.1259/bjr/18700614 ’ 2006 The British Institute of Radiology

to overcome hypoxia in the clinical setting. Specific quantitative dose guidelines are needed that can be applied to the clinical setting, particularly to dose escalation trials aimed to combat hypoxia in prostate cancer. The linear-quadratic (LQ) and tumour control probability (TCP) models extended to address tumour hypoxia were used to analyse the clinical data. Recently, the a/b ratio of the LQ model for prostate cancer has become a highly debated topic in the radiation therapy community [11–22]. By taking into account the effect of tumour repopulation, Wang et al [14, 15] and Kal and Van Gellekom [19] took two different approaches, and obtained similar results of a/b ratio (around 3.1 Gy). Based on the clinical data of external-beam radiotherapy (EBRT) and low- and highdose rate (LDR/HDR) brachytherapy [12, 13, 23], Wang et al [14, 15] have derived a new set of LQ parameters (a50.15 Gy21, a/b53.1 Gy and repair half-time Tr516 min) with estimated numbers of clonogens around 106,107, depending on the patient risk level. The new results provide reasonable estimates of radiosensitivity and the number of clonogens for human prostate tumours [14, 15, 24–26]. The radiobiological modelling with these parameters provides a consistent interpretation for most currently available clinical data, including EBRT, permanent brachytherapy, HDR brachytherapy and their combinations [14, 15, 26]. Fowler et al [16, 21] argued that for slowly growing tumours (e.g. prostate cancer), the onset of tumour 905

J Z Wang, X A Li and N A Mayr

repopulation may be delayed significantly to over 200 days. Wang et al [17, 22] pointed out that there is no strong biological evidence for such arguments. The clinical results reported so far have been controversial. In a recently reported study, Perez et al [27] did observe an impact of elapsed treatment time on outcome of EBRT for T2 prostate tumours, implying the onset time of repopulation should be relatively short. Therefore in our study, the original assumption of the onset time of tumour repopulation will be followed for various sets of LQ parameters. Recently, the effect of hypoxia has been taken into account by Nahum et al [18] in their modelling study. However, their model parameters (a, b and oxygen enhancement ratio (OER)) were completely based on in vitro measurements, which may not be fully applicable to the clinical micro-environment of prostate cancer in vivo [28]. Tumour heterogeneity of radiosensitivity has been studied [20] and due to the large uncertainty existing in the current clinical data and the large number of parameters used, it is not possible to give a reasonable estimate of a/b ratio for any statistical significance using the heterogeneous model. For this reason, we chose the homogeneous model in this study. The purpose of our study is to quantify the dose escalation to address tumour hypoxia in prostate cancer. The LQ and TCP models are used to interpret the clinical data recently reported from Fox Chase [1] based on the concept of OER. The specific aim of this analysis is to derive the clinical OER (COER) and to determine the dose escalation necessary to offset the hypoxia effect using either permanent brachytherapy or EBRT.

Materials and methods LQ and TCP models for prostate cancer The general LQ model [29] was used in this study. In this model, the surviving fraction S of cells irradiated to a total dose D within an overall treatment time T is given by: S~ exp ({aD{bGD2 zcT)

ð1Þ

where a and b characterize intrinsic radiosensitivity, G is the dose-rate factor accounting for sub-lethal damage repair, c is the effective tumour repopulation rate [c5ln(2)/Td, and Td is the effective clonogen doubling time]. A median potential tumour doubling time of 42 days measured in vivo from 7 prostate cancer patients was used [30]. This repopulation rate corresponds to a biological effective dose (BED) of 0.11 Gy per day (for a50.15 Gy 21), compatible with the results of slowgrowing tumours presented in a repopulation-doseequivalent study by Jones et al [31]. The repopulation time T can be replaced by (T–Tk), where Tk is the onset time of tumour repopulation. Because the actual value of this parameter is under debate [16, 17, 21, 22, 27] and no clinical data are available, we followed the original assumptions used by different investigators when they derived the LQ parameters from clinical data, i.e. Tk50 for the LQ parameters by Wang et al [14, 15] and by Kal and Van Gellekom [19], and Tk.200 days (i.e. ignoring 906

the repopulation effect) for the LQ parameters by Brenner et al [11, 12] and by Fowler et al [13]. For EBRT, the total dose D equals nd and the dose-rate factor G equals 1/n, where n is the number of dose fractions and d is the dose per fraction. The treatment time T of EBRT can be approximated as the number of treatment fractions multiplied by 1.4 (5 fractions per week). For permanent brachytherapy, D, G and T have been derived in many papers (e.g. [14, 29]) as follows: R0 1{e{lT , l 2R2 1 1 1{e{2lT { 1{e{(mzl)T ; ð2Þ G~ 2 0 D (m{l) 2l mzl 1 c T&{ ln l aR0 D~

where m is the repair rate of tumour cells [m5ln(2)/Tr], R0 is the initial dose rate (R05D0?l, D0 is the prescribed dose) and l is the decay constant for the implanted isotopes [l5ln(2)/Ts, Ts is the half-life time of the radioisotope]. Ts559.4 days for 125I implants. The Poisson TCP model is often used to link cell killing of a radiotherapy scheme to treatment outcome [32], TCP~ exp ({KS)

ð3Þ

where K is the initial number of tumour clonogens. We consider it as an effective number of clonogens associated with the patient risk level. Two representative sets of LQ parameters derived from clinical data [11–15, 19] and one set of generic LQ parameters for tumour [10] were used in this study (Table 1). The numbers shown in the parentheses of Table 1 indicate the standard confidence intervals (CI) [14] and they were used to estimate the uncertainties of the COER and dose escalations. The clonogen number K of 1.66106 (CI: 5.66104,8.86107) derived from clinical data [14] was used in the calculations with the first parameter set (see the following subsection ‘‘Fox Chase clinical data’’). Due to limited data for the other two parameter sets, the clonogen numbers were arbitrarily set to match the clinical data and the CIs were not considered. Our calculation was mainly based on the first parameter set. The later two parameter sets were used to test the model dependency of the final results presented in this study.

LQ model extended for tumour hypoxia OER is a dose-modifying parameter, which is defined as the ratio of doses needed to produce the same biological effect for cells without and with oxygen. Traditionally, OER was commonly used to describe the cell survival curves obtained in in vitro measurements. The OER of X-ray irradiation for most mammalian cells in vitro is about 3 at high doses and is possibly lower (about 2) at doses below approximately 2 Gy [10]. In this work, we apply the OER concept to clinical data (i.e. in vivo) to characterize tumours with or without hypoxia The British Journal of Radiology, November 2006

Hypoxia in prostate cancer Table 1. Radiobiological parameters used in this study LQ/TCP parameters a,b

Wang et al Brenner et alb Generic parametersb

a (Gy

21

)

0.15 (0.11–0.19) 0.04 0.3

a/b (Gy)

Tr (min)

Td (days)

3.1 (2.6–3.6) 1.5 10

16 (1–90) 114 60

42 (22–62) ‘c 42

a

The numbers in the parentheses show the standard confidence intervals (CIs). See text for detailed references. c The repopulation effect is ignored. Td5‘ is equivalent to Tk.200 days. LQ, linear-quadratic; TCP, tumour control probability. b

[5, 33–35]. Compared with in vitro experiments, the hypoxia in in vivo situations may be quite different. The population for tumours in vivo always reflects a spectrum of hypoxia. Even for well-identified hypoxic tumours, the hypoxic cells may exist only in subvolumes of tumours. Therefore, the OER obtained retrospectively from clinical data represents only a statistical average and may be different from the corresponding in vitro values for the same type of tumours. Zolzer and Streffer [36] argued that tumours in vivo should be dominated by chronic hypoxia and the radiosensitivity for cells under chronic hypoxia conditions would be higher than those under acute hypoxia conditions. They reported that the OER measurements of two human tumour cell lines under continued/chronic hypoxia conditions (1.3 and 1.5 for MeWo and squamous carcinoma 4451, respectively) were quite different from the OER values measured under acute hypoxia conditions. To distinguish this feature, we term herewith the OER obtained from clinical data as clinical OER (COER). For simplicity, the COER is represented by H throughout this paper. Given a certain amount of clonogens, to achieve the same TCP for two patient groups, the cell killing efficiency should be the same. That is: S~ exp ({ah Dh {bh GD2h zcT) ~ exp ({ao Do {bo GD2o zcT)

ð4Þ

where subscripts h and o label the parameters for tumours with and without hypoxia, respectively. Assuming Dh and Do represent the doses to achieve the same TCP for the two patient groups, respectively, we have Dh5H?Do. To be consistent with Equation (4), we assume that ah5ao/H, bh5bo/H2, and they are valid to apply to the LDR 125I brachytherapy and the EBRT with low dose fractions (d52 Gy). It appears that the hypoxia impacts more significantly on the quadratic term than on the linear term. Such relationships of the LQ parameters between aerated and hypoxic cells have been observed in in vitro experiments of low dose irradiation (0.5–3 Gy) for Chinese hamster V79-171 cells [37, 38]. Therefore, the LQ model extended to the hypoxia effect can be expressed as (Sh for hypoxic cells): Sh ~ exp ({aD=H{bGD2 H 2 zcT)

ð5Þ

Similar formulae to describe the hypoxia effect can be found in the literature [5, 33–35]. Because of the limited clinical data, we will focus on the effect of hypoxia on radiosensitivity only, and neglect The British Journal of Radiology, November 2006

the influence of hypoxia on repair time or repopulation rate. Hill et al presented data of animal and human carcinoma of the cervix showing no evidence for a difference in repopulation kinetics between aerated or hypoxic tumours [39, 40]. It is generally assumed that the transient hypoxia can be overcome by tumour reoxygenation by protracting radiotherapy (i.e. fractionated EBRT or LDR brachytherapy). The reoxygenation effect has been modelled in previous studies (e.g. [41]). Therefore, it is believed that the oxygen effect on the clinical outcome observed in the 125I data (see next subsection) should be mainly due to the chronic hypoxia [36]. In this work, we use the above formulae to study the effect of chronic hypoxia. For patients with tumours that are only partially hypoxic, i.e. containing oxygenated and hypoxic subvolumes, the two-compartment model was used to calculate the overall surviving fraction [18, 33, 35], S~(1{fh )So zfh Sh

ð6Þ

where fh is the hypoxic fraction of cells in the tumour and So is for aerated cells. Animal tumour data and human tumour data indicate that the hypoxic fraction ranges from 0 to 50%, with an average value of 15% [5, 10, 42]. This average value along with the range from 5% to 50% was used in this study. The uncertainty in estimating the COER originates from the uncertainties of the LQ/TCP parameters (as shown in Table 1), the hypoxia fraction and the clinical data. The LQ/TCP parameters are strongly correlated to each other and constrained by clinical data; therefore, their uncertainties are not independent from each other [14] and have been considered in this study. The code autoEUD described in Wang and Li [26] was used to propagate the different uncertainties to determine the CIs of COER and dose escalation. For independent parameters, their contributions to the final CIs were calculated based on the root of the quadratic-summation. For a given patient group of prostate cancer, based on Equations (1) to (6), we could calculate the TCP for a given radiation treatment or determine the required radiation dose D to achieve a given TCP. The software AutoEUD and the typical dose–volume histogram (DVH) of EBRT and brachytherapy described in a previous paper [26] were used to automate these calculations.

Fox Chase clinical data Since 1999, Movsas et al have published a series of papers to address the hypoxia of prostate cancer [1–3]. 907

J Z Wang, X A Li and N A Mayr

For the first time, they demonstrated the existence of hypoxia in human prostate cancer based on the in vivo electrode measurements of oxygen levels, and that the degree of hypoxia in prostate cancer correlated to the treatment outcome of radiotherapy. 57 patients with localized prostate cancer were included in their study [1]. Before radiotherapy treatments, custom-made Eppendorf microelectrodes were used to measure the PO2 in both pathologically involved regions of the prostate and normal muscle (as an internal control). Real-time ultrasound imaging was used to guide the microelectrode measurement. For each patient, approximately 100 PO2 readings were obtained along 3 to 5 tracks in both prostate and muscle [3]. Following the Eppendorf PO2 measurement, 48 patients received permanent 125I brachytherapy. The dose prescribed to the prostate was 145 Gy. Nine patients were treated on a dose escalation protocol involving a 46 Gy (in 23 daily fractions) of EBRT to the pelvic region plus two HDR boost implants (8.75 Gy or 9.75 Gy per fraction). Hormonal therapy prior to radiation therapy was given to nine patients. In their study, the biological failure was defined as two consecutive rises in PSA without a return to baseline. With a median follow-up time of 19 months (ranging from 4 months to 31 months), the overall 2 year bNED (biologically no evidence of disease) was estimated as 81% for the entire patient group (as shown in Figure 1 of this paper and Figure 2 of [1]). In order to study the impact of hypoxia, the P/M ratio of PO2, which eliminates the potential interpatient and technical variations, was used to analyse the treatment outcome. The Fox Chase experience demonstrated that the P/M ratio strongly correlated to the bNED outcome. A threshold analysis showed that the 2 year bNED differed significantly at a P/M ratio of ,0.05 (hypoxic group) versus >0.05 (aerated group): 31% versus 92% (p,0.0001) (see Figure 1). There were 12 and 45 patients in the two groups, respectively. The classic prognosis factors, including clinical stage, Gleason score, pretreatment PSA and age at treatment were similar in the two patient groups. Based on the numbers of all patients and the patients at risk for each group, using the binominal distribution, we estimated the standard CI of the 2 year bNED for the hypoxic tumour group to be (15%, 47%). Because of the large numbers of patients in the aerated tumour group, the uncertainty of their 2 year bNED is negligible. According to the patient characteristics of the three prognosis factors, patients of the Fox Chase data were stratified into the low- or intermediate-risk groups. Because most patients presented low-risk prognosis factors, the clonogen number (1.66106) of low-risk patients was used for this study [14, 15]. Two radiotherapy modalities (125I brachytherapy vs combined EBRT and HDR) were used in the Fox Chase Study. Since the patient number of the combined modality (EBRT + HDR) is relatively small (9 patients) and the treatment outcome for low-risk patients is close for the two regimens, we ignored the difference of the two modalities and combined the data in the following study. For similar reasons, we also ignored the influence of hormone therapy applied to nine patients out of the total 57 patients. A representative DVH obtained from permanent 125I implant (Figure 1 of Ref. [26]) was used 908

Figure 1. 2 year biologically no evidence of disease (bNED) versus prescription dose of 125I brachytherapy. The symbols represent the Fox Chase clinical data [1] and the curves represent the model calculations. The three points (circle, square and triangle), as well as the three curves (solid, dashed and dash-dotted), are for patient groups with hypoxic, mixed and aerated prostate tumours, respectively. The prescription dose of the clinical data was 145 Gy of 125I implant. The vertical dotted line indicates the prescription dose for the hypoxic group to achieve a bNED of 81%. The vertical error bar shows the standard confidence interval (CI) of the clinical data for the hypoxic group (the CIs for the other two groups are not shown) and the horizontal error bar shows the CI of dose escalation for the hypoxic group to achieve a target bNED of 81%.

to account for the dose inhomogeneity of the Fox Chase data.

Results Clinical OER for hypoxic prostate cancer The clinical data and the model calculation are summarized in Figure 1. It has been reported that, based on LQ and TCP models, the 145 Gy 125I implant has a TCP of 82% for low-risk prostate patients [14, 26]. This calculated TCP value is consistent with the clinically observed 2 year bNED (81%) for the mixed patient group in the Fox Chase study (see ‘‘Fox Chase clinical data’’ in the Materials and Methods section). Using the bNED data for the aerated and hypoxic groups, we derived the COER and the CI for the hypoxic tumours: H51.4 (CI: 1.2,1.8) (Table 2). This result is very consistent with the in vitro OER measurements of two human tumour cell lines (1.3 and 1.5 for MeWo and squamous carcinoma 4451, respectively) under chronic/continued hypoxia conditions [36]. The large uncertainty is mainly related to the limited information for the hypoxic fraction in the hypoxic tumours. If the uncertainty of hypoxic fraction is not considered, the uncertainty of COER is greatly reduced to¡0.13, i.e., in the range of (1.27, 1.53). Similar calculations, except for the uncertainty estimation, were performed for other sets of LQ parameters. In the COER derivation with the parameters of Brenner et al (i.e. a50.04 Gy 21, a/b51.5 Gy, Tr51.9 h with repopulation ignored), only the hypoxic clonogens were considered [43]. The COER obtained with this parameter set is The British Journal of Radiology, November 2006

Hypoxia in prostate cancer

Figure 2. (a) Tumour control probability (TCP) as a function of the fraction of tumour hypoxia (fh) for individual patients of prostate cancer. The prescription dose is 145 Gy for 125I brachytherapy. (b) Dose escalation of 125I brachytherapy required overcoming tumour hypoxia (to achieve a TCP of 81%) as a function of hypoxia fraction fh.

H51.5. Using the LQ parameters that are considered generic for tumours (i.e. a50.3 Gy 21, a/b510 Gy, Tr51 h and Td542 days), we obtained the COER: H51.2. Interestingly, the result derived with the parameters of Wang et al lies between the values obtained with these two sets of LQ parameters, and the differences among the COER values obtained with the three parameter sets are smaller than the CIs determined by various uncertainties. More clinical data with better statistics and better knowledge of the hypoxic fraction are required to further narrow down the CIs. To summarize the results, the COERs, as well as the required dose escalations (see the following subsection), derived with different LQ/ TCP parameters are listed in Table 2.

Dose escalation for hypoxic tumours Based on the derived COER, the dose required to compensate for tumour hypoxia was calculated using the code autoEUD [26]. To bring the TCP for hypoxic tumours to the same level as that for the general population of prostate patients (TCP581%), the prescription dose required for permanent 125I implant should be 169 Gy. This dose may be used to target the hypoxia region if it can be identified. Otherwise, a reduced dose of 165 Gy (CI: 153,186 Gy) should be prescribed to the entire prostate, including the nonhypoxic regions, to achieve the same treatment outcome (a TCP of 81%) as shown in Figure 1. This dose escalation is practically feasible, as demonstrated by a recent study [44]. The calculations performed so far provide a guide of dose escalation for a cohort of patients with hypoxic prostate tumours. If the hypoxic subvolume of individual tumours can be detected, for example by biological/ functional imaging, the data presented may also be used to target these subvolumes and design the individualized radiation plan. Figure 2 shows the TCP and dose prescription as a function of the hypoxic fraction fh for individual prostate tumours. The TCP presented in Figure 2a is for the dose prescription (145 Gy of 125I implant) used in the current clinical practice. Figure 2b shows the dose escalation required to overcome the effect of hypoxia and to bring the TCP to 81% as that of the general population of prostate cancer patients. For individual patients with hypoxic fraction changing from 5% to 50%, the prescription dose is increased from 157 Gy to 181 Gy. The British Journal of Radiology, November 2006

The above results obtained from the Fox Chase data (brachytherapy) data set may be used to derive the dose escalation of EBRT. Various clinical studies have shown that a biological equivalence exists between EBRT and permanent brachytherapy for prostate cancer [13, 26]. For patients with similar risk-levels, the 145 Gy brachytherapy yielded a similar treatment outcome as 71 Gy EBRT in 2 Gy fractions [13]. Based on the LQ and TCP models and the COER derived in this study, we calculated the EBRT dose required to compensate for the effect of hypoxia. This EBRT dose is found to be 88 Gy (CI: 74,118 Gy) in 2 Gy daily fractions. Similarly, based on the COERs derived with the parameters of Brenner et al or with the generic LQ parameters for tumours (see Table 1), we obtained the dose escalation necessary to compensate for prostate hypoxia. With Brenner et al’s parameters, the prescription dose would be 193 Gy for 125I brachytherapy and 109 Gy for EBRT. With the generic tumour parameters, these doses would be 155 Gy for the 125I brachytherapy and 102 Gy for EBRT. Because the clinical data used to derive the COERs are from 125I brachytherapy, the dose escalations calculated for the 125I brachytherapy have smaller CIs and they match the sequence of the a/b ratio in the three parameter-sets. However, the extrapolated EBRT doses show large uncertainty and the results based on the parameters of Brenner et al and the generic tumour parameters are unusually high; therefore, the prediction of EBRT dose escalation has limited clinical value.

Conclusion and discussion In this study, we have analysed the reported clinical data for hypoxic prostate cancer. Based on the LQ and TCP models extended to account for tumour hypoxia, the clinical OER for prostate tumour was obtained: COER51.4 (CI: 1.2,1.8). The result agrees with the in vitro OER measurements under chronic/continued hypoxia conditions [36]. To overcome tumour hypoxia, a dose escalation to 165 Gy (CI: 153,186 Gy) (instead of 145 Gy) is required for 125I permanent implants. A similar calculation was performed for EBRT. The EBRT dose needs to be escalated from 71 Gy to 88 Gy in 2 Gy fractions in order to achieve a TCP of 81% for the hypoxic tumour group. There is in vivo evidence showing that increasing levels of hypoxia in prostate cancer correlate significantly with 909

J Z Wang, X A Li and N A Mayr Table 2. Impact of radiobiological parameters on the calculation of COERs and dose escalations LQ/TCP parameters a,b

Wang et al Brenner et alb Generic parametersb

COER

125

I dose (Gy)

EBRT dose (Gy)

1.4 (1.2–1.8) 1.5 1.2

165 (153–186) 193 155

88 (74–118) 109 102

a

The numbers in the parentheses show the standard confidence intervals (CIs). The corresponding LQ/TCP parameters used in this study are shown in Table 1. LQ, linear-quadratic; TCP, tumour control probability; COER, clinical oxygen enhancement ratio; EBRT, external-beam radiotherapy. b

the increasing clinical stage and patient age [2]. The P/M ratio of PO2 may be used as a prognostic factor for highrisk patients. Dose escalation in both EBRT and brachytherapy has been proven effective to treat highrisk patients of prostate cancer [12, 23, 26, 44, 45]. This study used similar approaches to address the unsatisfactory treatment outcome of hypoxic tumours. The required dose escalations (165 Gy for 125I implant and 88 Gy for EBRT) are practically deliverable [23, 44], and clinical trials with similar dose levels or even higher are being conducted in several institutions. The results obtained in this work are based on the Fox Chase clinical data [1] that, for the first time, shows the existence of hypoxia in human prostate tumours and the correlation of the degree of hypoxia with treatment outcome. There are several limitations in the Fox Chase study that may compromise the accuracy in the current estimation of the COER and the corresponding target doses. First, the PO2 measurement with Eppendorf microeletrodes is a sampling method which measures only a sampled portion of the prostate. It may not reflect the overall PO2 distribution in the entire tumour. The method also fails to discriminate cell type and viability, resulting in PO2 readings from less significant tissues. Second, this method cannot distinguish the hypoxia types: chronic and transient. The transient hypoxia may be reduced or overcome by providing enough time for reoxygenation during the treatment course. Third, the study contained only a small number of patient samples with limited follow-up time, and furthermore mixed treatment modalities (two radiation methods and hormone therapy) were involved in the small data sample. Because of the limited statistical power, only two patient groups could be identified with an artificially determined threshold of the P/M ratio. More clinical data with larger numbers of patients, longer follow-up time and consistent treatment are certainly desired to further refine our conclusions. Because of the limited clinical data, it was not possible to derive independent relationships between a, b and COER in this study. Instead, the relationships, ah5ao/H, bh5bo/H2, from the literature [5, 33–35] were used. Although similar relationships have been observed in in vitro measurements [37, 38], the results obtained in this analysis would be affected if such relationships were not applicable to prostate cancer. The COER calculation was based on an average hypoxic fraction of 15% [5, 10, 42]. A change of this value would change the results of the COER as well as the dose escalation. There are also concerns about the OER changing with the irradiation dose fraction. In this study, we assume that the COER keeps constant for the LDR brachytherapy and the 910

low-dose-fraction (2 Gy) EBRT. The use of alternative models, which include temporal variations in the hypoxic fraction, might also lead to quite different conclusions. Therefore, the results obtained in this study may be limited by the approximations used in the models and the uncertainties shown in the clinical data. Caution needs to be exercised in using the presented data for clinical decision-making purposes.

References 1. Movsas B, Chapman JD, Hanlon AL, et al. Hypoxic prostate/muscle Po2 ratio predicts for biochemical failure in patients with prostate cancer: preliminary findings. Urology 2002;60:634–9. 2. Movsas B, Chapman JD, Greenberg RE, et al. Increasing levels of hypoxia in prostate carcinoma correlate significantly with increasing clinical stage and patient age. Cancer 2000;89:2018–24. 3. Movsas B, Chapman JD, Horwitz EM, et al. Hypoxic regions exist in human prostate carcinoma. Urology 1999;53:11–8. 4. Parker CP, Milosevic M, Toi A, et al. Polarographic electrode study of tumor oxygenation in clinically localized prostate cancer. Int J Radiat Oncol Biol Phys 2004;58:750–7. 5. Okunieff P, Hoeckel M, Dunphy EP, et al. Oxygen tension distributions are sufficient to explain the local response of human breast tumors treated with radiation alone. Int J Radiat Oncol Biol Phys 1993;26:631–6. 6. Ho¨ckel M, Knoop C, Schlenger K, et al. Intratumoral pO2 predicts survival in advanced cancer of the uterine cervix. Radiother Oncol 1993;26:45–50. 7. Ho¨ckel M, Schlenger K, Mitze M, et al. Hypoxia and radiation response in human tumors. Semin Radiat Oncol 1996;6:3–9. 8. Nordsmark M, Overgaard M, Overgaard J. Oxygenation status predicts radiation response in advanced squamous cell carcinoma of head and neck. Radiother Oncol 1996;41:31–9. 9. Brizel DM, Sibley GS, Prosnitz LR, et al. Tumor hypoxia adversely affects the prognosis of carcinoma of the head and neck. Int J Radiat Oncol Biol Phys 1997;38:285–9. 10. Hall EJ. Radiobiology for the radiologist. 5th edn. Philadelphia, PA: Lippincott Williams and Wilkins, 2000:91–111. 11. Brenner DJ, Hall EJ. Fractionation and protraction for radiotherapy of prostate carcinoma. Int J Radiat Oncol Biol Phys 1999;43:1095–101. 12. Brenner DJ, Martinez AA, Edmundson GK, et al. Direct evidence that prostate tumors show high sensitivity to fractionation (low a/b ratio), similar to late-responding normal tissue. Int J Radiat Oncol Biol Phys 2002;52:6–13. 13. Fowler J, Chappell R, Ritter M. Is a/b for prostate cancer really low? Int J Radiat Oncol Biol Phys 2001;50:1021–31.


Hypoxia in prostate cancer 14. Wang JZ, Guerrero M, Li XA. How low is the a/b ratio for prostate cancer? Int J Radiat Oncol Biol Phys 2003;55:194–203. 15. Wang JZ, Li XA, Yu CX, et al. The low a/b ratio for prostate cancer: what does the clinical outcome of HDR brachytherapy tell us? Int J Radiat Oncol Biol Phys 2003;57:1101–8. 16. Fowler JF, Ritter MA, Fenwick JD, et al. How low is the a/b ratio for prostate cancer? In regard to Wang et al. [letter]. Int J Radiat Oncol Biol Phys 2003;57:593–5. 17. Wang JZ, Guerrero M, Li XA. Low a/b ratio for prostate cancer: in response to Dr. Fowler et al. [letter]. Int J Radiat Oncol Biol Phys 2003;57:595–6. 18. Nahum AE, Movsas B, Horwitz EM, et al. Incorporating clinical measurements of hypoxia into tumor local control modeling of prostate cancer: Implications for the a/b ratio. Int J Radiat Oncol Biol Phys 2003;57:391–401. 19. Kal HB, Van Gellekom MPR. How low is the a/b ratio for prostate cancer? Int J Radiat Oncol Biol Phys 2003;57:1116–21. 20. Carlone M, Wilkins D, Nyiri B, et al. Comparison of a/b estimates from homogeneous (individual) and heterogeneous (population) tumor control models for early stage prostate cancer. Med Phys 2003;30:2832–48. 21. Dasu A, Fowler JF. Comments on ’Comparison of in vitro and in vivo a/b ratios for prostate cancer’. Phys Med Biol 2005;50:L1–4. 22. Wang JZ, Stewart RD, Carlson DJ, et al. Reply to ‘‘comments on ‘comparison of in vitro and in vivo a/b ratios for prostate cancer’’’. Phys Med Biol 2005;50: L5–8. 23. Levegru¨n S, Jackson A, Zelefsky MJ, et al. Fitting tumor control probability models to biopsy outcome after threedimensional conformal radiation therapy of prostate cancer: pitfalls in deducing radiobiologic parameters for tumors from clinical data. Int J Radiat Oncol Biol Phys 2001;51:1064–80. 24. King CR, Mayo CS. Is the prostate a/b ratio of 1.5 from Brenner and Hall a modeling artifact? [letter]. Int J Radiat Oncol Biol Phys 2000;47:536–8. 25. Wyatt RM, Beddoe AH, Dale RG. The effects of delays in radiotherapy treatment on tumor control. Phys Med Biol 2003;48:139–55. 26. Wang JZ, Li XA. Evaluation of external beam radiotherapy and brachytherapy for localized prostate cancer using equivalent uniform dose. Med Phys 2003;30:34–40. 27. Perez CA, Michalski J, Mansur D, et al. Impact of elapsed treatment time on outcome of external-beam radiation therapy for localized carcinoma of the prostate. Cancer, J 2004;10:349–56. 28. Wang JZ, Mayr NA, Li XA, et al. Modeling prostate cancer: in regard to Nahum et al. [letter]. Int J Radiat Oncol Biol Phys 2005;61:309–10. 29. Dale RG. The application of the linear-quadratic dose-effect equation to fractionated and protracted radiotherapy. Br J Radiol 1985;58:515–28.


30. Haustermans K, Fowler JF. A comment on proliferation rates in human prostate cancer [letter]. Int J Radiat Oncol Biol Phys 2000;48:303. 31. Jones B, Cominos M, Dale RG. Application of biological effective dose (BED) to estimate the duration of symptomatic relief and repopulation dose equivalent in palliative radiotherapy and chemotherapy. Int J Radiat Oncol Biol Phys 2003;55:736–42. 32. Porter EH. The statistics of dose/cure relationships for irradiated tumors. Br J Radiol 1980;53:210–27. 33. Buffa FM, West C, Byrne K, et al. Radiation response and cure rate of human colon adenocarcinoma spheroids of different size: the significance of hypoxia on tumor control modeling. Int J Radiat Oncol Biol Phys 2001;49: 1109–18. 34. Hendry JH, Thames HD. Fractionation sensitivity and the oxygen effect. Br J Radiol 1990;63:79–80. 35. Popple RA, Ove R, Shen S. Tumor control probability for selective boosting of hypoxic subvolumes, including the effect of reoxygenation. Int J Radiat Oncol Biol Phys 2002;54:921–7. 36. Zolzer F, Streffer C. Increased radiosensitivity with chronic hypoxia in four human tumor cell lines. Int J Radiat Oncol Biol Phys 2002;54:910–20. 37. Watts ME, Hodgkiss RJ, Jones NR, et al. Radiosensitization of Chinese hamster cells by oxygen and misonidazole at low X-ray doses. Int J Radiat Biol 1986;50:1009–21. 38. Skarsgard LD, Harrison I, Durand RE. The radiation response of asynchronous cells at low dose: evidence of substructure. Radiat Res 1991;127:248–56. 39. Speke AK, Hill RP. The effects of claiming and reoxygenation on repopulation during fractionated irradiation. Int J Radiat Oncol Biol Phys 1995;31:857–63. 40. Hill RP, Fyles W, Milosevic M, et al. Is there a relationship between repopulation and hypoxia/reoxygenation? Results from human carcinoma of the cervix. Int J Radiat Biol 2003;79:487–94. 41. Brenner DJ, Hlatky LR, Hahnfeldt PJ, et al. A convenient extension of the linear-quadratic model to include redistribution and reoxygenation. Int J Radiat Oncol Biol Phys 1995;32:379–90. 42. Jones B, Dale G. Estimation of tumor hypoxia fraction from clinical data sets compatible with accelerated repopulation. Acta Oncologica 1998;37:263–8. 43. Brenner DJ, Hall EJ. In response to Drs. King and Mayo: Low a/b values for prostate appear to be independent of modeling details [letter]. Int J Radiat Oncol Biol Phys 2000;47:2753–65. 44. Li XA, Wang JZ, Stewart RD, et al. Dose escalation in permanent brachytherapy for prostate cancer: dosimetric and biological considerations. Phys Med Biol 2003;48:2753–65. 45. Stock RG, Stone NN, Tabert A, et al. A dose-response study for I-125 prostate implants. Int J Radiat Oncol Biol Phys 1999;41:101–8.

911


SHORT COMMUNICATION

Scatter from radiotherapy beams emerging from primary barriers: an aid to bunker design 1

D C BENITO,

MSc

and 2A L MCKENZIE,

DSc

1

H H Wills Physics Laboratory, University of Bristol, Tyndall Avenue, Bristol BS8 1TL and Department of Medical Physics and Bioengineering, Bristol Haematology and Oncology Centre, Horfield Road, Bristol BS2 8ED, UK 2

ABSTRACT. Charts to assist in linear accelerator bunker design have been produced using the Monte Carlo framework, Geant4. These charts assess the amount of forward scatter produced at different angles to the beam axis by concrete and steel barriers irradiated by 6 MV and 10 MV photon beams at normal incidence. These new charts complement existing charts of broad-beam transmission through walls. This is because the existing charts give no indication of the amount of scatter emerging at large angles from the beam axis, for example, towards the maze entrance.

Designers of linear accelerator bunkers commonly refer to charts such as those given in references [1, 2]. These plot primary beam attenuation against wall thickness at given X-ray energies. Experience shows that such charts are sufficiently accurate for design purposes. The relative intensity of back-scatter from the bunker walls may also be determined from charts in the above publications to a level of accuracy that is sufficient for radiation protection calculations. Such back-scatter may be used to estimate dose rates at maze entrances, provided that the linear accelerator is not orientated in the room in such a way that the primary beam can be aimed at the inner maze wall. In such cases, if the inner maze wall has been designed insufficiently thick, the dose rate from wide-angle scatter from the X-ray beam emerging from the wall can be more than an order of magnitude greater than the dose rates, typically a few mGy h21, arising from the multiple-order scatter that eventually finds its way down the maze (Figure 1). An example is discussed later. Surprisingly, perhaps, no charts have been published to cater for this situation. While primary-beam charts can be used to calculate the intensity of broad-beam transmission through an inner maze wall, no charts exist to show the level of scatter that may emerge at discrete angles and contribute to the dose at the maze entrance (Figure 1). Designers who have access to software packages based on Monte Carlo programs may use these to find the scatter from such beam configurations. However, such packages are not widely available, and require a period of training and familiarization. For those who do not design bunkers regularly, these inconveniences militate against their use. In attempting to provide a simple, practical solution to what is a relatively common problem we have calculated a series of plots, illustrated in Figures 2–5. These show 912

Received 10 January 2006 Revised 24 May 2006 Accepted 30 May 2006 DOI: 10.1259/bjr/94993761 ’ 2006 The British Institute of Radiology

the intensity of radiation scattered at different angles from a primary beam as it emerges from a wall made of concrete or steel. The data are plotted as ratios of scattered dose rate at 1 m from the exit point relative to the dose rate of the primary beam as it enters the wall, where it is assumed to have an area of 100 cm2. In practical use, the bunker designer will need to know the scatter not at 1 m, but at several metres from the centre of the area of the beam as it emerges from the wall. This distance would typically be the distance to the maze entrance. The dose rate there will be found by applying the inverse-square law. The inaccuracies of using a distance measured from a single point in the centre of a finite beam area are minimized because the distance to the maze entrance is generally significantly greater than the size of the exit beam.

Method In order to calculate the scatter charts, we used the Geant4 Monte Carlo framework code [3] to simulate scattering within the wall. We chose 6 MV and 10 MV spectra [4] from Elekta (Elekta; Elekta Ltd, Crawley, West Sussex, UK) and Varian (Varian; Varian Medical Systems UK Ltd, Crawley, West Sussex, UK) linear accelerators to create beams incident on concrete of density 2.35 tonnes m23 and steel of density 7.85 tonnes m23. These materials are commonly used in bunker walls. Where practicable, we checked that our use of the code generated results that were consistent with our experiments or with published charts. Comparison with existing charts was only indirect, since no charts of the kind we have produced were previously available. Experimental verification was similarly not without difficulty, because dose rates were generally low and the level of background leakage and scatter tended to be The British Journal of Radiology, November 2006

Short communication: Aiding bunker design

Figure 1. A simplified plan view of a radiotherapy treatment room.

of the same order as that of scatter from the maze wall. This is discussed in more detail below.

Plots The calculated plots are shown in Figures 2–5. The statistics were such that the plots were smooth, apart

from data points at the largest barrier thicknesses at the greatest scatter angles where random deviations from smoothed data were in the order of 30%. Because of this, some plots have been replaced with the best fitting (exponential) dotted line. The position of the calculated data points indicates the closeness of fit. All four charts show the expected decrease in scatter dose rate with increasing wall thickness at given angles

Figure 2. Dose ratio D1/Do, as illustrated in Figure 1 (concrete, 6 MV).


913

D C Benito and A L McKenzie

Figure 3. Dose ratio D1/Do, as illustrated in Figure 1 (steel, 6 MV).

of scatter. The agreement between the gradients of these charts and those of primary attenuation is striking. For instance, referring to Figure 2, for 6 MV radiation through concrete, the gradient of the intensity of scatter exiting at 60 ˚ is three decades per metre of concrete. This is the same as the gradient found in published charts [1, 2] of the attenuation of primary 6 MV radiation through concrete. Similarly, for 6 MV radiation through steel (Figure 3), the gradient of the intensity of scatter exiting at 60 ˚ is three decades for approximately 300 mm of steel. After correcting for the relative densities of steel and concrete, this is essentially the same as the gradient in Figure 2 for 6 MV radiation through concrete. Consideration of the physics shows that the similarity of the gradients of scatter and primary plots is to be expected. Suppose that the intensity of 60 ˚ scatter exiting a concrete wall of thickness 0.2 m is measured. Suppose that an extra metre of concrete is now added ‘‘upstream’’, making a total thickness of 1.2 m. Any scatter from this extra metre will be so attenuated by the final 0.2 m when it exits the wall that it will contribute a negligible amount to the measured scatter, relative to that arising from the 0.2 m nearest the detector. However, after the primary beam has travelled the added metre and enters the final 0.2 m, it is diminished in intensity by three decades. Therefore, the scatter arising within the final 0.2 m will also have diminished by three decades. This, in broad terms, explains why plots of both the exit scatter and primary attenuation have the same gradient. An example will serve to illustrate the use of the charts. Suppose that the inner maze wall in Figure 1 is made of concrete of density 2.35 tonnes m23 and is 0.6 m thick. This was a common thickness for secondary barriers in bunkers 914

designed for cobalt treatment. Suppose that a 6 MV linear accelerator is to be installed in the bunker. Assuming an instantaneous dose rate of 4 Gy min21 at the isocentre, the dose rate at the entrance to the maze wall, 4 m from the source, is 15 Gy h21. For a collimator setting of 10 cm610 cm, the area at the wall is 1600 cm2. Using Figure 2, the dose rate scattered at 60 ˚ to 1 m from the centre of the exit beam is 561026615 Gy h21 61651200 mGy h21. The factor 16 arises because the charts are presented for a beam area of 100 cm2 on entry to the wall. Using the inverse-square law to find the dose at the maze entrance, which also happens to be 4 m from the centre of the exit beam, and using a factor of unity to convert whole-body dose to effective dose, the effective dose rate at the maze entrance is then 75 mSv h21. Such a calculation shows that additional protection might be needed for the inner maze wall in this case, possibly provided by steel. From the discussion above, the same Figure 2 could be used to determine the scatter exiting from a wall of concrete and steel by finding the equivalent thickness of an all-concrete wall.

Comparison with measurement Before these plots can be used with confidence to design bunkers, the level of scattered dose that they predict must be compared with measurement. The reason why direct comparison with experiment is difficult is explained below. We were able to check that elements of our implementation of the Geant4 code were in accordance with existing charts by comparing our predictions of backscatter with those of the experimental data summarized in Figure 19 of Reference [2]. The middle column of Table 1 shows our Geant4 estimates of the intensity of 6 MV radiation back-scattered to 1 m The British Journal of Radiology, November 2006


Figure 4. Dose ratio D1/Do, as illustrated in Figure 1 (concrete, 10 MV).

Figure 5. Dose ratio D1/Do, as illustrated in Figure 1 (steel, 10 MV).


915

D C Benito and A L McKenzie Table 1. Comparison between the Geant 4 simulation and data from reference [2] for back-scattered doses for three different scattering angles Scattering angle (degrees)

Relative back-scattered dose estimated from Geant4

Data from Reference [2]

180 150 120

6.6261025 ¡ 0.861025 6.3761025 ¡ 0.761025 4.5261025 ¡ 0.461025

7.4061025 6.8061025 5.1061025

from a concrete barrier irradiated by a 100 cm2 beam. The figures are shown relative to the intensity of the incident beam. The statistical uncertainties in the Monte-Carlo results are indicated by two standard deviations. The estimates agree with published data to within about 10%, which suggests that the code was functioning as intended, as far as could be tested using results for back-scatter. The difficulty in verifying the plots by direct experiment is that (1) any barrier from which oblique scatter is measured will also transmit head leakage from the linear accelerator into the detector and (2) back-scatter from the barrier tends to be multipally scattered around the barrier, and again adds to the signal due to the oblique scatter. The intensity of both the leakage and the multiple scatter are generally of the same order as, or greater than, that of the oblique scatter, and it is difficult to identify the individual components in the detected signal. This neutralizes the effectiveness of building temporary barriers within a linear accelerator room in order to test the oblique-scatter plots. Experiments based on bunkers in clinical use run into difficulty because dose rates from scatter from inner maze walls in existing installations tend to be very low. Inner maze walls may well be over-protected: maze walls that were inadvertently designed too thin would have had extra protection added at the first opportunity. These low dose rates are not only of the same order as the background from leakage and multiple scatter, but they are also at the detection limit of most readily available ionization chambers. It might be perceived that the presence of head leakage and multiple scatter down a maze would detract from the usefulness of oblique-scatter plots. However, in practice, it is the dose beyond the maze entrance that is of interest, where the head leakage is insignificant because of attenuation by the outside secondary wall, and where the multiple scattered radiation has diminished with distance down the maze and can, in any case, be estimated from existing charts. It is here, beyond the maze entrance, that the oblique-scatter plots should prove useful in ensuring that scatter from the inner maze wall to the outside is kept to acceptably low levels. Instead of building barriers inside the linear accelerator bunker, another way to test the plots is to measure the dose rate beyond the maze entrance where the detector is protected from head leakage by the outside secondary wall. The detector should be positioned slightly to one side of the maze barrier, so that it can ‘‘see’’ the area where the primary beam exits the inner maze wall, but not the far end of the maze, which is a source of multiple scatter. There is only one linear accelerator bunker in our radiotherapy centre where the primary beam can be 916

directed at the inner maze wall. This wall consists of a 0.6 m concrete barrier to which has been added 0.25 m of steel on the inside of the room. A Southern Scientific RO10 survey meter using a 400 cm3 ionization chamber was placed outside the end of the maze, and slightly to one side of it, in order to minimize the level of multiple scatter entering the detector. The instrument had been calibrated less than 2 months previously in accordance with the laboratory accreditation requirements of the UK Accreditation Service. This detector was 4.43 m from the centre of the linear accelerator beam at the point where it exited the inner maze wall. The scatter angle from the centre of the beam to the detector was 62.5 ˚. The primary beam of the linear accelerator was aimed at normal incidence onto the inner maze wall, and set to deliver a 40 cm620 cm beam as defined at a target distance of 1 m. The beam was narrower in the horizontal direction so that the oblique scatter entering the detector would be confined within a smaller range of scatter angles. The measured dose rate was 0.7 mGy h21. When the detector was moved out of the line of sight of the exit beam, the measured dose rate fell to between 0.1 mGy h21 and 0.2 mGy h21. This is representative of the dose rate from multiple scatter reaching the detector even although it was within the ‘‘shadow’’ of the main component of multiple scatter from the far end of the maze. Hence the dose rate attributable to oblique scatter from the inner maze wall was in the order of 0.5 mGy h21. This dose rate has to be compared with that predicted by the plots. We accounted for the two-component nature of our maze wall by multiplying by the ratio of the densities of steel and concrete to estimate the thickness of an equivalent wall built either entirely of concrete or entirely of steel. The calculated thicknesses are 1.435 m and 0.43 m, respectively. For an oblique angle of 62.5 ˚, the charts in Figures 2 and 3 predict a relative scattered dose rate of 2.161028 and 1.261028, respectively. Using the linear accelerator output of 240 Gy h21 at 1 m from the source, and taking care to estimate the incident beam size at the same distance from the source as that at which the output is calculated, the predicted dose rates at the detector are 1.1 mGy h21 or 1.9 mGy h21 depending upon whether an all-concrete or an all-steel wall is assumed. Hence, in comparison with the measured rate of 0.5 mGy h21, the plots predict between two and four times the actual dose rate. Several uncertainties attend this conclusion, however. The measured dose rates were at the very limits of detection of the instrument. In addition to the concrete and steel composition of the wall, there is a layer of plaster of indeterminate thickness, but which will reduce the measured dose rate. The predicted dose rates had to be made using a calculated effective thickness, which turned out to be so large that the plots even had to be extrapolated slightly to determine the relative scattered dose.

Conclusion The charts presented here are for calculating the intensity of radiation emerging from a wall and scattered in different directions, and are particularly helpful in The British Journal of Radiology, November 2006


designing bunkers where a primary beam is directed at a maze wall. A ratio of between two and four between the predicted and the measured dose rate is equivalent to an uncertainty of 100–200 mm concrete or 30–60 mm steel. The lower values in these ranges may be acceptable, particularly since the plots appear to be conservative, that is, a design based on the plots might tend to overestimate the protection required. The discrepancy may arise because of the very low dose rates that are available and because of the uncertainty of the effect of the additional plaster layer to the wall which would reduce the amount of the discrepancy. In the light of the remaining uncertainties, we should welcome experimental verification by other centres who may happen to have thinner maze walls, ideally of uniform composition. Such independent corroboration would lend weight to these plots as a primary tool in room design.


Acknowledgments We are grateful for the assistance of Paul Stevens and Henry Lawrence in implementing the Geant4 code.

References 1. Handbook of Radiological Protection. Part 1: Data. SBN 11 360079 8. London, UK: HMSO, 1971. 2. Recommendation for Data on Shielding from Ionizing Radiation, Part 2 Shielding from X-radiation. British Standards: BS4094. ISBN 058006522-7. London, UK: British Standards Institution, 1971. 3. GEANT4 (Agostinelli S, Allison J, Amako K, Apostolakis J, Araujo H, Arce P, et al), http://cern.ch/geant4 [Accessed 18 September 2006]. 4. Sheikh-Bagheri D, Rogers DW. Monte Carlo calculations of nine megavoltage photon beam spectra using the BEAM code. Med Phys 2002;29:391–402.

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SHORT COMMUNICATION

The use of megavoltage cone-beam CT to complement CT for target definition in pelvic radiotherapy in the presence of hip replacement M AUBIN, MSc, O MORIN, BSc, J CHEN, PhD, A GILLIS, MD, B PICKETT, MSc, J F AUBRY, C AKAZAWA, CMD, J SPEIGHT, MD, M ROACH III, MD and J POULIOT, PhD

MSc,

Department of Radiation Oncology, University of California San Francisco, Comprehensive Cancer Center, San Francisco, CA 94143, USA ABSTRACT. In Europe and the USA combined, over half a million people had a hip joint replaced in 2005, contributing to the increasing number of radiotherapy patients with metallic hip prostheses. The treatment plan for external beam radiation therapy is based on the delineation of the anatomy in the planning CT scan. When implanted objects of high atomic number (Z) material are present, however, severe image artefacts are generated in conventional CT, strongly hindering the ability to delineate some organs. This is particularly the case for the planning of prostate patients with hip prostheses. This short communication presents the use of a new imaging modality, megavoltage cone-beam CT, to complement the regular CT for target definition of prostate cancer treatment of patients with hip replacements.

Treatment planning for external beam radiation therapy is based on the delineation of the anatomy visualized in the planning CT scan. However, image artefacts caused by the presence of a hip replacement often render CT images useless for prostate delineation (Figure 1, left) and preclude precise dose calculation. It was recently suggested that CT-MR image registration could facilitate target definition for a prostate patient with hip replacements [1]. Effectively, MR images provide an accurate definition of the clinical target volume (CTV) and better visualization of normal structures. A number of factors, however, were found to affect image quality and/or accuracy of target definition. The standard MR couch, different from a CT or linac treatment couch, might result in different patient position, and the presence of the metallic implants may create significant distortion. Finally, in the presence of a hip replacement, neither the regular CT nor the MR can provide accurate electron density information for heterogeneous dose calculation. Promising artefact reduction techniques on regular CT are being developed to minimize the impact of streaking artefacts [2]. At that time, they may require manual image post-processing and most CT scanners available in radiation oncology departments are not equipped with these features. We report on the use of megavoltage cone-beam CT (MV CBCT) to complement the conventional CT for Address correspondence to: Dr Jean Pouliot , Department of Radiation Oncology, University of California San Francisco, 1600 Divisadero Street, Suite H1031, San Francisco, CA 94941-1708, USA. E-mail: [email protected]. This research was supported by Siemens Oncology Care Systems.

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Received 20 January 2006 Revised 22 June 2006 Accepted 4 July 2006 DOI: 10.1259/bjr/19559792 ’ 2006 The British Institute of Radiology

target organ definition in seven patients treated for prostate cancer with external beam radiation therapy. In this study, we exploited the predominantly Compton scattering of high-energy photons delivered in MV CBCT systems to obtain 3D images of the anatomy in the presence of unilateral or bilateral hip replacements and complement the planning CT during the target delineation process.

Method A cone-beam CT image is reconstructed from a large set of projection images acquired at various angles around a patient in a process similar to that of conventional CT. In cone-beam CT, a two-dimensional (2D) array of detectors, in our case a portal imager attached to the linear accelerator (linac), is used to reconstruct a three-dimensional (3D) image. For MV CBCT, the 6 MV treatment beam of the accelerator, containing photons primarily in the MeV range, is used for the imaging. A MV CBCT system [3, 4] integrated onto an ONCOR clinical Linac (Siemens Oncology Care Systems, Concord, CA) was used to acquire 3D images in treatment position for seven prostate patients with unilateral or bilateral hip prostheses. MV CBCT acquisitions were performed by rotating the linac in a continuous 200 ˚ arc (270 ˚ to 110 ˚, clockwise) acquiring one portal image for each degree. Because MV CBCT uses the treatment beam, the treatment planning system (Pinnacle, Phillips, Best, The Netherlands) was used to evaluate the dose delivered during an MV CBCT The British Journal of Radiology, November 2006

Short communication: Target definition in pelvic radiotherapy with MV CBCT

Figure 1. Comparison of a conventional CT (left) and megavoltage cone beam CT (MV CBCT) (right). (A) Axial and (B) coronal views are shown for a unilateral hip replacement. (C) Axial view for a bilateral hip replacement.

acquisition. For a typical acquisition procedure, the dose at the isocentre was 0.05 Gy and the acquisition lasted 45 s. The reconstructed image, a typical 25662566274 volume (27.4 cm627.4 cm627.4 cm), was available less than 2 min after the acquisition of the first portal image. A first order scatter correction was applied to facilitate the adjustment of window/levels. For each patient, the MV CBCT images were imported into the treatment planning system and registered with the original planning CT using bony anatomy contoured on each image set. The target volumes and organs at risk for prostate treatment were contoured using both the CT and the MV CBCT for single hip replacement, and using only the MV CBCT for bi-lateral hip prostheses.

Results The MV CBCT images could be used to visualize clearly the hip prosthesis and bony anatomy and provide sufficient soft-tissue contrast to help delineate the The British Journal of Radiology, November 2006

prostate, bladder and rectum. The artefacts on the regular CT obscure the border between the prostate and anterior wall of the rectum (Figure 1A, left) and the interface between the prostate base and the bladder neck (Figure 1B, left). The MV CBCT images were particularly useful to help delineate these structures as well as the lateral extension of the prostate in the axial plane, the seminal vesicles and the lymph nodes. Also, normal anatomy such as pelvic bones, penile bulb, bladder, femoral heads, rectum and small bowel can be delineated with higher accuracy as well. An example of organ segmentation is presented in Figure 2. The change of shape of the prostate (red), bladder (yellow) and rectum (blue) between the CT (left) and the MV CBCT (right) can be easily observed. In this study, the time intervals between the CT and the MV CBCT ranged from a few hours to 1 week. By itself, this can explain the change in rectum and bladder volumes due to different fillings. For the six patients with a single hip replacement, the posterior limit of the prostate was found to be more anterior, and therefore the prostate 919

M Aubin, O Morin, J Chen et al

Figure 2. Segmentation of the bladder (yellow), prostate (red) and rectum (bule) using the conventional CT (left) and the megavoltage cone beam CT (MV CBCT) (right) shown on the (A) axial, (B) sagittal and (C) coronal views.

volumes contoured with the help of MV CBCT were generally smaller than the volumes that would have been estimated using only the regular CT containing severe artefacts. These smaller prostate volumes may prevent overdosage of the rectum. Target delineation for the patient with bilateral hip prostheses was entirely performed using the MV CBCT, since the relevant organs were totally obscured due to the severity of the artefacts on the conventional CT.

Discussion Image artefacts (Figures 1 and 2, left) caused by the presence of hip replacements render regular CT images difficult to use for prostate delineation. In contrast, the presence of high-Z material has relatively little impact on the image quality of MV CBCT. Tests performed on phantoms [5] showed that the presence of a metallic object strongly impacts on Hounsfield numbers (up to 70% error) of a conventional CT image and therefore disturbs the electron density even far away from the object, making CT inaccurate for dose calculation. For this reason, CT treatment plans of patients with hip replacements are generally produced without density correction. Similar tests performed with MV CBCT on a phantom with and without a metallic object demonstrated that Hounsfield numbers remain unchanged (within 3%) in the presence of metallic objects, allowing for significantly more accurate dose calculation. Thus, the next step after using MV CBCT for image segmentation will be to use MV CBCT for dose calculation. Research to develop calibration procedures 920

to use MV CBCT for dose calculations is being performed. There is also ongoing technical development to increase the field of view of the current version of MV CBCT (27 cm627 cm) to encompass the entire pelvic region. The possibility of considering dose escalation protocols depends on the ability to identify the prostate volume and critical structures for treatment planning, and the accuracy of the dose calculation. Because of degradation of image quality in the presence of high-Z material, treatment planning and dose calculations are limited in these settings. Consequently, patients with hip prostheses may not be candidates for advanced treatment planning like intensity-modulated radiotherapy (IMRT). With the advance capability of using MV CBCT in the treatment planning software, it is now possible to deliver accurately higher doses of radiation to the prostate in patients with hip prostheses. While MV CBCT acquisition dose is two to three times higher than a conventional CT, it represents only a very small fraction (,0.1%) of the treatment dose. With dose escalation these patients may benefit from a decreased risk of treatment failure. MV CBCT in the presence of high-Z material may improve treatment planning, allowing patients with pelvic malignancies and hip prostheses the most advanced form of radiation therapy.

Conclusion MV CBCT provides 3D anatomical information of the patient in the treatment position, even in the presence of The British Journal of Radiology, November 2006

Short communication: Target definition in pelvic radiotherapy with MV CBCT

‘‘CT non-compatible’’ objects. MV CBCT registered with the planning CT can complement missing information and facilitate segmentation for planning purposes when hip prostheses are present.

References 1. Charnley N, Morgan A, Thomas E, Wilson S, Bacon S, Wilson D, et al. The use of CT-MR image registration to define target volumes in pelvic radiotherapy in the presence of bilateral hip replacements. Br J Radiol 2005;78:634–6. 2. Yazdia M, Gingras L, Beaulieu L. An adaptive approach to metal artifact reduction in helical computed tomography


for radiation therapy treatment planning: experimental and clinical studies. Int J Radiat Oncol Biol Phys 2005;62: 1224–31. 3. Pouliot J, Bani-Hashemi A, Chen J, et al. Low-dose megavoltage cone-beam CT for radiation therapy. Int J Radiat Oncol Biol Phys 2005;61:238–46. 4. Morin O, Gillis A, Chen J, Aubin M, Bucci MK, Pouliot J. Megavoltage cone-beam CT: system description and IGRT clinical applications. Special issue on image-guided radiation therapy (IGRT). Med Dosim 2006;31:51–61. 5. Aubin M, Morin O, Bucci K, Chan A, Chen J, Ghelmansarai F, et al. Megavoltage conebeam CT to complement prostate planning CT in presence of hip prosthesis. ESTRO Annual Meeting (abstract), Lisbon, 2005.

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PICTORIAL REVIEW

CT of thoracic lymph nodes. Part I: anatomy and drainage 1

T SUWATANAPONGCHED,

MD

and 2D S GIERADA,

MD

1

Department of Radiology, Faculty of Medicine, Ramathibodi Hospital, Mahidol University, 270, Rama VI Road, Rajthevi, Bangkok 10400, Thailand and 2Mallinckrodt Institute of Radiology, Washington University School of Medicine, 510 South Kingshighway Blvd, Campus Box 8131, Saint Louis, Missouri 63110, USA ABSTRACT. CT is the primary non-invasive technique for the diagnostic evaluation of thoracic lymph nodes. The CT patterns and anatomic location of thoracic lymph node involvement can provide important clues in the diagnosis of many diseases. Part I of the pictorial review illustrates the anatomic location and drainage of thoracic lymph nodes in the chest wall, mediastinum and lungs through examples of pathologic involvement. Part II of the pictorial review focuses on CT patterns of lymph node involvement in various pulmonary and extrapulmonary diseases, differential diagnoses based on CT findings and pitfalls.

CT is the primary non-invasive technique for the diagnostic evaluation of thoracic lymph nodes. Lymph node abnormalities are depicted by CT as an increase in nodal size and/or number or change in attenuation. Although these findings are non-specific, patterns of thoracic lymph node involvement can provide important clues in the diagnosis of many pulmonary and extrapulmonary diseases. Part I of this pictorial review illustrates the anatomic location and drainage of thoracic lymph nodes in the chest wall, mediastinum and lungs through examples of pathologic involvement. Part II focuses on CT patterns of lymph node involvement in various pulmonary and extrapulmonary diseases.

Classification of thoracic lymph nodes As in other parts of the body, thoracic lymph nodes are named using descriptive terminology according to the blood vessels or visceral structures to which they are most closely related, or by their general anatomic location. Although there are slight differences in the classification of the thoracic nodes [1–5], they can be divided into those of the chest wall and those of the intrathoracic contents. To facilitate accurate pathologic staging and analysis of treatment outcomes in lung cancer, a classification scheme for mediastinal and pulmonary lymph nodes (Figure 1) has been devised by the American Joint Committee on Cancer (AJCC) and the Union Internationale Contre le Cancer (UICC) [5], based on surgically recognizable anatomic landmarks.

Chest wall nodes The axillary nodes (Figures 2 and 3) receive superficial lymphatic drainage from the upper limbs, breasts and Address correspondence to: D S Gierada.

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Received 11 April 2005 Revised 23 June 2005 Accepted 11 July 2005 DOI: 10.1259/bjr/26411607 ’ 2006 The British Institute of Radiology

pectoral muscles excluding their medial portions, parietal pleura, and skin and muscles of the trunk above the umbilicus and iliac crest [1, 2]. The lymph flow is directed toward the terminal nodal group in the axillary apices. The efferent vessels from this group unite as the subclavian trunk, which finally drains directly or indirectly into the jugulo-subclavian venous confluence [1, 2, 6]. A few efferents usually reach the supraclavicular nodes, a well-recognized route for the spread of breast cancer [1, 2, 6]. The internal mammary (internal thoracic or parasternal) nodes (Figure 4) lie at the anterior ends of the intercostal spaces, along the internal mammary (internal thoracic) vessels. They receive lymphatic drainage from the anterior diaphragmatic nodes, anterosuperior portion of the liver, medial part of the breasts, and deeper structures of the anterior chest and upper anterior abdominal wall [2]. Their efferent channels may empty into the right lymphatic duct, the thoracic duct, or the inferior deep cervical nodes [3, 6]. The posterior intercostal nodes (Figures 5 and 6), located near the heads and necks of the posterior ribs, receive lymphatic drainage from the posterolateral intercostal spaces, posterolateral breasts, parietal pleura, vertebrae and spinal muscles [2–4]. The efferent vessels from the upper intercostal spaces end in the thoracic duct on the left, and in one of the lymphatic trunks on the right [2–4]. Those from the lower four to seven intercostal spaces unite to form a common trunk, which empties into the thoracic duct or cisterna chyli [2–4]. The juxtavertebral (pre-vertebral or paravertebral) nodes lie along the anterior and lateral aspects of the vertebral bodies, most numerous from T8 to T12 (Figures 5 and 6) [3, 4]. They communicate with posterior mediastinal lymph nodes [3] and the posterior intercostal nodes, and similarly drain to the right lymphatic duct or thoracic duct [3, 4]. The British Journal of Radiology, November 2006

Pictorial review: CT of thoracic lymph nodes

(a)

(b)

Figure 1. Revised American Joint Committee on Cancer (AJCC) and the Union Internationale Contre le Cancer (UICC) regional nodal stations for lung cancer staging. (From Mountain CF, Dresler CM. Regional lymph node classification for lung cancer staging. Chest 1997;111:1718–23 [5]. Reprinted with permission). (a) Drawing illustrates mediastinum lymph node stations in the frontal projection. Ao 5 aortic arch, PA 5 main pulmonary artery, 1 (red) 5 highest mediastinal nodes, 2R and 2L (dark blue) 5 right and left upper paratracheal nodes, 4R and 4L (orange) 5 right and left lower paratracheal nodes, 7 (blue) 5 subcarinal nodes, 8 (grey) 5 para-oesophageal nodes, 9 (brown) 5 pulmonary ligament nodes, 10R and 10L (yellow) 5 right and left hilar nodes, 11R and 11L (green) 5 right and left interlobar nodes, 12R and 12L (pink) 5 right and left lobar nodes, 13R and 13L (pink) 5 right and left segmental nodes, 14R and 14L (pink) 5 right and left subsegmental nodes. (b) Illustration of mediastinum lymph node stations in the left anterior oblique projection. Ao 5 aortic arch, PA 5 main pulmonary artery, 3 (pink) 5 pre-vascular and retrotracheal nodes, 5 (black) 5 subaortic nodes, 6 (red) 5 para-aortic nodes.

The diaphragmatic nodes are located on or just above the thoracic surface of the diaphragm and are divided into three groups [2–4, 7]. The anterior (pre-pericardial or cardiophrenic) group (Figure 6) is located anterior to the pericardium, posterior to the xiphoid process, and in the right and left cardiophrenic fat. This node group receives afferent drainage from the anterior part of the diaphragm and its pleura, and the anterosuperior portion of the liver. They drain to the internal mammary nodes alongside the xiphoid and can provide a route for retrograde spread of breast cancer to the liver, via lymphatics of the rectus abdominis muscle when the upper internal thoracic trunks are blocked [4]. The middle (juxtaphrenic or lateral) (Figure 7) group receives lymph from the central diaphragm and from the convex surface of the liver on the right [2]. The posterior (retrocrural) group (Figure 8), lying behind the diaphragmatic crura and anterior to the spine, receives lymph from the posterior part of the diaphragm and communicates with the posterior mediastinal nodes and para-aortic nodes in the upper abdomen [2, 4]. When diaphragmatic nodes are enlarged, widespread disease in other locations is usually present, so biopsy of these sites is uncommon [7]. The British Journal of Radiology, November 2006

Mediastinal lymph nodes Anterior mediastinal group This group includes the highest mediastinal (station 1, Figures 1 and 3a), pre-vascular (station 3A, Figures 1 and 3b), and para-aortic (station 6, Figures 1 and 9)

Figure 2. Enhanced CT scan in a 66-year-old woman with lymphoma showing multiple enlarged bilateral axillary lymph nodes (arrows).

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T Suwatanapongched and D S Gierada

(a)

(b)

Figure 3. A 65-year-old man with chronic lymphocytic leukaemia. (a) Enhanced CT scan demonstrates enlarged right axillary nodes (arrowheads) and right interpectoral (Rotter) node (black arrow) lying between pectoralis major (M) and minor (m) muscles. Nodes in the subpectoral and interpectoral regions are included in the axillary nodal group. Also seen are enlarged highest mediastinal nodes (station 1; white arrows) defined by their location cranial to the superior margin of the left brachiocephalic vein, behind and to the right and left sides of the trachea. (b) Enhanced CT scan at the lower level shows bilaterally enlarged axillary nodes (arrowheads), including left subpectoral nodes (open arrow) underneath the left pectoralis minor muscle (m). There are enlarged pre-vascular nodes (station 3A; white arrows), which lie between the superior margin of the left brachiocephalic vein (V) and the superior margin of the aortic arch, and anterior to its large arterial branches; enlarged retrotracheal node (station 3P; black arrow), which lies behind the trachea and above the inferior aspect of azygos vein arch; and enlarged right upper paratracheal nodes (station 2R; wavy arrow), which are located above the superior margin of the aortic arch.

Figure 5. Enhanced CT scan of a 31-year-old man with Figure 4. Enhanced CT scan at the level of the main pulmonary artery in a 55-year-old woman with left breast cancer demonstrating enlarged left internal mammary node (arrow). Note normal right internal mammary vessels (wavy arrow) and a portion of primary cancer in the left breast (asterisk).

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lymphoma showing enlarged, necrotic right and left intercostal nodes (white arrows) as well as enlarged left paravertebral (arrowheads) and retrocrural (black arrows) nodes. Note a left pleural effusion (E) with pleural nodules (small white arrows), splenectomy clips and coeliac adenopathy (N). A 5 aorta.



Figure 6. Enhanced CT scan in a 69-year-old woman with lymphoma showing enlarged bilateral paravertebral nodes (white arrows), left intercostal node (open arrow) and anterior diaphragmatic nodes (black arrows). Note bilateral pleural effusions (E).

nodes [2, 3, 5, 8]. They receive afferent vessels from the thymus, thyroid, heart and pericardium, diaphragmatic and mediastinal pleura, and middle diaphragmatic nodes [2, 3]. Their efferent channels join those from the paratracheal, tracheobronchial and internal mammary nodes, to form the right and left bronchomediastinal trunks, which may empty to the right lymphatic duct, the thoracic duct, or open independently into the jugulosubclavian venous confluence [2, 3].

Paratracheal and tracheobronchial groups These groups receive drainage from most parts of the lungs and bronchi, thoracic trachea, heart and some efferents from the upper para-oesophageal nodes of the posterior mediastinal group [2, 4]. The nodes comprising

Figure 8. CT scan through the upper abdomen in a 45-yearold man with distal oesophageal carcinoma (not shown) revealing enlarged retrocrural lymph nodes (large arrows) and liver metastases (small arrows).

these groups include the upper (station 2R, 2L, Figures 1 and 3b) and lower (station 4R, 4L, Figures 1, 9 and 10) paratracheal, subaortic (aortopulmonary window, station 5, Figures 1, 11, and 12), retrotracheal (station 3P, Figures 1 and 3b), and subcarinal (station 7, Figures 1 and 12) nodes [2, 3, 5]. The azygos node, located medial to the azygos arch, is included in station 4R [5]. The upper paratracheal nodes link the lower paratracheal and inferior deep cervical nodes [10]. The subcarinal nodes are contiguous with the hilar nodes and drain to the paratracheal nodes, preferentially to the right [11].

Figure 9. Non-enhanced CT scan in the same patient as in Figure 7. Non-enhanced CT scan in a 28-year-old woman with metastatic papillary serous adenocarcinoma of the ovary revealing enlarged, densely calcified right middle diaphragmatic nodes (arrow), located lateral to the intrathoracic end of the inferior vena cava (V) and near the insertion of the right phrenic nerve.


Figure 7 revealing enlarged, calcified para-aortic nodes (station 6; arrows), lying anterior and lateral to the aortic arch (A) below its superior margin. Also seen is right lower paratracheal lymphadenopathy (station 4R; open arrow). V 5 superior vena cava. (From Glazer HS, Molina PL, Siegel MJ, Sagel SS. High-attenuation mediastinal masses on unenhanced CT. AJR Am J Roentgenol 1991;156:45–50 [8]. Reprinted with permission).

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Figure 10. Enhanced CT scan in a 73-year-old man with left lower lobe lung cancer (not shown) showing enlarged right lower paratracheal nodes (large arrow) lying medial to the azygos vein (V) and enlarged left lower paratracheal nodes (station 4L; open arrow) lying medial to ligamentum arteriosum (small arrows). Lower paratracheal nodes lie caudal to the top of the aortic arch. (From Sagel SS, Slone RM. Lung. In: Lee JKT, Sagel SS, Stanley RJ, Heiken JP, editors. Computed body tomography with MRI correlation, 3rd edn. Philadelphia, PA: Lippincott-Raven Publishers, 1998:351–454 [9]. Reprinted with permission).

Figure 12. Enhanced CT scan in a 65-year-old man with diffuse pulmonary lymphangitic carcinomatosis secondary to non-small cell lung cancer (not shown) demonstrating enlarged subcarinal (station 7; curved arrow), para-oesophageal (black arrow), right hilar (station 10R; large white arrows) and left hilar (station 10L; open arrow) nodes. Hilar nodes are outside the mediastinal pleura, below the top of the upper lobe bronchi. Note enlarged subaortic (arrowhead) and para-aortic (small white arrow) nodes. Oe 5 oesophagus.

Figure 13. Enhanced CT scan in a 65-year-old man with nonFigure 11. Enhanced CT scan in a 58-year-old woman with carcinoid tumour showing enhancing subaortic lymphadenopathy (station 5; arrows) within the aortopulmonary window region. This group is located lateral to the ligamentum arteriosum (not seen). Note primary tumour in the left upper lobe (open arrow).

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small cell lung cancer demonstrating metastasis to left pulmonary ligament node (station 9; curved arrow) from left lower lobe lung cancer (straight arrow). Oe 5 oesophagus, A 5 aorta. (From Sagel SS, Slone RM. Lung. In: Lee JKT, Sagel SS, Stanley RJ, Heiken JP, editors. Computed body tomography with MRI correlation, 3rd edn. Philadelphia, USA: Lippincott-Raven Publishers, 1998:351–454 [9]. Reprinted with permission).



Figure 14. Enhanced CT scan in a 29-year-old woman with sarcoidosis demonstrating enlarged right lobar node (station 12R; arrowhead) at the bifurcation of the bronchus intermedius, right segmental node (open arrow) adjacent to the right middle lobe lateral segmental bronchus, and left interlobar nodes (station 11R and 11L; white arrows) between the lingular and left lower lobe superior segmental bronchus. Note enlarged subcarinal nodes (black arrows) and bilateral pulmonary involvement.

Thus, the left lower lobe is the most common primary site for contralateral mediastinal lymph node metastasis in lung cancer.

Posterior mediastinal group The posterior mediastinal nodes are comprised of the para-oesophageal (station 8, Figures 1 and 12) and pulmonary ligament (station 9, Figures 1 and 13) nodes [4, 5]. The para-oesophageal nodes receive afferent vessels from the thoracic oesophagus, posterior pericardium, diaphragm, posterior diaphragmatic nodes and

Figure 16. Axial CT scan with lung-window setting in a 59year-old man with myocardial infarction showing a 1 cm, indeterminate, solitary pulmonary nodule containing an eccentric calcific focus in the right middle lobe (arrow). Wedge resection revealed a subsegmental lymph node (station 14R) with calcified granuloma.

the left hepatic lobe, and are more numerous on the left [2]. The pulmonary ligament nodes receive drainage from the basilar segments of the lower lobes and lower half of the oesophagus [4]. The efferents from the posterior mediastinal nodes communicate with the tracheobronchial group, particularly subcarinal nodes, and drain chiefly into the thoracic duct, but also drain to the subdiaphragmatic para-aortic or coeliac nodes [3, 4].

Lymph nodes of the lungs

Figure 15. Enhanced CT scan in the same patient as in Figure 12 showing enlarged right and left segmental nodes (station 13R and 13L; large white arrows) lying adjacent to the segmental bronchi (small white arrows) and enlarged para-oesophageal nodes (black arrows).


Lymph nodes are located along the bronchi and can be divided into hilar (station 10R, 10L, Figures 1 and 12) and intrapulmonary nodes [5, 10]. The latter consist of interlobar (station 11R, 11L, Figures 1 and 14), lobar (station 12R, 12L, Figures 1 and 14), segmental (station 13R, 13L, Figures 1, 14 and 15), subsegmental (station 14R, 14L, Figures 1 and 16) and intraparenchymal intrapulmonary (Figure 17) nodes [5, 9, 12]. Most of the lymphatic flow of the lungs is directed toward the interlobar and hilar nodes, which drain into the subcarinal nodes or directly into the lower paratracheal nodes [3, 4, 10, 11]. The normal hilar and interlobar nodes are frequently visible, particularly with thinner (1–3 mm) collimation and intravenous contrast administration [13]. Recognition of these nodes is important to avoid misdiagnosis of pulmonary embolism. Intraparenchymal intrapulmonary nodes may present as indeterminate subpleural pulmonary nodules in the lower parts of the lungs [12]. 927


(a)

(b)

Figure 17. A 58-year-old man with bronchioloalveolar carcinoma of the left upper lobe (not shown). (a) CT scan with lungwindow setting demonstrates a tiny, subpleural nodule in the lingular segment (arrow). (b) Histological examination reveals a normal lymph node (arrows), surrounded by alveolar tissue. It had capsule with visible germinal centres and contains histiocytes and carbon pigment (haematoxylin and eosin 640).

References 1. Johnson D, Ellis H. Pectoral girdle and upper limb. In: Standring S, editor. Gray’s anatomy, 39th edn. Edinburgh, Scotland: Churchill Livingstone, 2005:817–49. 2. Johnson D, Shah P. Thorax. In: Standring S, editor. Gray’s anatomy, 39th edn. Edinburgh, Scotland: Churchill Livingstone, 2005:951–1079. 3. Fraser RS, Mu¨ller NL, Colman N, Pare´ PD. The lymphatic system of the lungs, pleura, and mediastinum. In: Fraser RS, Mu¨ller NL, Colman N, Pare´ PD, editors. Fraser and Pare´’s diagnosis of diseases of the chest, 4th edn. Philadelphia, PA: W.B. Saunders Company, 1999:172–95. 4. Weinberg JA. The intrathoracic lymphatics. In: Haagensen CD, editor. The lymphatics in cancer. Philadelphia, PA: W.B. Saunders Company, 1972:231–99. 5. Mountain CF, Dresler CM. Regional lymph node classification for lung cancer staging. Chest 1997;111:1718–23. 6. Haagensen CD. Lymphatics of the breasts. In: Haagensen CD, editor. The lymphatics in cancer. Philadelphia, PA: W.B Saunders Company, 1972:300–98. 7. Aronberg DJ, Peterson RR, Glazer HS, Sagel SS. Superior diaphragmatic lymph nodes: CT assessment. J Comput Assist Tomogr 1986;10:937–41.

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8. Glazer HS, Molina PL, Siegel MJ, Sagel SS. High-attenuation mediastinal masses on unenhanced CT. AJR Am J Roentgenol 1991;156:45–50. 9. Sagel SS, Slone RM. Lung. In: Lee JKT, Sagel SS, Stanley RJ, Heiken JP, editors. Computed body tomography with MRI correlation, 3rd edn. Philadelphia, PA: Lippincott-Raven Publishers, 1998:351–454. 10. Nohl-Oser HC. Surgery of the lung, general management and operative technique. In: Nohl-Oser HC, editor. Surgery of the lung. New York, NY: Thieme-Stratton Inc., 1981:38– 184. 11. Shields TW. Mediastinal lymph nodes. In: Shields TW, editor. Mediastinal surgery. Philadelphia, PA: Lea & Febiger, 1991:14–8. 12. Bankoff MS, McEniff NJ, Bhadelia RA, Garcia-Moliner M, Daly BDT. Prevalence of pathologically proven intrapulmonary lymph nodes and their appearance on CT. AJR Am J Roentgenol 1996;167:629–30. 13. Shimoyama K, Murata K, Takahashi M, Morita R. Pulmonary hilar lymph node metastases from lung cancer: evaluation based on morphology at thinsection, incremental, dynamic CT. Radiology 1997;203: 187–95.



Book reviews Clark’s Positioning in Radiography, 12th edn. By A Stewart Whitley, Charles Sloane, Graham Hoadley, Adrian D Moore and Chrissie W Alsop. pp. 544, 2005 (Hodder Arnold, London, UK) £75.00 ISBN 0-340-76390-6 By any standards, Kathleen Clara Clark (1896–1968) was a remarkable woman. She entered radiography as the profession was developing. The Society of Radiographers had been set up in 1920 and letters had been written from the new society to the assistants in the various X-ray departments inviting applications for membership. Those who had been in active practice for over 10 years were given membership without examination; all other applicants had to take a new examination. The first group comprising 45 students was entered for examination in January 1922, 20 of whom passed and were duly awarded the certificate of the new Society (the MSR). Miss ‘Katie’ Clark completed her training course at Guy’s Hospital in 1921 and passed the first qualifying examination ever held by the Society of Radiographers. She then worked (1922–1927) at the Princess Mary’s Hospital (for Surgical Tuberculosis) and Margate (General) Hospital before moving to the Royal Northern Hospital in London (1927–1935). Kathleen Clark was aware of the lack of adequate training for radiographers, and so founded a School of Radiography at the Royal Northern Hospital, which became a model for schools elsewhere. She was also tutor to the X-ray Training School at the Royal Northern from 1930-1935. In 1935 she became co-founder and Radiographer-incharge of the Ilford Radiographic Technical and Demonstration Department based at Tavistock House, where she was involved in instruction and research into radiography and medical photography. Under her guidance, the department developed a worldwide reputation. Kathleen Clark was President of the Society of Radiographers from 1935 to 1937 (and the first woman President). She was also the first President to wear the President’s chain of office, having been presented with it by her predecessor Dr Leo Rowden. The first edition of her classic book Positioning in Radiography (Heinemann Medical Books) was published in 1939. The book became the standard work of reference for radiographers and has been through many editions, most of which are held in the Information Centre at the British Institute of Radiology. Generations of radiographers will have used the textbook. Her slide collection has been preserved and is also available in the Information Centre at the British Institute of Radiology as the K C Clark Slide Library. Although primarily of historical interest, the slide collection remains a useful teaching resource and may be copied. The original Positioning in Radiography is a very remarkable book for several reasons. First, Katie Clark was keen to standardize both positioning and exposure, and in her book she standardized the radiographic The British Journal of Radiology, November 2006

projections so that similar projections could be made in all hospitals. Second, the book is very artistic. The illustrations do not come across as cold and entirely objective scientific images. It is therefore not surprising that the artist Francis Bacon acknowledged Positioning in Radiography as a crucial source of inspiration and that it was his favorite medical textbook. Lawrence Gowling indicated that Bacon repeatedly borrowed from the photographs in the book for his work. The images of the body that Francis Bacon made have an almost radiographic quality and there is the impression that multiple layers of the body are seen at the same time and that one is not just looking at the skin’s surface. This was well illustrated by Professor Robert Clark from the University of South Florida in his inaugural lecture ‘‘The Art of Radiography: how Positioning in Radiography by radiographer Kathleen Clara Clark influenced the art of Francis Bacon’’, for the British Society for the History of Radiology held recently at the British Institute of Radiology. Kathleen Clark was awarded the MBE in 1945 for her services to radiography, particularly for mass miniature radiography of the chest. Her book Mass Miniature Radiography of Civilians (MRC Special Report, series No. 251) was written jointly with P D’Arcy Hart, Peter Kerley and Brian Thompson, appeared in 1945 and was the definitive guide to mass radiography of the chest. She was committed to fostering co-operation and contact between radiographers throughout the world, and was a driving spirit behind the formation of the International Society of Radiographers and Radiological Technologists (ISRRT). She remained as principal at Ilford until 1958 and acted as Consultant in Radiography until 1964. The story of her life remains an inspiration to radiographers everywhere. In front of me is the first edition of Positioning in Radiography (1939) and the new Clark’s Positioning in Radiography’ (2005). The new book certainly feels very different from the first edition; however, comparisons are always difficult. It is rather like comparing an ‘‘Old Master’’ such as the Flemish baroque painter Peter Paul Rubens (1577–1640), who was probably the most renowned northern European artist of his day, to modern artists such as the American pop artist Andy Warhol (1928–1987) or to the contemporary British conceptual artist Tracey Emin (born 1963). A comparison can be made, but is any comparison really meaningful? Certainly, the new work by Stewart Whitley and his coauthors feels less artistic that the first edition. However, the world has changed and a different approach is needed. The new book is an excellent account of current plain radiographic techniques in a single volume. Techniques that are no longer used or that are associated with a high radiation dose to the patient have been removed from the text and a modern alternative is suggested. Many of the radiographic positions in the older editions are no longer used. There are excellent sections on mammography and dental radiography, and there is a new 929

Book reviews

section on forensic radiography. The trauma section is excellent and there is a reference to Advanced Trauma Life Support (ATLS). The introduction of digital radiography has profoundly affected our departments and this is acknowledged in the book. The expression ‘‘film’’ has been replaced by ‘‘cassette’’ or ‘‘‘image receptor’’. I would have liked to see a little more detail on post-processing and image handling, and marking in a digital environment. The radiographic images in the modern text have less contrast than the images in the first edition; however, the earlier images were obtained using old high silver film. The older images look very different from our modern digital images. I am pleased that the book also reflects the extended role of the technologist/radiographer and there is a useful ‘‘Radiological Considerations’’ sub-heading to all of the chapters, giving the clinical context and an understanding of image interpretation requirements. The anatomical diagrams are good and the labelling is clear. I showed the book to the radiographers in my department and the book was universally greeted with enthusiasm. Everyone thought that the department needed a new book on radiographic technique to reflect modern practice. It is easy for a department to feel that as they already have a radiographic positioning book, they do not need a new one since positioning does not change. This is not really the case as techniques develop and are refined as time passes. Whilst we should keep our old volumes of Positioning in Radiography (and many with missing pages), I would recommend that all departments buy a copy of Clark’s Positioning in Radiography – my hospital Trust has ordered four copies for our various departments.

A THOMAS MRI of the Musculoskeletal System, 5th edn. By TH Berquist. pp. 1008, 2005 (Lippincott Williams & Wilkins, Philadelphia, PA) $199.00 ISBN 0-7817-5502-6 This is an updated edition of a well-established text of MRI of the musculoskeletal system. This is essentially an in-house text with sixteen of the eighteen contributors coming from the Mayo clinic. The text and most of the images have been updated from the fourth edition, which was published in 2001. The book is an excellent authoritative text on MR of the appendicular system. It has a traditional textbook format, with the chapters being divided mainly along anatomical lines. The book is quite comprehensive, covering most conditions that one is likely to see in a musculoskeletal practice. Loose ends are dealt with in chapters on musculoskeletal neoplasms, infections, marrow disease and a final chapter entitled ‘‘Miscellaneous conditions’’. The book has many strengths, not least the comprehensive and readable initial chapters on the basic principles of MRI. Chapters are introduced with good descriptions of the basic anatomy and are illustrated with good quality images.

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If you are looking for an excellent reference book on MR of the appendicular musculoskeletal system, you will not be disappointed. However, do not expect anything more than a tome dedicated to MRI: what you see on the title is what you get. Only lip service is paid to the varying merits of MRI with other imaging techniques or the clinical application of MRI in everyday practice. The chapter on the spine is written by neuroradiologists and there are large gaps related to musculoskeletal aspects of spinal imaging. There is no mention of some conditions that are frequently encountered in a musculoskeletal practice, such as ankylosing spondylitis and pars defects. Primary bone tumours of the spine such as osteoblastoma, giant cell tumour or chordoma do not feature. There is a disproportionate emphasis on intradural disease, reflecting the interest of the authors. I judge this to be a first class reference book for MRI of the musculoskeletal system, although the spine chapter would benefit from some input by a musculoskeletal radiologist to fill some significant gaps. The book represents excellent value, being priced at just under $200.00.

S OSTLERE NeuroPET: Positron emission tomography in neuroscience and clinical neurology. By Karl Herholz, Peter Herscovitch and Wolf-Dieter Heiss. pp. xv+297, 2004 (Springer-Verlag, Heidelberg, Germany) ISBN 3-540-00691-5 This is a very well prepared and concise book at just over 250 pages. Very heavily referenced, it provides the reader with an excellent and well-illustrated introduction to the concepts associated with the use of positron emission tomography (PET) in the human brain. It has a number of sections; in the first part of the book, these are related to individual disease states, with chapters on dementia, movement disorders, brain tumours, etc. The next section is related to the imaging of brain function, with specific chapters on blood flow, glucose consumption, oxygen consumption, amino acid transport, etc. The third and final portion of the book is technical, dealing with the data acquisition, reconstruction, modelling and statistics associated with the collection of PET data. A colleague once told me that PET is the technology of the future and always will be. At least with this book you can have some understanding as to the basic concepts, applications and potential values of this increasingly important clinical methodology. The book is probably too detailed for the clinicians starting to do a little bit of PET reporting, but I think those are few and far between with the majority of people involved in this technology wanting to understand it at a significant level, which this book is ideal for. The authors are to be congratulated on putting together a very concise and well presented book, which I would recommend to anyone involved in PET imaging. A JACKSON



CASE OF THE MONTH

Various holes and lesions 1

S PUNEKAR,

MBBS, MRCP,

2

D J A BUTTERISS,

BSc, MBBS (Hons), FRCR

and 2D BIRCHALL,

MB BChir, FRCR

1

Freeman Hospital, Newcastle upon Tyne and 2Newcastle General Hospital, Newcastle upon Tyne, UK Received 21 July 2005 Accepted 11 October 2005 DOI: 10.1259/bjr/62312165 ’ 2006 The British Institute of Radiology

A 14-year girl presented with progressive painless visual loss in her right eye. She was otherwise well and on examination abnormal findings were localized to the right globe. Further questioning revealed that her mother had undergone surgery for a brain tumour several years previously. MRI of the brain and orbits pre- and postadministration of gadolinium was performed (Figure 1a–d). What are the MRI findings? What is the most likely diagnosis?

(a)

As a result of the cranial MRI findings, imaging of the abdomen was also performed (Figure 2). What abnormality is shown? Does this help to confirm or disprove the diagnosis? What other abnormalities would you wish to exclude in this case? Address correspondence to: Dr Samad Punekar, Radiology, Freeman Hospital, Newcastle Upon Tyne, High Heaton NE7 7DN, UK. E-mail: [email protected].

(b)

(c)

Figure 1. (a) Cranial axial proton density-weighted image. (b) Axial post-

(d)

gadolinium T1 weighted image. (c) Axial T2 weighted image. (d) Coronal postgadolinium T1 weighted image.


931

S Punekar, D J A Butteriss and D Birchall

fluid is usually slightly hyperdense compared with cerebrospinal fluid (CSF). Simple cysts of multiple intra-abdominal organs including liver, pancreas, kidney and adrenal are common, with renal cell carcinoma (RCC) and phaeochromocytoma the most common nonCNS neoplastic abnormalities. vHL-associated RCCs act differently from sporadic RCCs and may be multiple, bilateral and cystic. Phaeochromocytomas in vHL occur in certain lineages only, and this has been used to classify the condition: Diagnostic criteria are summarized below:

Figure 2. Abdominal axial T2 weighted image.

Discussion Figure 1 shows the presence of two synchronous central nervous system (CNS) lesions. First, there is a proliferative process affecting the right retina, which fills most of the vitreous (Figure 1a). This demonstrates enhancement following gadolinium administration (Figure 1b). Second, there is a cystic abnormality in the right cerebellar hemisphere (Figure 1c), with an associated enhancing mural nodule laterally (Figures 1b,d). MRI of the abdomen showed several pancreatic cysts (Figure 2). These appearances are in keeping with the presence of retinal and cerebellar haemangioblastomas, features that in combination with a positive family history allow the diagnosis of von Hippel-Lindau disease (vHL). Pancreatic cysts are found in approximately 75% of patients with vHL. vHL is an inherited multisystem disorder characterized by a variety of highly vascularized tumours of the CNS and the viscera, and is an eponym that carries the names of two eminent European physicians [1, 2]. Eugen von Hippel, a late 19th century German ophthalmologist, described familial retinal capillary angiomatosis in 1895, a lesion that has subsequently been coined ‘‘the von Hippel tumour’’ [3]. In 1926, Arvid Lindau, a Swedish neurologist, reported the occurrence of hereditary cerebellar haemangioblastoma (later termed ‘‘the Lindau tumour’’) and, having noted an association with retinal angiomatosis and renal cell carcinoma, linked the cerebellar, retinal and visceral components into a single coherent entity [4]. He corresponded with von Hippel, and their joint observations were subsequently widely disseminated, largely through the work of Harvey Cushing. The term ‘‘von Hippel-Lindau disease’’ was first used in 1936 by Davison et al [5]. The most characteristic lesion of vHL is the CNS haemangioblastoma, and these are the cause of the majority of morbidity and mortality associated with the condition. They are most common in the cerebellum, but may be found anywhere in the neuraxis from the cerebral hemispheres and optic nerves to the spinal cord. The typical appearance is of a cystic lesion with an avidly enhancing mural nodule, but there is a spectrum of appearances to a completely solid lesion. The cyst

932

1. CNS and retinal haemangioblastoma. 2. Haemangioblastoma and at least one of the following: renal, pancreatic, hepatic, epididymal cyst. Phaeochromocytoma, renal carcinoma. 3. Family history and at least one of the following: haemangioblastoma, visceral changes, phaeochromocytoma, renal cancer. This disorder has an incidence of approximately 1 in 36 000 live births. It is a highly penetrant autosomal dominant trait, and has been attributed to a germline mutation of a tumour suppressor gene on the short arm of chromosome 3. Active research continues into the molecular consequences of the vHL gene. Screening of affected patients and at-risk relatives is necessary because of the high incidence of RCC (occurring in up to 45% of cases) and because of the morbidity associated with CNS lesions. The Cambridge protocol was devised by Maher et al for screening patients with vHL disease or at-risk relatives [6]. Affected asymptomatic patients and at-risk relatives should have annual physical examination, ophthalmoscopy and urine testing, brain imaging every 3 years, and abdominal CT scanning every 3 years. Treatment involves resection of the offending tumour, aspiration of the cysts causing pressure-related symptoms and cryotherapy or phototherapy of retinal lesions. Radiology has a central role in managing vHL. Because a conservative approach to the treatment of some vHL lesions is now more widely accepted, ongoing follow-up with careful cross-sectional imaging plays a central role in evaluating the progression of disease.

References 1. Richard S, Graff J, Lindau J, Resche F. Von Hippel-Lindau disease. Lancet 2004;363:1231–4. 2. Lonser RR, Glenn G, Walther M, et al. Von Hippel-Lindau disease. Lancet 2003;361:2059–67. 3. Von Hippel E. Vorstellung eines Patienten mit einer sehr ungewohnlichen Netzhaut. XXIV Verstellung der ophthalmolgischen Gesellschaft (Heidelberg, 1895). Wisebaden, Germany: Bergmann Verlag, 1896:269. 4. Lindau A. Studien eber Kleinhirnzysten. Bau, Pathogenese und Beziehungen zur Angiomatosis retinae. Acta Pathol Microbiol Scand 1926;S1:1–128. 5. Davison C, Brock S, Dyke CG. Retinal and central nervous hemangioblastomatosis with visceral changes (von Hippel-Lindau’s disease). Bull Neurol Instit NY 1936;5: 72–93. 6. Maher ER, et al. Clinical features and natural history of von Hippel-Lindau disease. Quart J Med 1990;77:1151–63.



Book reviews Basics of PET imaging: Physics, Chemistry and Regulations. By Gopal B Saha. pp. xv+206 (Springer, NY) £35.00 ISBN 0-387-21307-4 Gopal Saha is a World expert in radiopharmacy. His text book on Fundamentals of Nuclear Pharmacy is now in its fifth edition. Here, he has produced a very readable text covering all aspects of the basics of PET imaging. The book is divided into 11 chapters covering radioactive decay and interaction of radiation with matter; PET scanning systems; data acquisition and corrections; image reconstruction, storage and display; performance characteristics of PET scanners; cyclotron and production of PET radionuclides; synthesis of PET radiopharmaceuticals; regulations governing PET radiopharmaceuticals; reimbursement for PET procedures; design and cost of PET centre; and procedures for PET studies;. There are several appendices, including a useful glossary of terms and answers to the numerical and yes/no questions at the end of each chapter. There is a also good index. The lay-out of the book is logical and, for anyone with some relevant background knowledge either in radiation physics or nuclear medicine, the text is easy to follow. Some features I particularly liked were a good, brief description of factors affecting acquired data –normalisation, photon attenuation, random coincidences, scatter coincidences, dead time and radial elongation; a good discussion of the performance characteristics of PET scanners – spatial resolution, sensitivity, noise equivalent count rate, scatter fraction and contrast. There is a very clear description of the various factors that contribute to limiting spatial resolution, including positron range and non-collinearity of the coincidence gammas; also, I enjoyed the useful list, with brief details, of the physicochemical and biological tests that are necessary as part of a quality assurance programme for PET pharmaceuticals. The text is generally accurate and has been carefully proofread, although one or two numerical errors seem to have crept in, especially in Table 1.3 – half value layers for 511 keV photons. Also, the quoted effective dose (assumed to be in mSv MBq21) for an 18F-FDG examination cannot be correct if the stated organ doses are right (p. 193). UK readers may have a few problems with this book, written for the US market, including use of units – although the SI equivalent of the units still currently used in the US (curies, rads etc.) is frequently given, a reader who cannot readily switch from one system to the other may have difficulty. Also, radiation protection is a major problem. In a recent review of PET: molecular imaging and its biological applications, edited by ME Phelps; see Br J Radiol 78:2005;871, Lesley Malone commented that radiation safety had been neglected. It is covered at length by Saha in the longest chapter in the book. However, organisational details, with the involvement of both Federal and State Regulatory Bodies, are completely different in the US from those in the UK. Also, even the regulations themselves show variations, e.g. some differences in dose limits and descriptors of designated areas. Since the section on principles of radiation protection is not specific to PET and The British Journal of Radiology, November 2006

the same information can be found in a number of other texts, there is a case for skipping this chapter completely. The subject of patient doses appears to receive rather less attention in the US than in the UK and Europe. By combining the information in Appendix D with that in Chapter 11, the effective dose for a few clinical PET procedures can be worked out. However, there is no attempt to put these into context with other nuclear medicine and radiological examinations. In conclusion, who should read this book? The preface suggests that it is suitable for nuclear medicine technologists, radiographers, radiation therapy technologists and physicians studying PET courses. This is unlikely to be the situation in the UK because PET technology has not diffused so rapidly and most persons in the abovementioned groups would find that there were just too many unfamiliar concepts. However, the book can be recommended to experienced nuclear medicine staff and more junior medical physics staff who wish to get involved in PET or simply to find out more about it.

P DENDY Imaging of the head and neck, 2nd edn. By Mahmood F Mafee, Galdino E Valvassori and Minerva Becker. pp. 866, 2004 (Thieme Medical Publishers Stuttgart, Germany) J229.95 ISBN 3 13 100942 X This is the second edition of this superb reference text which is, against my expectations, improved over the original first edition. The book is quite capable of causing significant damage if you try and read it on your knee, as I did, with over 800 pages of heavily and well-illustrated clinical material, presented in excellent and logical order. The information in the book is presented in six anatomically defined sections, starting with the temporal bone, the eye, orbit and base of skull, nasal cavity and paranasal sinuses, the masticatory system, the suprahyoid neck, and the infrahyoid neck. In each of these there is excellent illustrated, descriptive anatomy of both the pure anatomy and the radiological important components. No time or effort has been spared to ensure that the reader is provided with any audio visual aid in terms of diagrams, 3D renderings, histological sections or images that could possibly help their comprehension of the underlying anatomical components. The anatomical descriptions are then followed by detailed discussion of the relevant pathologies affecting the area, lavishly illustrated with an absolutely excellent series of cases, the majority of which are collected on recent generation imaging equipment and are of very high quality. The discussion and presentation is of more than adequate detail for the specialist radiologist interested in these areas, whilst at the same time, on the whole, being easily so well presented that it could be perused by the non-specialist radiologist or the radiologist in training without them feeling particularly threatened. Overall, I have to admit that I am fairly jealous that I did not write this book and it certainly has a strong place on my desk top from now on. A JACKSON 933


September 2006 Volume 79 Special Issue

The British Journal of Radiology, 79 (2006), S1

PREFACE

Imaging in radiotherapy treatment planning and delivery R M HARRISON Deputy Editor (Physics & Technology)

DOI: 10.1259/bjr/14035252 ’ 2006 The British Institute of Radiology

Recent years have seen major developments in radiotherapy techniques and equipment, focusing on conformal external beam techniques with the emphasis on intensity-modulated radiotherapy (IMRT). The progress of IMRT implementation has been summarized in a series of review articles in the British Journal of Radiology during 2003 and 2004. One of the fundamental prerequisites for conformal radiotherapy is the localization of the target, starting with the gross tumour volume (GTV) and moving outwards to the planning target volume (PTV). The judgement of how to define these volumes is crucial. The future value of IMRT and related techniques may depend on the underpinning images of these volumes and how they are interpreted. Fortunately, in parallel with the developments in radiotherapy delivery techniques has come an equally impressive and timely development in associated imaging. Although CT remains the gold standard for radiotherapy planning, MRI has much to offer. Superior soft tissue characterization and new developments in dynamic contrast enhanced, diffusion-weighted and diffusion tensor imaging are making their mark in target delineation. Magnetic resonance spectroscopy (MRS) promises insights into tissue metabolism and, although currently limited by large voxel sizes, may provide information on tumour response and recurrence to complement the structural information of MRI and CT. The traditional role of the simulator is now complemented by the use of the CT scanner as a virtual simulator, thus making 3D image datasets available.

The British Journal of Radiology, Special Issue 2006

Verification on a fraction-by-fraction basis is now possible by several means. CT scanners within the treatment room have been used to compare images at treatment time with planning CT scans and also to measure and correct for interfraction variations in anatomy or positioning. Portal imaging has seen equally impressive developments as a result of the introduction of flat panel imaging systems. As well as their improved image quality for traditional radiographic projection imaging, when used in conjunction with cone beam reconstruction algorithms, they have made megavoltage CT a reality. Alternatively, if used with a kilovoltage X-ray source and cone beam CT algorithm, the CT reconstructions can harness the improved tissue contrast associated with the lower X-ray energies. The development of PET/CT imaging for radiotherapy planning provides, like MRS, another example of the imaging of metabolic function, but in this case the functional information is displayed within the same coordinate system as the familiar structure provided by CT. This is an unusual luxury, however, and the problems of extracting and overlaying data from different imaging studies and modalities has prompted several advances in image registration and fusion algorithms. So these are exciting times in radiotherapy imaging and this special issue reflects both the current practice and future opportunities for exploiting advances in several imaging modalities in order to increase the confidence with which modern conformal radiotherapy treatments may be planned and delivered.

S1

The British Journal of Radiology, 79 (2006), S2–S15

New developments in MRI for target volume delineation in radiotherapy 1,2

V S KHOO,

FRACR, FRCR, MD

and 3D L JOON,

FRACR

1

Royal Marsden Hospital, Institute of Cancer Research, Fulham Road, London SW3 6JJ, 2University of Manchester, Manchester, UK and 3Austin Health Radiation Oncology Centre, Heidelberg Repatriation Hospital, Victoria, Australia ABSTRACT. MRI is being increasingly used in oncology for staging, assessing tumour response and also for treatment planning in radiotherapy. Both conformal and intensity-modulated radiotherapy requires improved means of defining target volumes for treatment planning in order to achieve its intended benefits. MRI can add to the radiotherapy treatment planning (RTP) process by providing excellent and improved characterization of soft tissues compared with CT. Together with its multiplanar capability and increased imaging functionality, these advantages for target volume delineation outweigh its drawbacks of lacking electron density information and potential image distortion. Efficient MR distortion assessment and correction algorithms together with image co-registration and fusion programs can overcome these limitations and permit its use for RTP. MRI developments using new contrast media, such as ultrasmall superparamagnetic iron oxide particles for abnormal lymph node identification, techniques such as dynamic contrast enhanced MRI and diffusion MRI to better characterize tissue and tumour regions as well as ultrafast volumetric or cine MR sequences to define temporal patterns of target and organ at risk deformity and variations in spatial location have all increased the scope and utility of MRI for RTP. Information from these MR developments may permit treatment individualization, strategies of dose escalation and image-guided radiotherapy. These developments will be reviewed to assess their current and potential use for RTP and precision high dose radiotherapy.

The increased sophistication of modern radiotherapy planning techniques such as conformal (CFRT) and intensity-modulated radiotherapy (IMRT) necessitates improved means of defining target volumes for treatment. This is needed to achieve the intended benefits of using CFRT and IMRT. This step remains the most crucial and difficult part of the radiotherapy planning process, otherwise a geographical miss of the tumour or a systematic error will be perpetuated throughout therapy. MRI is being increasingly used in oncology for staging, assessing tumour response and evaluating disease recurrence. As a result of the enhanced imaging properties of MR, it has been estimated to be a more cost effective diagnostic tool in the management of some diseases [1]. Similarly, the improved characterization of soft tissues and visualization of tumour extent using MRI can be used to benefit the radiotherapy treatment planning (RTP) process from delineation of target volumes to determining planning margins and treatment response [2]. There are many current areas of development in MRI. These include developments in hardware technology, such as 3 Tesla machines, and the use of new MR contrast media, such as ultrasmall superparamagnetic iron oxide particles for lymph node evaluation [3]. MR techniques and sequences previously used for research are now becoming available for general use. MR techniques such as dynamic contrast enhanced and S2

Received 9 August 2005 Revised 6 January 2006 Accepted 10 March 2006 DOI: 10.1259/bjr/41321492 ’ 2006 The British Institute of Radiology

diffusion weighted MRI may provide further characterization of tissue and tumour regions [4, 5]. MR sequences such as ultrafast volumetric and 3D cine sequences can offer the opportunity to assess target/organ motion and deformity [6, 7]. Temporal-spatial information gleaned from MRI can then be used for image-guided strategies in radiotherapy delivery. All these features have the potential to increase the scope and utility of MRI for RTP. It is worthwhile briefly reviewing the background to the use of MRI for RTP in order to understand the rationale and issues with its use. Some methods of utilizing MRI in RTP will be outlined. This article will then discuss the new MRI developments in terms of their current and potential impact in target volume definition for treatment planning with examples of applications at some cancer subsites.

MRI rationale for RTP Any additional procedures used for RTP must add value to the planning process. Standard RTP uses CT data. CT images are good at distinguishing between structures that have substantially different X-ray attenuation properties or Hounsfield units, such as between air, tissue and bone. It is more difficult to discriminate between adjoining soft tissue structures using CT if these soft tissue structures possess similar The British Journal of Radiology, Special Issue 2006

MRI for target volume delineation in radiotherapy

Hounsfield units unless there is a fat, air or bone interface between these structures. The imaging parameters for CT scanning are much more limited compared with the range available with MRI. In the case of MRI, the contrast from soft tissue structures can be widely varied by extensively manipulating the imaging parameters, which include proton densities and tissue relaxation times (spin-lattice or T1 and spin–spin or T2). This increased flexibility in varying tissue contrast or signal intensities offers much better characterization of soft tissues even when these structures possess very similar X-ray attenuation properties or electron densities. Tumours often have similar electron densities to their neighbouring soft tissues. By using different MRI sequences, better tissue discrimination can be obtained between the extent of tumour with its boundaries of infiltration and the adjacent normal structures. In this manner, MRI provides improved target delineation for RTP. This utility of MRI applies not only to the initial radiotherapy treatment of tumours but also potentially for re-treatments by being able to differentiate between changes due to recurrent cancer or that secondary to post-treatment fibrosis. It can also provide better delineation of organs at risk (OARs) for dose avoidance in RTP. An obvious benefit of enhancing the visualization of volumes of interest (VOI) is the increased reliability and consistency of target definition. This will improve both interobserver and intraobserver variability for outlining. This has value for institutional and multicentre trials in radiotherapy where it is important to maintain consistent and accurate target and OAR volumes. Substantial and inappropriate variations in target volumes can impact on trial outcomes, with geographical misses for poorer local control rates or unnecessary inclusion of normal tissue for higher toxicity rates. MRI can avoid bony and metal artefacts seen with CT. Large thick bony sections attenuate X-rays and reduce the adjacent soft tissue image quality. This can obscure identification of nearby tumours and internal anatomy.

MR images are not affected by this. MRI can thus further improve the delineation of both tumour and OAR volumes for RTP in these regions. Another feature of the increased functionality of MRI is its true multiplanar capability. This ability to image in any oblique plane can reduce the ‘‘partial volume’’ imaging effect that often results from conventional transaxial CT imaging, particularly where the 3D shape of the target is extreme or changes substantially between conventional transaxial CT slices. Para-sagittal or paracoronal views can also permit better understanding of the boundaries of target volumes with the surrounding normal tissues leading to better target volume delineation (Figures 1 and 2). Furthermore, MRI can provide functional and biological information for tumour regions that may improve target definition and permit new opportunities for novel radiotherapy strategies. Some of the salient features of using MRI for RTP are summarized in Table 1. However, it is important to be aware that implicit in the use of MRI for RTP is that oncology clinicians should have the necessary training to comprehend MR images and understand how to use them appropriately for defining VOIs. Thus it is important to undertake suitable supervised training [8]. Even if there is relevant experience, it is still beneficial to liaise closely with local diagnostic colleagues who have MR expertise for the cancer subsite(s) in question. Ideally there should be an oncology team collaboration for the definition of target volumes in radiotherapy similar to the multidisciplinary team arrangement that exists for general cancer management. This approach has been endorsed by national bodies, such as the Royal College of Radiologists in a recent publication [9]. All of these advantages add to the RTP process and outweigh its current drawbacks, which include the lack of electron density information, potential image distortion and specific patient considerations with MR scanners. Some examples of the current impact of using MRI for target volume delineation and some of these

(a)

(b)

Figure 1. A comparison of sagittal views of the pelvis for prostate radiotherapy with (a) CT reconstructed from 2.5 mm slices and (b) MR image obtained in-plane in the same patient. Some of the relevant structures of interest for radiotherapy are labelled on the MR image. These structures are not visualized well enough on CT to provide confident determination of the prostate boundaries for radiotherapy. The British Journal of Radiology, Special Issue 2006

S3

V S Khoo and D L Joon

(a)

(b)

Figure 2. A comparison of coronal views of the pelvis for prostate radiotherapy with (a) CT reconstructed from 2.5 mm slices and (b) MR image obtained in-plane in the same patient. Definition of the prostate gland boundaries and the adjacent structures is better visualized on MRI than with CT.

Table 1. Advantages and disadvantages of MRI for radiotherapy planning (RTP) Features

Advantages

Patient

Non or minimally invasive procedure Claustrophobia due to the smaller patient bore Few patient risks Contraindicated in patients with loose metal No radiation associated with imaging foreign bodies within the body, particularly This may be advantageous to paediatric patients and pregnant the orbits or pacemakers women This may be a useful for follow-up scanning Increased number of imaging parameters for more MR image distortion imaging flexibility Systems Superior soft tissue imaging with excellent spatial resolution Object induced distortions to provide better visualization for the following: determining Lack of electron density information for dosithe tumour/GTV extent and degree of tumour infiltration metry and needs additional steps to permit Understanding the surgical bed or altered anatomy secondary dose calculations to surgery Lack of cortical bone information to create Distinguishing between post-treatment fibrosis or tumour digitally reconstructed radiographs (DRR) in recurrence radiotherapy Improved definition of normal soft tissue structures and May have longer scan times than CT with more tissue planes avoidance of image artefact from metal potential for motion artefacts prosthesis and large bony regions Need for specific training to comprehend and True multiplanar capability to image in any oblique plane and understand MR images for RTP use reduction of the ‘‘partial volume’’ imaging effect increased RTP systems can only import transverse MR accuracy, reliability and consistency of target definition to images and cannot take full advantage of reduce both interobserver and intraobserver variability sagittal and coronal in-plane MR images Providing functional and biological information for Immobilization devices used in radiotherapy functional avoidance or biological targeting may not be MR compatible Ultra-fast volumetric and cine mode acquisitions to assess temporal-spatial variations in target positioning or deformation Can be registered with CT information for use in RTP systems New contrast agents (i.e. USPIO) to define nodal status Less incidence of allergic reactions to gadolinium than iodine-based contrast agents New bore flange openings to reduce patient claustrophobia Not as readily available and accessible as CT Open MR systems for easier patient access, tolerance Smaller bore than CT (52 cm vs 82–85 cm) and positioning for radiotherapy Curved table top

Imaging

Contrast agents Machine

Disadvantages

GTV, gross tumour volume; USPIO, ultrasmall paramagnetic iron oxide.

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issues in using MRI for RTP will be briefly outlined below. A more in-depth review has been published and the reader is advised to peruse this [10].

Current impact of MRI for treatment planning In many areas of oncology, diagnostic MRI is the gold standard imaging modality for the staging and assessment of cancer patients. This is true for imaging of cerebral and spinal tissues, soft tissue sarcomas and pelvic tumours. Some examples will be highlighted here to illustrate the current status of the added advantage of MRI for target volume delineation in the planning process. MRI has been used extensively for central nervous system (CNS) radiotherapy. Many investigators have reported quantitative improvements of up to 80% of cases in target volume definition with the addition of MRI to 3D CT based treatment planning [11–13]. Despite the obvious advantage of MRI for target volume definition at this site, there are clear situations where the use of both CT and MRI data is valuable and can provide for more consistent volume delineation than compared with either MRI or CT alone [14]. In a study of base of skull meningiomas, MRI was able to delineate tumour volumes that were present close to the base of skull bones as the X-ray attenuation from these large bones can obscure soft tissue detail using CT alone [15]. However, CT was able to provide information on the extent of bony erosion from tumour that was not available with MRI. In some cases, the individual CT or MR volumes were vastly different, with each modality providing separate but complementary information (Figure 3). This supports the use of combined CT and MRI data to provide the optimum target volume delineation. Segmentation algorithms, automated and atlas based, may further aid the delineation process and this methodology is currently being investigated [16, 17]. In many centres providing CFRT and IMRT for CNS tumours, the use of CT-MRI fusion for RTP may now be considered as standard practice. The major thrust of current developments in CNS treatment planning is to integrate the use of functional data for target volume determination [18]. Head and neck anatomy is complex and the extent of the infiltrating tumour can be difficult to define. MRI can assist in delineation of radiotherapy volumes here. It is useful for defining (1) longitudinal tumour infiltration along the upper aero-digestive tract and adjacent fascial planes, e.g. pre-vertebral fascia, (2) tumour infiltration of soft tissue structures and tissue planes such as the pterygoids and tongue (Figure 4), (3) the extent of perineural infiltration and intracranial extension, e.g. nasopharyngeal tumours and (4) nodal metastases. This situation is best illustrated by nasopharyngeal tumours where the use of multimodality imaging with MRI can change disease staging in about 50% of cases impacting on RTP [19, 20]. In a study of over 250 patients, up to 40% of intracranial infiltration detected on MRI was missed by CT [21]. Segmentation algorithms for MRI are being developed to help in target volume delineation based on the contrast enhancement ratio of T1 weighted images and signal intensity of T2 weighted images, and The British Journal of Radiology, Special Issue 2006

Figure 3. A case from a study of meningiomas of the skull base evaluating the differences between MRI and CT assessment of the clinical target volume (CTV) for radiotherapy. The 3D reconstructed view of the CT-defined CTV (red outlines) and MR-defined CTV (yellow outlines) illustrates the spatial differences in CTV definition by the two different imaging modalities where the MR-defined CTV demonstrates tumour extending laterally along the petrous ridge that was not seen using CT [15].

this methodology may further assist the RTP process [22]. In the pelvis, MRI has provided improved target delineation for urological, gynaecological and gastrointestinal cancers. For prostate cancer, MRI can provide better internal organ assessment than CT for disease extent, capsular and seminal vesicle involvement [23–25]. MRI can overcome some of the limitations of CT definition of prostate treatment volumes (Figures 1 and 2). MRI, particularly using sagittal views, can be useful in defining the prostatic apex and distinguishing between the boundaries of the prostate with the base of the bladder and the anterior wall of the rectum [26, 27]. The prostatic capsule, which cannot be distinguished from adjacent normal tissue on CT, can be seen as a thin rim of low signal intensity on T2 weighted MRI thereby permitting boundary definition (Figure 5). Comparative MRI-CT planning studies using MRI-defined prostate volumes as the gold standard have reported that CTdefined prostate volumes tend to overestimate the planning volume by as much as 27–43% due to the soft tissue uncertainty in CT delineation [28–31]. In our prospective study of 105 men with prostate cancer, we found that MRI-CT fusion for RTP can often up-stage the disease extent by clarifying extracapsular spread, seminal vesicle involvement and early adjacent organ invasion that resulted in substantial changes to the target S5


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Figure 4. A case of a base of tongue cancer imaged with (a) CT and (b) MRI showing a large mass on the left side of the oropharynx involving the base of tongue and the left tonsillar fossa with invasion of the left parapharyngeal space. The base of tongue mass extends past the midline. These features are better visualized using MRI than CT.

volumes in 29% of cases [32]. In a retrospective review of 199 patients treated with radiotherapy, a comparable upstaging effect of MRI was seen in 52% [33]. For postprostatectomy patients, a visible mass was noted within the operative bed using MRI in 50% of cases, that was not seen using CT [32]. This improved ability to delineate prostate and seminal vesicles can also reduce interobserver and intraobserver variation. MRI can also be

useful where the internal pelvic anatomy has been substantially altered due to previous extensive surgery such as abdominal-perineal resections [34]. It can also aid delineation of adjacent normal tissue structures such as rectal wall, recto-vesicle fascia of Denonvillier, urogenital diaphragm, penile bulb, periprostatic venous plexus, neurovascular bundle, levator ani and anal sphincters. The use of MR-based prostate planning volumes can

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Figure 5. A comparison of transaxial co-registered views of the pelvis for prostate radiotherapy with (a) CT and (b) MR in the same patient. The boundaries of the prostate are better visualized using MR than with CT, notably the anterior rectal wall/rectovesical fasia and prostate capsule. S6



result in more appropriate treatment volumes, leading to better shaping of the treatment fields that can reduce the risk of treatment related complications to important normal structures such as the rectum and penile bulb [35]. In addition, co-registered CT/MRI has also been used for permanent and high dose rate prostate brachytherapy to better define target volumes, reduce delineation uncertainty for RTP and to guide needle insertion compared with other imaging methods used such as ultrasound or CT [36, 37]. This is particularly important for post-implant dosimetry where the prostate gland delineation may be more difficult following the implant procedure and where CT assessments may be hampered by its poor intrinsic tissue contrast and seed induced artefact [38, 39]. MR-CT fusion has also been found to be beneficial for brachytherapy in other tumour sites such as head and neck, sarcomas and gynaecology [40]. In cervical brachytherapy, the use of MRI has been strongly supported by the Gynaecological GEC-ESTRO Working Group [41]. For rectal cancer radiotherapy, traditional planning relies on visualizing a filling defect using rectal contrast with the treatment fields usually placed according to bony landmarks, but rectal contrast does not define the circumferential thickness of the tumour and bony landmarks do not accurately define the anterior margin as governed by the rectal lymphatic drainage and mesorectum. CT planning can improve delineation of the rectal tumour by providing visualization of the increased thickness of the rectal wall and provide better definition of the lymphatic region by depicting the vascular structures, visible lymph nodes and boundaries of the mesorectum [42]. However, inaccuracies in the CT definition of tumour may occur because of poor contrast between faeces and tumour, partial volume effects due to the curves/valves of Houston in the rectum and imaging of the horizontal sigmoid. Infiltration into the anus can

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also be difficult to assess in a low rectal tumour unless there is an obvious mass effect (Figure 6). MRI can avoid these CT identification issues and better define the depth of invasion through the rectal wall [43]. MRI can provide visualization similar to endorectal ultrasound investigations, but ultrasound data cannot be imported into treatment planning systems. MRI can aid the CT based RTP process by defining the longitudinal spread of the cancer superiorly and inferiorly, as well as the extent of infiltration of the mesorectum [44]. It may also provide better assessment of early invasion into local adjacent structures such the bladder, prostate or seminal vesicles in men, and vagina and uterus in women, in addition to anal sphincter infiltration in both [45–47]. This better estimation of the gross tumour volume may permit strategies of anal sphincter sparing, tumour boosting and dose escalation with or without concurrent chemotherapy.

MRI issues and schemas for treatment planning A comprehensive step-by-step guide and methodology review for the integration of MR into treatment planning systems is outside the remit of this article. However, it is worthwhile to outline briefly some of the relevant issues faced when using MR images for RTP and some of the available methods used. In RTP, geometrically accurate images are needed for precision radiotherapy and electron density information is required to take into account tissue inhomogeneities when calculating dose distributions. CT data provide the necessary information for both these requirements. MRI suffers from a lack of electron density information and potential spatial image distortion. As a result, MR images cannot be imported alone into RTP systems to readily create and plan three-dimensional (3D) radiotherapy.

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Figure 6. Co-registered (a) CT and (b) MRI scans showing a rectal cancer in the lower rectum extending to the anorectal junction with invasion of the left posterolateral wall. This is easily seen on the T2 weighted MRI scan as a hyperintense signal region that is not visible on CT. The British Journal of Radiology, Special Issue 2006

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Unlike CT images where the Hounsfield units can be directly related to electron density values, the signal intensities in MR images do not have such a relationship. If MR images are to be used alone for RTP, then X-ray attenuation coefficients have to be manually assigned to all regions within the MR image in order to allow dosimetry calculations. In some regions, such as in the brain where the internal tissue is more homogeneous and the skull can be easily segmented, this may be easily achieved by assuming an homogeneous attenuation value within the skull vault. On MR images, investigators have calculated that doses were within 2% of those calculated using CT data [48]. Similar work has been recently performed in the pelvis for prostate radiotherapy with 2–3% between CT-based and MR-based dose calculations [49]. Other investigators noted dosimetry discrepancies that were .2% compared with CT dosimetry for the pelvis if bone and water CT number bulk-assigned values were not assigned to MR images [50]. Simple methods of CT density value assignments may more difficult to implement if the anatomy is varied or complex such as in the head and neck regions or thorax/abdominal regions. An alternative method is to integrate MR images with CT data. This requires bringing the MR data into spatial 3D alignment with planning CT data or image registration. Then the different image modalities can be ‘‘fused’’ so that a common image set is produced with both sets of imaging information available for clinical interrogation. This method is favoured as the data can be transferred and implemented within a treatment planning system and provides for optimum evaluation of the imaging information to create the most appropriate VOI for RTP. The methodology for image co-registration and image fusion is considered by Kessler in this issue.

MR image distortion Before image co-registration can occur, it is important to ensure that the MR data are suitable. MR image distortion is one potential concern. MR image distortions can be grouped into two main categories; system-related and object-induced. An illustration of these distortion effects is shown in Figure 7. System-related distortions are due to the imperfections of the magnet, its operating system and imaging sequences. The effects of magnet field inhomogeneity, gradient field non-linearities and eddy currents can contribute to system-related image distortion. In general, the effects of system related distortion are smallest at the centre of the magnet and worsen with increasing distance from the magnet centre. The distortion magnitude is largest at the periphery of the field-of-view (FOV). Object-induced distortion arises when any object (i.e. the patient) is placed within a magnetic field. This type of distortion results from magnetic susceptibility and chemical shift effects. Different body tissues have different magnetic susceptibilities and susceptibility artefacts can be pronounced at different tissue boundaries, such as between air cavities and soft tissues. Chemical shift effects result from the different behaviour of protons in fat and tissue. Fat protons ‘‘precess’’ at a S8

Figure 7. An illustration of various forms of distortion in MRI using a phantom consisting of a coplanar array of waterfilled tubes embedded within in a circular solid plastic (PMMA) block. System distortion effects are seen in the apparent curvature of the tubes at A and their disappearance at B, which was due to warping distortion of the imaging plane. Magnetic susceptibility differences due to the presence of the plastic support block at C give rise to object-induced distortions in the form of discontinuities at the point where each tube enters the support block [10]. (Reproduced with permission from Elsevier).

slower rate than water protons and this can result in a chemical shift effect where the positions of fat/water protons are shifted from their true spatial locations. In order to utilize MR images for RTP, these image distortions must be evaluated, minimized, and/or corrected especially if they are to be co-registered with CT, otherwise a systematic error will be incorporated into the treatment plan. There are many methods to deal with MR image distortion, but the primary task is to quantify the presence and extent of any distortion. This methodology has been previously described in detail [51–54]. This system provides for evaluation and correction of both system-related and object-induced distortions, but can also facilitate quality assurance programs. It is important to ensure that the same MR scanner and imaging sequence used for mapping is also used for patient imaging. System-related distortion is best quantified and mapped using a phantom (linearity test object) with a known array of markers in 3D to provide spatial assessments of the whole imaging volume used for RTP. A reference frame with imbedded markers provides another set of marker positions that covers the periphery of the imaging volume. The linearity test object or the patient is housed within this reference frame. A separate set of markers placed on the patient’s surface provides assessment of object-induced effects. In brief, the positions of all markers are mapped within the imaging volume using a dedicated automated algorithm. The use of read-out gradient reversal imaging and postprocessing image corrections can account for objectinduced effects. Distortional shifts of up to 5 mm can be corrected [51–54]. Other correction methods have also been developed and can reduce distortions by a factor of two [55]. Site specific phantoms have also been created The British Journal of Radiology, Special Issue 2006


with air cavities to assess the use of MRI for lung RTP [56].

Other MR considerations It is important to review the image protocols to be used and to evaluate those sequences which can offer the best combination of image quality/resolution and minimal image distortion. These sequences will differ depending on the anatomical site being imaged and treated. As previously mentioned, the effect of distortion is least within the centre of the magnet and VOI for RTP should be imaged within this zone whenever possible. If larger FOVs are used, the central region of imaging may be extracted and this region can then be co-registered with CT data to aid target volume definition. In determining imaging protocols for multi-modality image co-registration, attention must be paid to some of the issues listed below to optimize the appropriate imaging acquisition parameters. 1. What is the appropriate sequence to be used for the cancer type and anatomical region? For example, will it be T1, T2 weighted or a hybrid sequence, standard or ultra-fast acquisition, gated or volumetric, and contrast or not? This will depend on what imaging information is needed by the clinician. 2. What is the optimal FOV and relevant imaging volume? 3. What is the resolution needed? 4. What is the appropriate imaging slice orientation, slice thickness and slice gaps? 5. Is there a need for multislice and/or non-coplanar image reconstruction or oblique plane imaging? 6. What is the influence of various body coils or internal body MR probes on the MR images for co-registration? 7. What is the appropriate quality assurance program to ensure reliability of image quality and data transfer? Any MRI for RTP should also mimic the CT planning procedures such as scanning the patient in the treatment position with a flat bed insert, using the same but MR compatible immobilization devices where specified, providing the same instructions to the patient, e.g. full or empty bladder, and minimizing internal organ motion by breath held procedures, bowel relaxants or reducing the scanning time whenever indicated and possible. Ideally the MRI scans should be timed as close as possible to the CT planning scans.

New developments in MRI Developments in MR hardware 3 T MRI scanners The field strengths of MRI scanners have been increasing since their initial development from 0.5 T (Tesla) scanners to 1 T and now 1.5 T and 3 T. The advantages of increasing the field strength are severalfold. The quality of the MR image is related to the signal The British Journal of Radiology, Special Issue 2006

obtained relative to any imaging deficiencies using the MR scanner, such as machine imperfections and object/ patient induced errors. This relationship is termed the signal-to-noise ratio (SNR). The image quality or resolution increases when the signal is strengthened and the noise lowered. The SNR approximately increases in a linear manner with field strength. In simplistic terms, higher field strengths would permit higher resolution images with improved tissue contrast that can lead to better tissue diagnosis and definition of tumour volumes for RTP. The image acquisition time may also be reduced and this can potentially minimize motion artefacts. With these advantages, the development and use of 3 T MRI machines may become more commonplace as they can further improve the image quality using external phased array coils. In prostate MRI, the image quality from external phased array coils using 3 T are comparable with endorectal coils at 1.5 T [57] and in some cases exceeding them [58]. This provides more imaging options and avoids the internal deformation that occurs with endorectal coils that may limit its use for RTP. These 3 T machines can also benefit the application of functional MRI and MR spectroscopy (MRS) by providing better resolution for assessed metabolites that can define tumour regions, e.g. for brain [59] and prostate radiotherapy [60]. The latter subject is covered in more detail by Payne and Leach in this issue. There are several issues associated with 3 T MR scanners that may limit their use for RTP. Amongst these is the exacerbation of magnetic susceptibility effects, doubling of the chemical shift effect, patient safety and engineering challenges. Currently, the clinical role of these 3 T scanners including their utility for RTP remains to be defined. Magnetic susceptibility effects. Higher field strengths can exacerbate the magnetic susceptibility artefacts, particularly at borders of different structures; for example, at tissue/bone/air boundaries, such as in the head and neck regions [61]. These effects can lead to increased signal intensities that can cause misdiagnosis using contrast studies unless pre- and post-contrast views are used for appropriate comparison. This effect can potentially result in misidentification of abnormal areas leading to erroneous volumes for RTP. Chemical shift effect. The chemical shift effect (see above) is doubled using 3 T field strengths. This effect causes misregistration of fat and water tissues and will hinder the definition of radiotherapy volumes when coregistered with planning CT images. However, this aspect may be of considerable use in MRS as it can increase the resolution of identifying tissue metabolites. This may permit improved and more reliable localization of tumour regions for boosting, such as the use of MRS citrate-choline assessment in prostate cancer. It is also very useful for brain and MR angiography assessments because of the longer, more variable T1 relaxation time. Patient safety. There may be patient safety concerns as the energy deposited using 3 T scanners may be up to 4 times that of 1.5 machines. This may be a factor when using fast or intensive pulse sequences. Potential S9


problems may be reduced by appropriately modifying the pulse sequences and/or reducing the volume of tissue being imaged. Machine factors. 3 T scanners are still not commonplace and are substantially more expensive than 1.5 T machines. There are also engineering issues that include the use of higher gradient systems and increased shielding, as well as smaller scanner bores. These may restrict the size of the patient scanned and also increase the potential for patient claustrophobia.

MR simulators Open and low field MR scanners can act as radiotherapy simulators providing a radiotherapy environment similar to CT simulation for RTP. Other advantages include the lower cost of these machines, greater T1 image contrast, reduced vessel flow/ghosting artefacts and the opportunity to use patient immobilization devices that were previously restricted by size with conventional MR scanners. Although open low field MR scanners may have larger system-related distortions which can be compensated for, object-induced distortions are substantially less due to the lower magnetic field strength [40, 62]. Individual MR sequences may be longer using low field MRI and provide more opportunity for internal organ motion, but appropriate selection of sequences can reduce the scanning required. One limitation of MR simulation is that there is inadequate image detail of bony structures from MR to automatically create digitally reconstructed radiographs needed for treatment verification. Software to delineate bone regions can overcome this issue. The lower SNR may limit diagnostic quality images, but good quality images have been obtained using low field 0.2 T MRI in regions of interest such as the prostate [63]. A recent study of 243 patients revealed that open low field MR simulation provided adequate images for RTP in up to 95% of cases [40]. The greater target volume delineation from MRI led to RTP improvements in up to 33% of lung cases and 40% of prostate cases. MR simulation can better delineate erectile soft tissues to permit dose sparing of these structures by IMRT [64].

Developments in MR sequences Manipulation of the relaxation times of protons in tissues provides the two basic T1 and T2 weighted MR sequences. Whilst this provides superior imaging of soft tissues, the imaging parameters can also be manipulated to benefit RTP by providing ultrafast imaging, volumetric sequences and cine-mode acquisition. These sequences can be used to provide information on target motion, OAR displacement to modify planning margins for RTP and to initiate image-guided radiotherapy strategies. A series of ultrafast imaging can be obtained by sequences such as echoplanar imaging. Evaluation of the position of OARs and tumours that are influenced by respiration or other internal organ activity during irradiation can be performed and modelled. MRI can provide better delineation of internal soft tissue structures S10

such as between myocardium and ventricular space compared with CT for heart dose–volume assessment by breath hold for left breast radiotherapy [65] or ultrafast acquisitions can be performed with normal respiration to assess the impact of irradiation compared with other treatment image-guided strategies. Ultrafast acquisitions may also be used to obtain gated images to model organ and/or target motion for RTP [6]. Sequential volumetric and cine-mode acquisitions can greatly aid image-guided radiotherapy strategies by providing data for the implementation of appropriate site-specific planning margins for both the target and OARs. Cine MRI can evaluate intrathoracic tumour mobility for patient individualization of treatment margins [66] and determination of the efficacy of freebreathing gating techniques for lung radiotherapy [67]. Cine MR has been used to assess intrafraction motion in prostate cancer [7, 68]. This intrafraction information can be used not only to determine internal margin size for RTP but also to estimate the degree of organ deformation that may occur during radiotherapy [69]. This issue of target volume deformation is also currently being investigated for bladder cancers in an image-guided program (POLO or Predictive Organ LOcalization) whereby cine MR studies are obtained to assess the temporal-spatial changes of the bladder as it fills during radiotherapy and the degree of tumour deformation that can occur during this period [70]. Ultrafast, volumetric and cine MRI can provide noninvasive means to evaluate not only variability in target volume positioning during radiotherapy but also the temporal variation in target volume deformation that may occur interfractionally and intrafractionally. These are pertinent issues that currently limit precision radiotherapy and justify the development of image-guided radiotherapy.

Developments in MR contrast agents Local-regional control is an important issue in radiotherapy treatment for many cancers. Adequate dose to involved lymph nodes can increase the probability of local control and this may translate into improved survival. Reducing dose appropriately to uninvolved local-regional nodal regions can substantially lower the probability of radiation related side-effects and also allow combination with systemic chemotherapy or radiosensitizers to be better tolerated or permit dose intensification. However, the determination of pathological lymph nodes is poor using current imaging techniques and surgery remains the gold standard in establishing the lymph node status. The difficulty in using simple size criteria alone to assess evidence of subclinical tumour involvement is well known. Up to 20% of normal size lymph nodes can be positive for microscopic disease whilst up to 30% of enlarged lymph nodes may only show inflammation. The likelihood of pathological nodal involvement increases with higher stage of disease. Furthermore, in MRI the signal intensity and degree of contrast enhancement between benign and malignant lymph nodes is not reliable enough to provide satisfactory discrimination. The British Journal of Radiology, Special Issue 2006


Ultrasmall superparamagnetic iron oxide particles (USPIO) USPIO particles have recently been reported to be suitable as contrast agents for the identification of and discrimination between normal and abnormal lymph nodes. USPIO particles are injected and taken up by macrophages and transported via the lymphatic system to the lymph nodes. In the reticuloendothelial tissues of normal lymph nodes, these macrophage ingested USPIO particles produce a reduction in signal intensity within the node due to the negative enhancement from the iron oxide particles. This lowering of signal intensity when scanned 24–26 h following administration can then be compared with the pre-USPIO scan. An example of this negative enhancement seen in normal lymph nodes is shown in Figure 8. In diseased nodes that are replaced by tumour, these USPIO particles within macrophages are prevented from occupying the node by the tumour and thus the signal intensity from pathologically involved nodes is preserved. Recent assessments using USPIO have reported a high sensitivity and specificity of up to 90% in cases with small volume tumour nodal involvement [3, 71]. If future studies confirm that this method of detecting early nodal involvement is reliable then the benefits for RTP are immense. This method using 3D sequences can provide visualization of lymph nodes along blood vessels and allow for mapping of lymph nodes according to surgical templates. This will optimize current target volume delineation for lymph nodes. There are many tumour regions where CT defined nodal volumes have only just superseded traditionally designed nodal volumes that are based on regional and bony landmarks. Inadequate nodal coverage can either compromise local control or unnecessarily irradiate normal structures [72, 73].

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The use of USPIO imaging can permit individualization of treatment fields. It can substantially influence local-regional volumes currently being prophylactically treated or permit dose escalation for involved lymph nodes. The treatment of head and neck tumours exemplifies this management challenge. USPIO imaging has been used for head and neck planning of surgery [71]. It can also complement the current recommendations for CT defined nodal volumes for head and neck radiotherapy [74]. It is important to note that whilst USPIO methods provide an advance in identification of pathological lymph nodes, it remains a morphological method and hence has its limitations. The threshold size for detecting pathological involvement may be 2–3 mm in a 5–10 mm node and false positives may occur if there is fibrosis or fatty replacement of the nodes [75]. It may be combined with other imaging methods such as PET to further increase its sensitivity and specificity.

Developments in MR techniques One of the well recognized advantages of MRI is its greater functionality in characterizing tissues. With this ability, MR may provide better knowledge of tumour extent through pathophysiology or tumour response compared with simple morphological assessments. This additional information on active tumour regions may be exploited in radiotherapy for the determination of boost volumes, dose escalation, combined therapy with chemotherapy or radiosensitizers, or to select prospective non-responders during a course of radiotherapy for more aggressive treatment. Cancer growth is usually accompanied by vascular extension and growth in order to meet the nutritional demands of rapid tumour expansion. These tumour features of vascular angiogenesis and increased cellular

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Figure 8. A case example of the use of ultrasmall paramagnetic iron oxide (USPIO) in MRI to evaluate pelvic lymph nodes in a man with prostate cancer. In this case, the lymph nodes (thick arrow) pre-USPIO (a) returned a negative MR signal following USPIO administration (thin arrow (b)) indicating normal lymph node architecture. This was later confirmed on lymph node sampling.


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growth are exploited by MR techniques such as MR diffusion and dynamic contrast enhanced MRI. These MR techniques have been available in the research domain for some time and are therefore technically not new developments. However, they are included here as they are not part of standard practice and their role in clinical management is still being assessed.

Dynamic contrast-enhanced MRI Dynamic contrast enhanced MRI (dcMRI) is a technique which exploits the vascular dynamics inherent in new blood vessel proliferation (neo-angiogenesis) and abnormal vasculature associated with tumour growth. Using fast MR sequences timed to capture the sequential changes in vascular perfusion following injection of lowmolecular weight contrast agents, the perfusion or enhancement of the examined tissues and/or organ can be characterized. Time dependent enhancement curves can be subsequently produced. These (qualitative) enhancement curves from different nominated regions of interest (ROI) within the tissues can then be compared and assessed. Different tissue enhancement curves can be recognized for normal, tumour and irradiated tissues [4]. In this manner, ROI exhibiting appropriate enhancement tumour curves can be selected or delineated for radiotherapy targeting or tumour boosting. The phases of the low-molecular weight contrast agents as they pass through the vascular system and tissues can be assessed by different MR sequences. T2 weighted sequences are more sensitive to the vascular phase and thus better display tissue perfusion and blood volume effects [76]. T1 weighted sequences are better at detecting the contrast agents in the extravascular to extracellular space and thus demonstrate microvessel perfusion, permeability and extracellular leakage effects [76]. They can be selected to provide different clinical assessments such as differentiating between benign or normal tissues, localizing active tumour regions and predicting and monitoring tumour response. T2 methods have been reported to be better suited for evaluating brain tumours, particularly gliomas [77], whereas T1 methods have been used more widely in breast, musculoskeletal, gynaecological and urological cancers [78–81]. These dcMRI methods have been used to detect tumour recurrence in previously irradiated breast [78] and prostate cancer sites [79, 82] as well as to predict response to radiotherapy in head and neck [83], rectal [80] and cervical cancers [81]. These methods can also be used to assess the nature of lymph nodes where pathologically involved nodes may possess appropriate tumour enhancement patterns on their signal intensity time curves. One limitation is that this technique can only assess a specific region and not all nodal sites.

Diffusion weighted MRI Diffusion weighted MRI (dwMRI) attempts to assess the diffusion capacity of tissue. This methodology relies on the tumour regions having increased cellular density due to tumour proliferation and therefore results in a S12

reduction of diffusion of water molecules through this abnormal region. Apparent diffusion coefficient (ADC) maps can be generated from different spatial regions of interest. A lower ADC is more likely to contain tumour than a high ADC. Early studies suggest that dwMRI may help distinguish between malignant and benign lesions in the brain [5], but dwMRI studies in other cancer types such as soft tissue sarcomas were not found to be as helpful due to substantial overlap in ADC values [84]. The use of dwMRI tumour defined areas may permit the assessment of tumour response during a course of therapy and regions of poor response may be selected for radiotherapy boosting. Preliminary dwMRI studies in rectal cancer treated using chemoradiation suggest that ADC values may provide indicators of tumour response [85, 86]. This can then be used to optimize treatment strategies during therapy or to initiate adjuvant therapy for the individual patient.

Diffusion tensor imaging Diffusion tensor MRI (DTI) is a technique that can demonstrate white matter abnormalities based on cerebral tissue anisotrophy (a measure of tissue disorganization) and provide information on brain tumour involvement on white matter tracts. Investigators using DTI suggest that this method may be useful for assessing white matter infiltration by occult tumour [87]. A recent study reported that DTI recorded larger white matter tract abnormalities than seen on T2 weighted images in 10 out of 13 high-grade gliomas and previously unrecognized contralateral involvement in 4 of these cases [87]. A further planning study suggested that treatment volumes may be optimized using DTI by reducing the PTV compared with CT planning alone and thus provide the opportunity of dose escalation whilst maintaining tissue tolerances [88]. If these findings are confirmed then this information will aid RTP and also provide prognostic information if the extent of invasion is a determinant of disease outcome. Magnetoencephalography and DTI can also be used in RTP to limit doses to relevant functional regions of the cerebrum or white tracts to reduce specific radiationinduced neurological dysfunction for each patient case to permit plan individualization [89]. DTI methods may offer complementary information to other imaging techniques including PET that are being used for target volume delineation in brain RTP.

Summary The superior characterization of soft tissues and visualization of tumour extent from MRI can benefit RTP by improving target volume delineation and assessment of planning margins in many cancer subtypes in sites such as the brain, spinal cord, soft tissues of the head and neck, trunk and limbs. In the past decade there have been many advances in MRI technology that can further aid the definition of volumes for both external beam radiotherapy and brachytherapy. Open and low field MR simulators can provide easier integration of MR for RTP by its lower cost, facilitating The British Journal of Radiology, Special Issue 2006


the use of patient immobilization devices and providing fewer image artefacts and less distortion. USPIO contrast agents can evaluate pathological lymph nodes for treatment. Ultrafast volumetric and cine mode sequences can provide temporal assessment of target volume deformity and positioning for image guided radiotherapy. 3 T scanners can provide higher resolution images for better tissue definition and can also benefit MRS applications. MR techniques using dcMR, dwMR and DTI can further assess target volumes with improved and complementary morphological, functional and biological data that can provide the opportunity to nominate biological target volumes and the potential to gauge treatment response. These techniques may also be combined with PET to further increase diagnostic sensitivity and specificity. This better estimation of target volumes may permit treatment individualization, organ sparing or functional avoidance, strategies of boosting and dose escalation with or without concurrent chemotherapy. All these new advances can increase the scope of MRI in radiotherapy, but currently their role and utility for RTP remains to be defined. However, just as important is the need for specific training for oncologists to understand how to utilize MR images for RTP. There must be close collaboration between diagnostic radiologists, physicists, radiographers and oncologists in order to effectively harness the benefits of MRI for radiotherapy. Ideally, this should be through a dedicated oncology team approach.

Acknowledgments We are grateful to Dr M Wada, Austin Health Radiation Oncology Centre, for providing the images demonstrating the base of tongue cancer case.

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Applications of magnetic resonance spectroscopy in radiotherapy treatment planning G S PAYNE,

DPhil

and M O LEACH,

PhD, FInstP, FMedSci

Cancer Research UK Clinical Magnetic Resonance Research Group, Institute of Cancer Research and Royal Marsden NHS Trust, Downs Road, Sutton, Surrey SM2 5PT, UK ABSTRACT. Following advances in conformal radiotherapy, a key problem now facing radiation oncologists is target definition. While MRI and CT provide images of excellent spatial resolution, they do not always provide sufficient contrast to identify tumour extent or to identify regions of high cellular activity that might be targeted with boost doses. Magnetic resonance spectroscopy (MRS) is an alternative approach that holds great promise for aiding target definition for radiotherapy treatment planning, and for evaluation of response and recurrence. MRS is able to detect signals from low molecular weight metabolites such as choline and creatine that are present at concentrations of a few mM in tissue. Spectra may be acquired from single voxels, or from a 2D or 3D array of voxels using spectroscopic imaging. The current state of the art achieves a spatial resolution of 6–10 mm in a scan time of about 10–15 min. Coregistered MR images are acquired in the same examination. The method is currently under evaluation, in particular in brain (where MRS has been shown to differentiate between many tumour types and grades) and in prostate (where cancer may be distinguished from normal tissue and benign prostatic hypertrophy). The contrast achieved with MRS, based on tissue biochemistry, therefore provides a promising alternative for identifying tumour extent and regions of high metabolic activity. It is anticipated that MRS will become an essential tool for treatment planning where other modalities lack the necessary contrast.

Magnetic resonance spectroscopy (MRS) is a noninvasive technique for measuring biochemicals in tissue. It uses the same general principles and equipment as its widely used partner, MRI. However, while MRI builds images using signals from 1H nuclei in tissue water (and sometimes lipid), present at concentrations of approximately 35 M, MRS is used to measure signals from magnetic nuclei (usually 1H, but 31P has also been extensively studied) of tissue metabolites such as choline, creatine and lactate that are present at much lower concentrations (typically of the order of a few mM). Example spectra from normal brain and from brain tumour are shown in Figure 1, to illustrate the type of data that are obtained. Further details of specific metabolites detected in brain and prostate are listed in Table 1 and discussed in the corresponding sections towards the end of the article. More detailed introductions to the use of in vivo MRS may be found in the literature [1, 2]. The frequency of MRS signals depends only on the gyromagnetic ratio (c) of the nucleus (which is a constant for a particular nuclear species, e.g. c(1H)52.6756108 rad/T/s; c(31P)51.0836108 rad/T/s), and on the local magnetic field, Blocal, experienced by the nucleus (Equation (1)). f (Hz)~c(rad=T=s)Blocal (T)=2p

ð1Þ

GSP is funded on a research grant from Cancer Research UK ([CUK] grant number C1060/A808).

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Received 3 January 2006 Revised 7 April 2006 Accepted 30 May 2006 DOI: 10.1259/bjr/84072695 ’ 2006 The British Institute of Radiology

Blocal is determined primarily by the applied magnetic field (e.g. 1.5 T for a standard clinical MR system). However, it is modified slightly by shielding effects of the electron cloud associated with covalent bonds to neighbouring atoms (usually of a few parts per million (ppm)), by magnetic field inhomogeneity, and by the presence of local nuclear or electronic magnetic dipoles. The horizontal ‘‘frequency’’ axis is usually expressed in parts-per-million of the main precession frequency (approximately 63 MHz at 1.5 T) so that spectra acquired at different magnetic field strengths may be compared. By convention (and for historical reasons), frequency increases from right to left. Following excitation of magnetic nuclei using a radiofrequency (RF) pulse, signal is acquired as the spins relax back towards their equilibrium state. The time taken for this relaxation to occur is described by two time constants, T1 (for longitudinal relaxation) and T2 (for transverse relaxation). T1 is typically of the order of 1s, and T2 (never longer than T1, and usually much shorter) may be about 100 ms. These relaxation time constants determine how quickly the measurement may be repeated to improve the signal-to-noise ratio. Incomplete relaxation can be used to weight signals according to their characteristic T1 or T2, which under some circumstances (and particularly for MRI) yields better contrast than concentration alone. One key advantage of MRS relative to radio-labelled studies is that the same type of magnetic nucleus (e.g. 1H) produces signals at different frequencies, depending on The British Journal of Radiology, Special Issue 2006

MR spectroscopy and radiotherapy treatment planning

Figure 1. Example of 1H MRS spectra from (a) normal brain and (b) brain stem tumour. The major peaks are from choline (Cho), Creatine (Cr), myo-Inositol (mI), N-acetyl aspartate (NAA) and lipids (lip). Repetition time (TR)51.5 s; echo time (TE)530 ms. Each spectrum shows acquired data (grey line), the fit (bold line) and fitted baseline. Note the increase in choline (Cho) and reduction in NAA in the tumour compared with normal brain.

the bonding to neighbouring nuclei. Thus, different nuclei of the same species within a single molecule produce distinct peaks within a spectrum (unless they are ‘‘magnetically equivalent’’), which usually permits different compounds of biochemical interest to be distinguished. In addition, one can study compounds naturally present in tissue without the requirement for specific labelling. The major disadvantage of MRS is that the signals from metabolites are relatively small, so that compared with MRI much larger ‘‘pixels’’ or ‘‘voxels’’ are required to obtain an adequate signal-to-noise ratio. Methods for data acquisition and some related issues are described in the section on ‘‘Technical Issues’’ below.

In contrast to CT and MRI, which provide morphological information about tissue, MRS gives biochemical information (Figure 1). Since the biochemistry of tumours is substantially different from that of normal tissue, MRS has the potential to aid identification of tumours when there is insufficient contrast in the morphological image. In particular, it can be used for differential diagnosis, both between tumours and benign pathology (e.g. prostate cancer and benign prostatic hypertrophy [3]), and between different tumour types (e.g. different brain tumours [4]). Since MR spectra are acquired using MRI scanners, they are automatically coregistered with MR images, which may then themselves

Table 1. Brief details of some metabolites seen in 1H spectra of tissues Metabolite

Chemical shift of main peak (ppm)

Number of equivalent H nuclei

Multiplicity

Description

‘‘Cho’’ includes contributions from choline, phosphocholine, glycerophosphocholine and other trimethylamines. These metabolites are involved in cell membrane lipid synthesis and breakdown, and are also affected by signalling pathways that can be upregulated in tumours. Since 9 magneticallyequivalent protons contribute to this peak, relatively low concentrations produce a measurable signal ‘‘Cr’’ includes creatine and phosphocreatine, which are both involved in energy metabolism

1

Cho

3.2

9

singlet

Cr

3.02

3

singlets

Lactate

3.9 1.33

2 2

doublet

Lipids myo-Inositol (mI)

1.3, 0.9 etc. 3.52

2

3.6

2

singlet doublet of doublets triplet

NAA

2.01

3

singlet

Citrate

2.6

1

multiplet


Lactate is a product of anaerobic glycolysis, a further aspect of energy metabolism, often being found in necrotic areas Often found in necrotic regions This can be detected in brain using shorter echo time acquisitions. Understood to be an essential ingredient for cell growth, an osmolite, and a storage form of glucose NAA is considered to be a neuronal marker, so only present in brain This is synthesized and accumulated by normal prostate epithelial tissue

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G S Payne and M O Leach

be co-registered with planning CT scans in the normal way if required (see also the paper by Kessler in this issue). In the context of radiotherapy, MRS has several potential applications. These include: 1. Identifying tumour extent and metabolically active regions to aid targeting of radiotherapy. This includes distinguishing tumour from normal tissue, benign lesions, or radiation necrosis. 2. Evaluating response to treatment 3. Identifying recurrence With the advent of new conformal techniques for radiotherapy, target definition has become one of the main issues in further improving the effectiveness of radiotherapy treatment. In particular, it is necessary to delineate accurately the extent of tumour tissue, and one may also wish to identify regions of high clonogenic cell density, which could be targeted with boost doses of radiotherapy using either intensity-modulated external beam radiotherapy or brachytherapy. Conventional imaging methods are often limited in this respect. For example, in the prostate, MRI has better soft tissue contrast than CT [5], but still only has a positive predictive value of 50% [6]. In addition, owing to the different sources of contrast in the two cases, the boundaries shown are often different [7, 8]. Since MRS measures tissue biochemistry, it has the potential to identify abnormalities and the nature of the abnormality more selectively than morphological imaging can achieve.

Current status of MRS in target planning for radiotherapy 1

H MRS packages are currently available for most commercial clinical MR scanners, usually including both single-voxel localization techniques and MR spectroscopic imaging (MRSI; for more details see subsection on localization strategies below). However, owing to the extra expense, and the time required to perform MRS examinations, many hospitals choose not to have this option. The packages are increasingly easy to use, but previous experience is important to achieve good results, and to understand the reasons for technical failures. Assuming diagnostic MR images have already been acquired, the extra time required for 1H MRS is typically a few minutes for single-voxel techniques, and 10– 15 min for MRSI. Some software for spectral processing is usually provided, but for quantitative work many users choose to export their data to specialist spectral processing software. MRS of nuclei other than 1H is less used owing to the requirement for additional hardware and (usually) longer scan times. While MRS has great potential to aid target identification in radiotherapy treatment planning, it is not currently used to any significant extent for this purpose. This is partly because it is only recently becoming available in a form that is relatively easy to use. In order to alter current treatment protocols to utilize MRS, the following steps need to be achieved: S18

N N

N N N

to validate spectroscopic observations against histology to demonstrate a correlation between MRS observations and survival or local recurrence based on existing treatments, and hence to define a measure of abnormality in different cases to use for delineating the treatment target volume to demonstrate improvement in survival in welldefined clinical trials to make the current state-of-the-art measurements routinely and robustly available, with tools to aid interpretation of spectroscopic data to educate the radiotherapy community

To improve spatial resolution (currently about 6– 10 mm) would be helpful but probably not essential, as practical radiotherapy techniques will include a margin of this order. As usual, there is a trade-off between signal amplitude and spatial resolution, so improving spatial resolution requires an improved methodology to increase the sensitivity of signal detection. Some of these issues are already being addressed, in particular in studies of the brain and prostate (see below).

Technical issues Magnetic field strength Most clinical MR studies are performed at a magnetic field strength of 1.5 Tesla and most published MRS work, including most of that referred to in this review, has been performed at this field strength. In research environments, 3 T scanners are becoming increasingly used, as the signal-to-noise ratio and spectral resolution both increase (approximately linearly) with field strength. The doubling of signal-to-noise ratio at 3 T can be traded for a 50% reduction in voxel volume (i.e. a reduction in voxel size of (0.5)1/3 < 0.8 in each spatial dimension) compared with 1.5 T scanners, but otherwise the functional information is very similar. Systems with much higher magnetic fields for human studies at 7 T and 8 T have been installed in a handful of research sites. However, these are very expensive, present a large number of technical challenges to operate (in particular the correspondingly higher gradient amplitudes and radiofrequency power deposition can exceed safety limits unless duty cycles are reduced) and generally (except in the brain) do not produce good diagnostic images owing to dielectric resonance effects [9]. Problems with image distortion (see below) will also increase at the higher fields. Shimming The sharpness of spectral lines depends on the homogeneity of the magnetic field. This is because a spread in values of the local magnetic field leads to a spread of frequencies (Equation (1)). The effect of placing a sample into a homogeneous field usually disturbs the field homogeneity. MR systems are equipped with a series of ‘‘shim coils’’ which create small additional magnetic fields that can be adjusted to help counteract this effect. This ‘‘shimming’’ process is usually performed at the beginning The British Journal of Radiology, Special Issue 2006


of each MRS study, and takes a few minutes. On current MR systems the process is normally automated.

Magnetic nuclei studied Because clinical MR scanners are designed to detect signal from 1H nuclei, only additional software is required to acquire 1H MR spectra. In addition 1H is the most sensitive magnetic nucleus, and hydrogen is present in nearly all biologically relevant compounds. A list of metabolites seen in 1H spectra of several tissues is given in Table 1. More details for brain metabolites can be found in the work of Govindaraju, Young and Maudsley [10]. If nuclei within a molecule are ‘‘equivalent’’ then they all contribute to the same peak. Peaks appear as multiplets (doublets, triplets etc.) when the nucleus of interest is covalently bound directly or through other nuclei to another (but non-equivalent) magnetic nucleus (1H, 31P etc.). This splitting (known as J-coupling) decreases as the number of bonds between the nuclei increases. In order to study compounds at millimolar concentrations, it is necessary to suppress the large signals from tissue water, and also sometimes from lipids. Suppression techniques are available to accomplish this. For comparison of spectra it is necessary to have a reference frequency. 1H MRS studies in solution often include tetramethylsilane (TMS) for this purpose, with a frequency (or ‘‘chemical shift’’) set to 0 ppm. TMS is used because it is stable and produces a sharp single peak. For consistency, the positions of spectral peaks in vivo are still expressed relative to TMS, even though it is not present in tissue; other peaks that are present, and which have shifts that are relatively independent of environmental conditions such as pH and metal ion concentration etc., are used as an internal secondary reference. Other nuclei that may produce useful MR spectra are 31 P, 13C and 19F. Phosphorylation is an important biochemical process, and several compounds of key importance therefore contain phosphorus. Several metabolites seen in 31P MR spectra are affected by signalling pathways up-regulated in cancer [11, 12] and therefore 31P MRS could play an important role in target identification and evaluation of therapy, especially of new therapeutic agents. One study [13] in soft tissue sarcoma treated with thermoradiotherapy has shown strong correlations between pre-treatment MRS data (in particular the ratios of phosphodiesters and of phosphomonoesters to inorganic phosphate) and pathological complete response. However, since 31P has only 6% of the sensitivity of 1H at a given field strength, larger voxels are required to obtain an adequate signal-to-noise ratio. In addition, extra hardware (RF amplifiers, filters, RF coils) is required that is tuned to the 31P resonance frequency. Currently, little work has been performed using 13C nuclei (owing to only 1% natural abundance and a lower sensitivity). While no endogenous metabolites contain MR-visible fluorine, several studies have demonstrated the value of 19F MRS in following drug metabolism [14]. Localization strategies MRS localization techniques fall into two groups. Single-voxel methods include those known as PRESS The British Journal of Radiology, Special Issue 2006

[15], STEAM [16] and ISIS [17], in which data are acquired from a single voxel positioned using MR images. The alternative is MRSI, also known as chemical shift imaging, (CSI) [18], in which a matrix of spectra are acquired either over a plane (2D-MRSI) or a volume (3DMRSI). The choice of localization method affects the efficiency of signal detection from within the target region and the contribution of unwanted signals from outside the volume (contamination). The effectiveness of the shimming method in optimizing the magnetic field homogeneity in the selected region, the effectiveness of suppressing the very large signals from water (and sometimes from lipid present in the selected volume) and the ability to suppress signals from outside the volume of interest can also affect spectral quality. In the brain, cerebral spinal fluid flowing into the selected region can also adversely affect shimming and water suppression. Sequence timing parameters determine the T1 and T2 weighting of the sequence, and also affect modulation of coupled spins. Common echo times (TE) are 270 ms (minimizing contributions from lipids and macromolecules), 135 ms (inverting the lactate doublet) and 20 ms (maximizing signal from macromolecules and lipids). A summary of the merits of single voxel MRS relative to MRSI is given in Table 2. For most applications in radiotherapy treatment planning use of MRSI will be preferred, with single voxel methods restricted to cases where it is required to characterize a single welldefined lesion, when time is short, or when adjacent structures make it hard to shim over larger regions.

Registration and motion issues The volumes selected for acquisition of the MR spectra are positioned using MR images acquired in the same examination. The MRS and MRI data are therefore inherently co-registered. However, although MR images have better soft tissue contrast compared with CT images, they have two major disadvantages for radiotherapy treatment planning – they suffer distortions and they do not intrinsically provide electron density information. These issues have been extensively discussed elsewhere [19]. There are two approaches to solve this problem. One approach under development is to correct the distortions [20, 21] and to make use of the MR images to estimate the electron density information [22]. The alternative is to co-register the MR images with CT images – usually using either external fiducial markers that can be seen by both imaging modalities, or using methods such as ‘‘mutual information’’ [23] based on the features within the images (see also the paper by Kessler in this issue). It is also necessary to ensure that the MR scanner is equipped with a flat-topped couch (similar to that used for treatment), and with suitable laser positioning beams. Motion during the scan is a potential problem. Being rigid, immobilization of the head is fairly straightforward, although there is still some residual pulsatile motion from the circulation of cerebral spinal fluid. Other parts of the body are more challenging. For studies of the prostate it is common to use agents like Buscopan to reduce peristalsis. Data acquisition can be synchronised to the heart or breathing cycle using cardiac or respiratory triggering, while ‘‘navigator echoes’’, an MR S19

G S Payne and M O Leach Table 2. Comparison of single voxel techniques with spectroscopic imaging

1. Specification of VOI No. of volumes Voxel specification Voxel shape 2. Voxel integrity Edge definition Chemical shift displacement artefact 3. Sensitivity Relaxation losses

Single voxel

Spectroscopic imaging

1 Must be specified before measurement Dimensions adjustable; also tilt

Typically 86868 (3d) or 16616 (2d) Many voxels available Grid shift can be performed retrospectively Orthogonal (usually square) grid

That of slice profile Yes – in each direction

Point spread function effects Not present

T2 (STEAM, PRESS)

T2 (in STEAM and PRESS pre-localized implementations) Phase-encoding loses approximately 13% in each spatial dimension Sometimes poor. Addition of small voxels does not recover SNR of the larger corresponding voxel

Other losses

Imperfect RF pulse profiles and flip angles

Conformation to target

Usually good

4. Other aspects Minimum number of acquisitions RF coils – uniform transmit coils and surface coil receiver RF coils – transmit/receive surface coil

1

Many (e.g. 512 for 86868 csi)

Good

Good

ISIS works well provided adiabatic RF pulses are used; PRESS and STEAM rely on slice selection which is not good with the non-uniform transmit fields from surface coils

The basic implementation works well if an adiabatic excitation pulse is used; additional localization (slice or volume selection) have same problems as PRESS and STEAM

VOI, volume of interest; RF, radiofrequency; SNR, signal to noise ratio.

technique that identifies a column of tissue and monitors movement along this column [24], may also be used. It is not possible to eliminate motion entirely and, as with conventional CT techniques, this must be remembered when prescribing the margins for irradiation. As the application of MRI and MRS for radiotherapy planning develops, it is likely that use of immobilization devices will be evaluated in more detail.

Validation Before radiation oncologists have the confidence to alter treatment plans based on MRS data, they need evidence that the abnormality detected using MRS represents tumour. The ‘‘gold standard’’ for comparison is histopathology (where available). In some cases, such as prostate (see below), studies have been performed where the prostate has been removed very shortly following the in vivo MRS examination, so that MR spectra can be directly compared with histopathology of the whole gland. This is the best possible situation for validation. In other cases biopsy samples are available, although it is often difficult to be certain of their precise location. Results of some of these studies are described below. Interpretation of the in vivo spectra can be aided by high-resolution ex vivo MRS of intact tissue samples acquired using magic angle spinning (MAS) at high magnetic fields [25]. Such spectra have much higher spectral resolution and sensitivity than spectra acquired in vivo, and permit detection and quantification of metabolites that in the in vivo spectrum are either at too low a concentration or are in regions of spectral over-lap. Other imaging methodologies (MRI, PET etc.) give S20

supporting evidence, but may themselves lack the necessary validation.

Example 1: MRS and brain cancer General issues regarding diagnosis and treatment There is a strong incentive for non-invasive assessment of brain cancer, particularly in children, and especially where sequential evaluations are required. In managing patients, the first priority is diagnosis, currently usually confirmed by histology, although this suffers from known limitations of sampling error and tumour heterogeneity. Guidance of biopsy is therefore also important. Assessment of tumour grade is important in defining treatment – this is usually limited to one time point, either at biopsy or during tumour debulking. Assessment of response following treatment is of increasing value, so that the treatment can be modified according to the response observed. MRI appearance and location alone have limited power to differentiate brain tumours. In particular, early stage disease and infiltrative disease may have little impact on the blood–brain barrier, limiting the sensitivity of contrast-enhanced MRI and leading to the requirement for an alternative method to identify tumour boundaries. Identification of residual disease from necrosis can also be difficult with MRI. The metabolic profile provided by MRS offers considerable potential to distinguish different tumour types and grade. Most MRS measurements of tumours have employed the 1H nucleus, owing to its greater sensitivity. 1H MRS The British Journal of Radiology, Special Issue 2006


of brain tumours has developed rapidly due to the less exacting technical requirements of spectroscopy in the brain compared with measurements elsewhere in the body. In particular, it is easier to control motion and most parts of the brain are fairly homogeneous, making it easier to achieve a good shim. Care should be taken with MRS when using contrast agents as this can broaden the choline peak, without affecting peak area. However, if shimming takes place immediately prior to MRS acquisition, this effect can be minimized or eliminated [26].

Features of 1H brain tumour spectra Spectra from brain tissues (Figure 2) include contributions from total creatine (Cr) (3.94 ppm and 3.01 ppm), total choline (Cho) (3.22 ppm), N-acetyl aspartate (NAA) (2.01 ppm), myo-inositol (mI) and glycine (3.55 ppm), lactate (Lac) (1.35 ppm), alanine (Ala) (1.47 ppm), contributions from a range of lipid resonances (lip) (0.9 ppm, 1.3 ppm), broad resonances due to macromolecules (2.05–2.8 ppm, 5.4 ppm) and poorly resolved amino acids such as glutamate and glutamine (Glx) (approximately 2.3 ppm). In tumours other resonances may be present, or may contribute to peaks, but their contribution can be hard to confirm unless tissue specimens are available.

Application in differential diagnosis and assessing the grade of brain tumours While MRI is an essential part of the evaluation of cranial neoplasms, indicating location, morphology, boundary definition and physiological characteristics, there remains considerable overlap between the appearance of primary tumours, metastases and radiation necrosis [27–33]. However, as illustrated in Figure 2, MR spectra do display differences between tumour types. For example, compared with normal brain, MR spectra of astrocytomas show elevated Cho, reduced Cre and significantly reduced NAA. A summary of the characteristics found in some common brain tumours is given in Table 3. More details can be found in Howe et al and Devos et al [28, 34]. While visual inspection of spectra can be helpful, there is considerable interest in automated evaluation of multiple spectral lines [29, 34, 35]. In particular this permits comparison of new spectra with those from tumours of known histology, provided the spectra are acquired under similar conditions (magnetic field, sequence timing parameters etc.). This approach is showing promising indications of improved discrimination of tumour type and grade. Since there are many tumour types, some of which are relatively rare, effective databases require many samples, thus necessitating multi-centre cooperation.

Figure 2. Mean and standard deviation (vertical lines) of normalized STEAM (echo time (TE)530 ms) spectra acquired from a number of subjects: Normal white matter (n56); meningioma (n58); astrocytoma grade II (n55); anaplastic astrocytoma (n57); glioblastoma (n513). Published with permission [28].


S21

G S Payne and M O Leach Table 3. Summary of 1H MRS characteristics of different brain tumours Tumour type

1

All tumour types Astrocytomas

Low NAA, and reduced Cr Elevated Cho, reduced Cre and significantly reduced NAA. With increasing WHO grade, methylene lipid (1.3) and lactate may be seen, indicating necrosis. Methylene levels correlate with contrast enhancement in MR images, and may allow the transformation from low to high grade to be detected prior to focal contrast enhancement [63]. Cho may also increase with the increased proliferative activity of higher grade tumours. Glioblastoma multiforme has the same spectral pattern as metastasis May show increased mI mI may also be elevated, together with Cho

Low grade tumours Oligodendoglioma and mixed oligoastrocytoma Meningioma Metastasis

Radiation necrosis

H MRS characteristics

Low [Cr] and [mIG], increased Cho, although there may be low levels of lipids and a characteristic presence of alanine (1.47) Similar features to astrocytomas, with increased lipid if necrosis is present. However, they may have a distinct spectroscopic boundary. Thus the presence of elevated choline/creatine ratio in the peritumoural region may suggest high grade glioma rather than a solitary metastasis [64] In some cases produces a peak at 2.4 ppm [65]

Therapeutic guidance, assessment of response and recent developments Recently there has been increasing interest in using functional imaging techniques, together with metabolic imaging using MRS [36] to aid target definition in radiotherapy, for example to identify areas of infiltration not evident with MRI. Figure 3 illustrates an example where the boundaries of the lesion detected using MRS (based on levels of Cho and NAA) are quite different from those suggested by contrast-enhanced MRI [37]. In a further study [38] of patients with high grade glioma who were scanned following surgery (but prior to radiotherapy), MRS indicated areas of residual abnormality that did not enhance with contrast at that time, but of which a proportion did enhance subsequently at

follow-up. In this study, MRS volumes were also greater than the T2 weighted area of abnormality. A study of survival following gamma knife surgery [39] showed that patients with poor overlap (,50%) between treatment volume and volume of metabolic abnormality had reduced survival compared with those having good overlap. Elevated Lip/Cre and Lac/Cre in the peri-tumoural region of high grade glioma may help to identify patients at risk of recurrence [40]. It is suggested that areas of relatively high Cho/NAA may indicate high cellular activity, and hence radio-sensitivity, and Lac may indicate hypoxic areas with reduced radio-sensitivity [37, 41]. The technique can also be helpful in identifying areas missed by radiation fields, and in separating

Figure 3. Comparison of lesion extent measured using T2 weighted MRI (red), T1 weighted MRI following injection of contrast agent (green), and MRS to measure high Cho/NAA (orange) for a patient with a grade 4 glioma. Published with permission [43]. S22



Figure 4. Centre: T2 weighted transverse image of prostate with tumour in right mid-gland, overlain with grid of voxels from which spectra were acquired. Left: Spectrum acquired from region of tumour, illustrating dramatically elevated choline and absence of citrate and polyamines. Right: Spectrum from healthy peripheral zone tissue with high citrate and presence of polyamines. (Reproduced with permission from J Kurhanewicz et al [46]).

recurrence from radiation necrosis. Reduction in Cho [42] and in lipid and lactate [43] can also reflect response to chemotherapy and radiotherapy. MRS has been compared with MRI and with 201TlClSPECT (a radio-nuclide technique that is widely used to assess brain tumours) in monitoring response to treatment in metastatic brain tumours. Reductions in Cho and Lac and an increase in Lip (believed to represent necrosis) in responding tumours were detected at an earlier time (1 week to 1 month) than contrast-enhanced MRI or 201TlCl [44]. A short term change in metabolites reflecting both pathology and normal brain tissues has been reported within 1 day of radiotherapy [45] although the significance of this is not clear, as a short term metabolic response to immediate damage delivered by radiotherapy may not translate into a therapeutic gain.

Example 2: MRS and prostate Introduction In the prostate, T2 weighted MRI has much superior contrast to CT, transrectal ultrasound and digital rectal examinations. However, the sensitivity and positive predictive value are still only of the order of 83% and 50%, respectively [6]. Several investigators are currently evaluating the use of 1H MRSI in the prostate to aid target definition. The main spectral peaks observed in normal prostate (Figure 4, right) are those of the choline-containing compounds (Cho) at 3.2 ppm, creatine and phosphocreatine (3.02 ppm) and a large peak from citrate (2.6 ppm). This latter peak is tightly coupled, and the appearance depends on the echo time used to acquire the data [47]. As in other tissues, choline compounds are associated primarily with membrane synthesis, while creatine is involved in energy metabolism. Citrate is a product of normal epithelial cell metabolism in the prostate, where high levels of zinc inhibit the enzyme aconitase and hence prevent the oxidation of citrate in the Krebs cycle that occurs in other cells. It is regulated by testosterone and prolactin [48]. Sometimes an additional peak may be detected between the choline and The British Journal of Radiology, Special Issue 2006

creatine peaks that arises from spermine, a polyamine characteristic of healthy prostate and which is also under androgen control. In prostate cancer (Figure 4, left) choline is elevated and the normal production of citrate is reduced. In contrast, benign prostatic hyperplasia, an enlargement of the prostate commonly found in older men, is characterized by high levels of citrate. Hence the choline/citrate ratio is a fairly reliable measure of the presence of cancer. Lipids are also sometimes seen in cancerous tissues, although the significance of these is yet to be established. 1 H MRSI data may be acquired from the prostate using an external phased-array coil. However, the best signalto-noise ratio is achieved using an endorectal coil. This is usually well tolerated by patients. The main disadvantage of the endorectal approach is slight deformation of the prostate which needs to be allowed for in using the images for radiotherapy treatment planning. One study suggests that rigid endorectal coils are less of a problem in this respect than the more usual balloon coils [49]. Buscopan is often used to reduce involuntary motion. While many studies [50, 51] have used slice-localized 2DMRSI, full 3D-MRSI is much preferred to obtain data from the whole of the prostate. Currently, the state-ofthe-art at 1.5 T is to achieve voxels with a 3D isotropic resolution of 6.25 mm in an acquisition time of 17 min [52]. Studies have shown that owing to haemorrhage there is some degradation of in vivo spectra in the 8 weeks following transrectal biopsy (18.5% of peripheral zone voxels have been reported as degraded within 8 weeks of biopsy, and 7% for those examined more than 8 weeks after biopsy [53]).

Validation of MRS in prostate for target definition A strong correlation has been found between negative MRSI and negative biopsy findings, and between positive MRSI and positive biopsy findings [54]. However, there is only a weak correlation between the concentration of prostate specific antigen (PSA, the current ‘‘gold standard’’) and either biopsy or MRSI findings [54]. Step-section pathological examination of radical prostatectomy specimens demonstrated that MRI S23


combined with MRSI yielded a significant improvement in cancer localization to a prostate sextant (left or right; base, mid-gland or apex) compared with MRI alone [55]. Several studies have shown that adding MRSI to an MRI examination increases the accuracy of diagnosis [55–57]. One particular area of high current interest is in discriminating the many patients who present with elevated PSA, but who have pathologically indolent cancer from those with aggressive disease; preliminary studies suggest MRSI has a useful role to play here as well [52]. High-resolution studies of tissue samples using MAS support these findings, with linear correlations measured between metabolite levels characteristic of normal epithelium or of prostate cancer, and the proportions of the corresponding cells as measured by computer-aided image analysis of prostate pathology slides [58].

Recurrence Use of MRSI together with MRI has been shown to improve substantially the identification of tumour recurrence following external beam radiotherapy (the area under the ROC curve, a measure of the effectiveness of a test, increased from 0.5 to 0.81 [59]). The presence of 3 or more suspicious voxels in a hemiprostate showed a sensitivity and specificity of 89% and 82%, respectively, for the diagnosis of local recurrence. Planning MRSI has been used in combination with MRI to define regions for dose escalation within the prostate [60–62], permitting a dose of . 90 Gy to the high-risk region while treating the remainder of the prostate to about 70 Gy.

Discussion MRI already plays a major role in identifying the extent and position of tumours, aiding delineation of target volumes. Increasingly, there is interest in defining the functional extent of tumours to better select target volumes, or areas that might receive a boost dose. MRI already provides the capability to image several characteristics of tissues which reflect the physiology and microenvironment, such as perfusion and diffusion. MRS provides the ability to observe aspects of tissue metabolism, which is a more direct reflection of tumour activity and of therapeutic response. The ability to obtain this information at the same examination as the anatomical and functional MRI images is a major advantage compared with other techniques. While the number of studies so far is limited, there is strong evidence that MRS may have a valuable role to play in radiotherapy treatment planning. The metabolic information provided by MRS provides signals associated principally with tumour (elevated Cho, Lac, Lip) and with normal tissue (NAA, Cre, Cit), where Lip is likely to be associated with necrosis, allowing identification of different tissue components, and providing measures that may reflect tumour grade, normal tissue response, and areas of mixed tumour and normal tissue. The mismatch between metabolic information, function and anatomical abnormality is not S24

surprising. In brain tumours, contrast enhancement depends on breakdown of the blood–brain barrier, which is not apparent in many low grade gliomas, and may not be present in early areas of infiltrative disease. Thus metabolism may be a better guide to early disease and disease extent compared with contrast enhancement. High signal on T2 weighted images reflects increased water content or local oedema, which may reflect inflammatory processes. Again, metabolic change could occur sooner, depending on the process controlling increased water content. The main limitation with MRS is that voxel sizes of typically 8 mm3 to 10 mm3 are required to achieve an adequate signal-to-noise ratio, with a large part of this volume needing to be occupied by abnormal tissue for a change in signal to be detected. Thus, small infiltrating lesions are unlikely to be detectable. It also needs to be remembered that the edges of target volumes defined from spectroscopic images for treatment planning, although smoothed and interpolated for presentation, are in fact limited by this same spatial resolution and will also include some blurring from point spread function effects. While some improvement in sensitivity and spectral specificity is expected with higher field scanners and improved sensitivity coils, this is likely to yield only a small improvement in spatial resolution. On the other hand, it is increasingly clear that while in principle both MRI and CT have much better spatial resolution, MRS, has the potential to improve identification of the gross tumour volume and hence improve treatment using radiotherapy.

Acknowledgments We wish to thank Dr Phil Evans and Dr Andrew Jackson (Institute of Cancer Research and Royal Marsden Hospital NHS Foundation Trust) for useful discussions.

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The role of PET/CT scanning in radiotherapy planning 1

P H JARRITT,

PhD, FIPEM,

1

K J CARSON,

PhD,

1

A R HOUNSELL,

PhD

and 2D VISVIKIS,

PhD

1

Northern Ireland Regional Medical Physics Agency, Royal Victoria Hospital, Belfast, UK and 2U650 INSERM, LaTIM, Brest, France

ABSTRACT. The introduction of functional data into the radiotherapy treatment planning process is currently the focus of significant commercial, technical, scientific and clinical development. The potential of such data from positron emission tomography (PET) was recognized at an early stage and was integrated into the radiotherapy treatment planning process through the use of image fusion software. The combination of PET and CT in a single system (PET/CT) to form an inherently fused anatomical and functional dataset has provided an imaging modality which could be used as the prime tool in the delineation of tumour volumes and the preparation of patient treatment plans, especially when integrated with virtual simulation. PET imaging typically using 18F-Fluorodeoxyglucose (18F-FDG) can provide data on metabolically active tumour volumes. These functional data have the potential to modify treatment volumes and to guide treatment delivery to cells with particular metabolic characteristics. This paper reviews the current status of the integration of PET and PET/CT data into the radiotherapy treatment process. Consideration is given to the requirements of PET/CT data acquisition with reference to patient positioning aids and the limitations imposed by the PET/CT system. It also reviews the approaches being taken to the definition of functional/tumour volumes and the mechanisms available to measure and include physiological motion into the imaging process. The use of PET data must be based upon a clear understanding of the interpretation and limitations of the functional signal. Protocols for the implementation of this development remain to be defined, and outcomes data based upon clinical trials are still awaited.

In considering the potential role of functional imaging in radiotherapy planning, it is necessary to put the development of PET/CT imaging into the context of developments in the delivery of radiotherapy treatments. Radiotherapy is in a period of rapid scientific and clinical development, in part due to the introduction of devices aimed at the better and more reproducible delivery of external beam therapy through the use of a number of technological advances. These include computer controlled linear accelerators, multileaf collimators (MLCs), electronic portal imaging devices (EPIDs) and kV imaging units attached to the linear accelerator for cone beam tomography (CBT) (see Moore et al in this issue). Portal images and CBT imaging provide ‘‘at treatment’’ images, allowing increased positioning verification and, ultimately, targeting in treatments requiring multiple irradiations. This is becoming known as image-guided radiotherapy (IGRT). The introduction of MLCs enables irregularly shaped treatment fields to be used to better conform the treatment field to the target. With the development of sophisticated software planning algorithms, MLCs can be used to deliver intensitymodulated radiation therapy (IMRT) treatments in which the intensity through the treatment beam is varied. Combining several IMRT treatment portals can result in complex cross-sectional dose distributions being achieved and even the delivery of high dose areas within the target, a technique that has been referred to as ‘‘dose painting’’ [1]. Beyond this, gating systems are being introduced which permit the radiation to be delivered in synchrony The British Journal of Radiology, Special Issue 2006

Received 11 July 2005 Revised 26 October 2005 Accepted 25 January 2006 DOI: 10.1259/bjr/35628509 ’ 2006 The British Institute of Radiology

with physiological motion such as respiration [2–5]. This is then coupled with the development of treatment plans and target volumes in equivalent physiological states requiring 4D plans including space and time. The raison d’etre for these technological advances is that a more accurate delivery of radiation to appropriate target volumes will lead to less damage to ‘‘normal’’ tissue whilst permitting higher, more targeted and non-uniform doses to be applied to the ‘‘diseased’’ tissue volume. In order to fully exploit these delivery system developments, the treatment volume planning process needs accurate information about the spatial location and extent of the target volume. In addition, information about intrafractional motion of the target and the impact of this on the delivery and treatment plan is required. It is necessary to define what is meant by radiotherapy treatment planning within the context of this review. Treatment planning may be regarded as encompassing the process that the oncologist undertakes when a patient is referred into their care. This includes gaining accurate diagnostic and staging information from a range of tests, planning the overall treatment as well as defining the target volume and specifying the radiation treatment portals. A more specific definition of treatment planning is to consider only the process of target delineation and the subsequent determination of the treatment beam shapes and orientation. In this context, it is assumed that the oncologist is fully aware of all the staging issues such that the role of the PET image is now to support the delineation of the tumour volume. For the purposes of S27

P H Jarritt, K J Carson, A R Hounsell and D Visvikis

this review, treatment planning shall be considered to be the process of target volume delineation. Radiotherapy treatment plans are developed using X-ray CT images. These provide data about the X-ray attenuation characteristics of the patient as well as identifying structures and organs for the specification and delineation of treatment volumes. It is clearly understood that this property may not reflect the underlying cellular and biochemical processes and that for the treatment to be effective it must also target those tissues where function rather than anatomy is abnormal. This has therefore led to the investigations into and the introduction of imaging methodologies which can supplement the purely anatomical information of X-ray imaging by providing data regarding molecular function. MRI can provide functional information, although limited in extent (see Khoo et al, and Payne and Leach in this issue). However, techniques based on the radioactive tracer method offer sensitivities to biochemical processes which are 103–106 higher than MRI, depending on the radioactive tracer and the detection technology. Positron emission tomography (PET) is one such technology. Current systems combine a PET detector with a CT scanner and can provide geometrically aligned anatomical data (the basis of the current treatment planning process) and functional information at the molecular level [6]. These functional data can be readily introduced into the planning process using current generation radiotherapy treatment planning systems. Oncologists have used PET images in the diagnostic and staging process, especially to make decisions about the inclusion of suspected nodal disease. These decisions can result in significant changes in treatment volumes when compared with a process where PET data has not been used [7]. The accurate staging of the disease using PET is known to be essential for the appropriate management of the patient [8]. To date, PET images have been applied in radiotherapy planning primarily for non-small-cell lung cancer (NSCLC) [2, 9–11]. However, studies into the use of PET images for other sites, such as oesophagus [12] and head and neck cancer [13] have also been published. All these studies have shown that the inclusion of PET data in the treatment planning process modified the target volumes in a significant number of patients. What is less clear is whether these changes are primarily due to the inclusion of previously unsuspected nodes in the target volume, as the patients had generally not had a staging PET scan in addition to the scan used for treatment planning purposes. It is well known that there may be high interobserver variability between gross tumour volumes (GTVs) defined using planning CT images. Several studies have shown that the inclusion of PET data reduces this interobserver variability [9, 11, 14]. Although this is a welcomed outcome, it does not necessarily mean that the volumes are being defined any more accurately [15]. From the above published studies, it can be seen that volumes can be delineated and treatment options changed based on the additional information. However, such an approach can only be justified if it is possible to adequately answer a number of key questions: (i) What does the molecular image represent, how specific is the functional signal and how sensitive to the disease process? (ii) What does the uptake of the radiotracer mean? (iii) How are the S28

PET data best acquired? Do the limitations imposed on the acquisition of PET/CT data impact upon clinical utility and outcomes? (iv) What criteria should be used to delineate target volumes on PET images? (v) Can the functional data acquired be synchronised to a physiological signal such as respiration? What will be the impact on volume delineation and treatment delivery? (vi) Does the incorporation of functional data change clinical outcomes by improving quality of life or by improving survival? The following sections review the current status of the answers to these questions and outline some of the work which is being undertaken to evaluate the impact of the unknowns on the introduction of PET into the radiotherapy planning process.

The PET image: its acquisition and interpretation PET tracers The most widely used PET tracer in oncology is a modified glucose molecule (fluorodeoxyglucose) labelled with radioactive fluorine (18F-FDG). This tracer is a marker of glucose metabolism within the body and is driven primarily by the expression of the glucose transporter molecule (GLUT-1) at the cell surface. When compared with CT for the diagnosis and staging of a range of cancers, it is known that PET has a significantly higher sensitivity and specificity for the detection of disease based upon the size and shape of structures in the body [16]. This is particularly marked in lung cancer, although false positive images can be caused by inflammatory or infectious processes and false negative images for a number of tumour types such as carcinoid and bronchoalveolar carcinoma. Sensitivity and specificity data are different for each tumour type and organ location. Although FDG is currently the predominant tracer used in oncology studies and its use will be the focus of this review, tracers mapping different metabolic processes may also have a role. Amino acid metabolism may prove to offer a useful non-specific index of tumour viability, especially in the brain where there are high rates of glucose metabolism in normal nervous tissue. This process can be probed with 11C-Methionine or 18F labelled tyrosine. Thymidine has also been proposed as a marker of cell proliferation based upon its incorporation into DNA [17]. A thymidine analogue, 39-deoxy-39[18F]fluorothymidine (FLT) can now be routinely manufactured and is entering clinical trials. Of growing interest is the possibility of differentiating tissue or cell types with images which provide data related to clonogen densities. Currently, tracers to image hypoxia, angiogenesis and apoptosis are under development. For hypoxia imaging, the most widely studied tracer is a misonidazole derivative (18F-MISO) together with a copper labelled compound, diacctyl-bis(N(4)-methylthiosemicarbazone (ATSM). The use of this latter compound to guide IMRT treatment planning has already been reported [18]. These and many other compounds will continue to be developed and each will need validation as an imaging and treatment guidance tool. The British Journal of Radiology, Special Issue 2006

PET/CT in radiotherapy planning

PET image interpretation How should the PET data be used in the radiotherapy treatment process and, in particular, in the delineation of the target volume? A recent editorial by Gregoire [19] highlights the issues in relation to false positive and false negative rates for detection of disease. It must be recognized that no single modality has 100% sensitivity and specificity for disease. A gold standard does thus not presently exist. The definition of treatment volumes, and hence treatments based upon a single diagnostic modality, may fail through the exclusion of diseased tissue not detected. Historical clinical trial data [20] shows that in the majority of lung cancer patients disease progression is due to recurrence local to the treatment volumes. This may be due to insufficient radiation dose being delivered to the tumour, with the prescribed dose generally being limited by the tolerance of normal tissue within the treatment field. Alternatively, it may be due to a geographical miss of the real tumour. Improved knowledge of the actual tumour size, shape and location may permit smaller treatment fields allowing increased radiation doses to be prescribed. Increasing volumes will minimize the risk of geographical misses. The delineation of volumes must be based on all the diagnostic information and knowledge available of the anatomy and physiology of disease. A new modality such as PET must be integrated into this knowledge base with decisions on treatment modification based upon probabilities of false positive and negative data within particular structures and locations [19]. Simple but logical approaches are currently being applied in clinical situations [21]. The following argument might be applied to a protocol for target volume modification in the treatment of lung cancer based upon CT data only; (a) if the PET scan is positive then volumes will be added into the treatment plan. This will result in non-neoplastic tissue being incorporated into the treatment volume on a number of occasions, but will limit the non-irradiation of neoplastic cells. The effectiveness of this strategy will be dependent on the increased sensitivity of the new modality (PET imaging) compared with CT. (b) If the PET is negative then volumes should be removed from the planned treatment volume. This approach requires a more specific PET image than can be obtained with CT. It will potentially result in neoplastic tissue being excluded from the treatment volume, but will minimize the volume of normal tissue irradiated with a therapeutic dose. The probability of a correct interpretation of a diagnostic image is dependent on a number of factors including: the spatial resolution of the modality; the signal to noise ratio of the image formation process; the affinity of the tracer or contrast material for neoplastic tissues; confounding processes such as inflammation, as well as the criteria used for image interpretation; and last, but not least, the interpretive skills of the reporting clinician and radiation oncologist.

PET and PET/CT images for treatment planning Images for radiotherapy treatment planning (RTP) purposes must be acquired differently from those for use in diagnosis and staging of patients. The standard The British Journal of Radiology, Special Issue 2006

imaging technique used in radiotherapy planning is X-ray CT as it provides both good anatomical detail for defining target volumes and the electron density data required for dose calculations. Any other images for use in RTP normally have to be registered to a planning CT scan. Images for use in RTP must be acquired with the patient in the treatment position, on a flat couch top and with the use of appropriate immobilization devices. Patient set-up must be carried out by therapy radiographers. The initial set-up of the patient in the treatment position may take some time. The internal target volume defined from the images must be related to the surface of the patient for subsequent set-up of the patient for treatment simulation and delivery. This is achieved using markers which are visible in the image to be located at skin positions which are semi-permanently marked. There are a number of different ways in which PET images can be used for treatment planning purposes. PET images could be available purely as a diagnostic aid; PET and CT images from separate scanners can be registered in software; PET-CT images from a combined scanner could be registered to a planning CT scan; or a planning PET-CT scan could be carried out on a combined scanner. Initial studies [22] of the use of PET images in treatment planning used PET and CT images which had been acquired on different scanners and registered afterwards. In general, these studies had been acquired with the patient in the treatment position. Software fusion of the PET and CT images has the advantage that the acquisition of the PET data does not necessarily need to be conducted in the treatment position and might be obtained from the diagnostic scan. The disadvantage is that the simulation scans and the PET scans will be acquired at different times, possibly on different couch tops (flat top compared with a concave top) and potentially with the patient in different positions. In a review of ‘‘organ motion and its management’’ Langen and Jones [23] refer to such position-related organ motion and the potential problems associated with imaging and treating with the patient in different positions. These can be caused in part by differences in deflection of the imaging and treatment couches as well as setup errors. If images are acquired in this way, for sites other than rigid organs such as the skull and brain, there will be a requirement for non-rigid body data transforms to register the two datasets [24]. Such image registration techniques are known to result in residual misalignment errors, the clinical significance of which remain unknown. Another option is to use the CT scan obtained for the attenuation correction of the PET images as the basis for registration with a separate radiotherapy planning CT scan. This has the advantage that the image registration is CT to CT and hence potentially has smaller errors associated with it. More recently, a number of studies have been published which have used combined PET-CT scanners [14]. Although the use of PET-CT scanners to provide a hardware solution to image registration has the advantage of reducing uncertainties in patient positioning, there is still an on-going debate about the efficacy of such integrated scanners [25] for radiation oncology. If a PET or PET-CT scanner is to be used directly for acquiring radiotherapy planning scans then a number of S29


issues need to be considered. The introduction of any new imaging modality for RTP requires the adaptation of scanning procedures for use with that modality. Close liaison between Oncology and Nuclear Medicine staff will be required to develop these protocols. Acquiring the PET/CT dataset in the treatment position requires the direct involvement of radiotherapy treatment radiographers to ensure that protocols are consistent across the planning and treatment processes. Set-up of the patient for treatment is a time consuming process and, if undertaken for PET scanning, it will be necessary to develop procedures that minimize contact between the therapy radiographers and the patient, after the patient has been injected with the radioactive tracer, in order to minimize radiation dose to staff [26, 27]. The patient will have to be positioned so that they fit through the bore of the PET-CT scanner. External wall or floor mounted positioning lasers may need to be fitted in the PET-CT scanning room. PET scans take considerably longer than CT scans, so the patient will need to maintain the treatment position for longer than for CT planning scans, simulation or treatment. Decisions will have to be made about whether dedicated PET-CT planning scans are going to be acquired, or whether whole body diagnostic scans are to be acquired in the treatment position and used for both staging and RTP. A potentially significant number of these patients will, however, not progress to radiotherapy treatment due to the outcome of the staging scan. PET scanning times are relatively long, depending on the acquisition protocol, and patient compliance may be an issue in terms of maintaining the position throughout the scan. The time taken to acquire the PET component of a PET-CT dataset depends on the number of fields of view required. A planning PET scan might require, for example in the case of the lung, two fields of view resulting in a minimum scan time of 8–10 min on current systems. Alternatively, the process could be integrated into the diagnostic work-up to preclude the need for a second radiotherapy planning study. In this scenario, the patient would need to maintain the treatment position for 30–35 min.

Target volume delineation for radiotherapy planning The radiotherapy planning process using CT data is based upon the definition of a number of volumes, usually defined in sequence and with each larger than the former. These volumes are defined in ICRU 50 and ICRU 62 [28, 29]. The first region defines a gross tumour volume (GTV), which is derived from identified disease locations with the CT or simulator images. This volume is then expanded, by the oncologist, to obtain a clinical target volume (CTV) using additional clinical information and experience. This CTV is then further modified by the addition of a variable margin to take account of patient and organ movement, and setup errors to provide the planning target volume (PTV). Of these volumes, only the GTV might be rigorously defined from structures within an image. The expansion of the GTV to CTV to the final PTV requires the subjective input of oncologists and information about organ movement and S30

set-up uncertainties. Whilst the delineation of volumes on CT images may present some difficulties, the definition of volumes on PET images is likely to be more problematic due to the poorer resolution and higher noise levels. There are a number of different options for defining lesion volumes using PET images. Volumes can be defined manually by the oncologist, as for conventional CT planning; alternatively, automatic outlining methods can be used. Thresholding is the most widely used method to determine volumes automatically from PET images [30– 32]. It is well known in nuclear medicine that the selection of the threshold depends on the lesion size, shape and contrast. Figure 1 illustrates the effect of different thresholding and analysis techniques on segmented volumes. It has been shown that when a priori knowledge of the size, shape and contrast of the lesion exist, threshold levels can be specified for accurate evaluation of the tumour volume. Most algorithms have been derived from the study of spheres of various sizes with differing levels of background signal. All groups have recognized the need to take account of the lesion to background ratio in defining volumes. All have recognized that the volume calculation is in error in small volumes (typically ,4–8 ml depending on the resolution of the system). However, little work has been published on the effect of noise levels within the images or on the validation of volume calculations from non-spherical objects. There has also been little validation of the use of PET imaging to delineate the GTV through the use of surgical samples [33, 34]. A number of groups have proceeded to implement volume definitions into the radiotherapy planning process based on this phantom validation work and used contours based upon the maximum in the lesion of interest with values ranging from 30% to 50%. Paulino et al [35] used a threshold of 50% in head and neck tumours and reported that 75% of GTV definitions were smaller using the PET image when compared with the CT derived volumes. Such a consistent finding would suggest an inappropriate choice of contour value. Other factors such as lesion volumes, reconstruction software, tracer sensitivity and specificity could all contribute to these findings. Other groups have chosen to use a threshold based upon the standardized uptake values (SUV) rather than upon a threshold based on maximum intensity [36]. What is not in doubt from these publications is that different GTVs have been obtained from the PET and the CT data and that treatment volumes have been modified. Bradley et al [7] reported a change in volume in 50% of patients based upon GTVs. The group continued to add margins to these data to provide the final PTV. In the chest for lymphoma, Lee et al [37] based volumes upon manual contours derived from PET images which had been thresholded to provide a PET defined lung volume which ‘‘matched’’ the CT lung volume. The functional image provides data beyond an enhanced tracer uptake used to define a tumour volume. The validation studies referenced above and the subsequent application of the findings are based upon the assumption of a uniform distribution of tracer within the visualized volume. This assumption is almost certainly not valid, especially in larger tumours. The PET image The British Journal of Radiology, Special Issue 2006


Figure 1. An example of the variations in segmented volumes for a lung lesion using different thresholding and analysis techniques (courtesy Mathew Hatt, University of Brest, France).

represents a metabolic process and different PET tracers are capable of providing evidence of different metabolic processes, as previously discussed. For each tracer an increase in uptake within the image equates to an increase in the functional process in that volume. Pilot studies [18] are being reported in which these additional data are being used to modify radiotherapy treatments to provide enhanced doses to specific areas, e.g. increased glucose metabolism in FDG images. PET images are being further segmented to inform an IMRT treatment plan [38]. The dose distribution has been modified on the basis of the regional distribution of FDG uptake within the tumour with a higher dose given to higher uptake volumes. Alternatively, multiple tracers have been used to identify metabolic areas with FDG and hypoxic areas with the potential for doses to be modified in the light of these different metabolic processes. It will be necessary to carefully validate these modifications to treatment protocols before adopting functional images into the routine radiotherapy process. In defining target volumes in the thorax and upper abdomen, respiratory motion must be taken into account. As described above, margins are added to the GTV to take this into account. There is also considerable interest in applying respiratory gating to both the planning PETCT scans and the radiotherapy treatment. The British Journal of Radiology, Special Issue 2006

Respiratory gating for PET/CT data in radiotherapy planning In defining a volume on the PET image it is essential to understand the acquisition process, especially in relation to physiological motion. For studies in the head and neck, immobilization techniques should eliminate potential gross movements and the alignment of PET and CT data from a combined PET/CT study should be ‘‘exact’’. For studies of the chest and abdomen, the use of immobilization devices will help eliminate gross patient movement. However, there remains a significant discrepancy in the way that physiological movement due to respiration and heartbeat impacts upon the PET and the CT images. The increasing use of high-speed, multislice CT scanners enables images of the chest to be acquired in periods of time which are short compared with the respiratory cycle and effectively provide a ‘‘snapshot’’ of the lungs in time. This technique can be further controlled by the use of breath-hold techniques. However, this situation does not pertain for the acquisition of the PET data. Data are acquired over a number of minutes and represent a time averaged distribution based upon the dwell time of the activity at any point in space during the study. Thus, objects which do not move with time will see no degradation in activity concentration, whereas those objects which move S31


significantly will exhibit a reduced activity concentration due to the distribution of activity throughout a larger apparent volume. It could therefore be argued that these PET images inherently include a margin for physiological motion and that no further allowance should be made in the definition of the PTV [39]. Other factors such as gross movement and repositioning errors will remain. This, however, is not the only degrading factor. The presence of respiratory motion introduces inaccuracies into the reconstructed images as a result of misregistration between PET and CT acquisitions [40, 41]. Since with these hybrid scanners, the CT maps are also used for the correction of the attenuation effects in the emission data, an extra inaccuracy may be introduced by using non-perfectly aligned CT and PET datasets as a result of the respiratory motion [42]. The definition of a PTV clearly remains a complex task and techniques for the definition of a GTV may have a limited impact on the final definition of a PTV, especially where significant physiological motion is known to occur. These uncertainties have led to the investigation of diagnostic and treatment methodologies which measure physiological motion and incorporate the data into the treatment plan and the delivery system. The solutions that have been proposed to date for taking into account the effects of respiratory motion concentrate on the acquisition of respiration synchronised PET and CT datasets. The use of breath hold protocols has been used as a means of improving registration between the PET and CT [43]. However, these will not aid in the delineation of target volumes as the radiotherapy treatment will be delivered over a few minutes. There is a lot of interest in the use of respiratory gating for both the PET-CT image acquisition and the treatment. Several studies have been carried out to investigate the feasibility of respiratory gating of PET of the upper chest and abdomen [44–46] and also to quantify the impact of respiratory motion on the underestimation of lesion activity [47]. Different detector systems have been proposed, including a transducer or an impedance electrocardiograph (ECG) monitor measuring changes in abdominal or thoracic circumference [48], a thermistor measuring the temperature of circulating air during patient respiration [45, 46], a spirometer measuring respiratory flow [49], the Varian Real Time position management (RPM, Varian Medical Systems, Palo Alto, CA) [47, 48], or the Polaris system tracking the displacement of infrared reflective markers in the patient chest [47]. An alternative approach to gating is to use an image derived respiratory signal through the acquisition of dynamic datasets [51, 52] or list mode data. One such respiratory correlated approach used a point source of 18 F-FDG attached to the patient’s skin to track respiratory motion [50]. Identifying the frames in which the point source fell within an operator defined region of interest (ROI) allowed PET images corresponding to different points within the respiratory cycle to be created. It was demonstrated that this technique produced similar results to gating. Another approach which uses time activity curves generated from a ROI drawn over a moving object in the image to recover the breathing frequency is currently undergoing clinical validation [49]. The advantage of these techniques is that the data S32

may be retrospectively reconstructed for any breathing phase or amplitude. Irrespective of the gating methodology implemented, the emission data acquired in each of the temporally gated frames is reasonably free of respiration-produced inaccuracies. However, the resulting individual frame images are of reduced resolution, as well as overall quality, as they contain only a fraction of the counts available throughout a PET acquisition [51]. Some groups have attempted to deal with this problem by acquiring gated data in 3D mode [45]. The need, therefore, exists for the development of correction methodologies making use of the gated datasets, in order to obtain respiration free PET images using all available data throughout a standard respiration average PET acquisition. This approach will also remove the need currently existing in terms of significantly increasing the time (over a factor of 3) of gated PET acquisitions in order to compensate for the presence of reduced statistics in the final reconstructed images. Very limited work is currently available in this domain. First, an emission driven solution through the combination of respiratory synchronised emission datasets and an iterative reconstruction algorithm can be envisaged, in a similar fashion to the methodology that has been previously suggested for SPECT cardiac imaging applications [52, 53]. The second option is based on a realignment methodology to ‘‘bring’’ all of the respiratory synchronised PET datasets to a particular phase in the respiratory cycle. This methodology is potentially applicable to both image and raw data domains, deriving the transformation parameters from the corresponding respiratory motion synchronised CT frames [48, 54].

The impact of PET imaging on patient outcomes There remains little evidence of the impact of the integration of PET into radiotherapy treatment planning on patient outcome. All published series’ are small and there are no studies which address the issue of patient outcomes and quality of life. These studies will be difficult to perform and the results seriously compromised by the fact that past published data have been based on patients who did not have the benefit of PET imaging as part of their disease staging [8]. There are a small number of studies which have investigated the potential impact that PET integration would have had on treatment volumes and normal tissue doses. Lee et al [37] considered the impact of PET directed treatment fields in thoracic lymphoma and demonstrated that if residual masses were reduced in line with PET positive masses then normal tissue doses could be significantly reduced and doses escalated without detriment to active disease areas. This approach should see a reduction in secondary tumour development in normal tissue subsequent to curative treatment [55]. Nioutsikou et al [56] have published an evaluation of the impact of including lung functional data into the radiotherapy planning process with a view to optimizing fields to exclude functional lung volume wherever possible. Whilst this is not directly related to PET volume definitions, it provides a methodology which The British Journal of Radiology, Special Issue 2006


could be applied to enable normal tissue toxicity to be evaluated when treatment plans are modified, thus building upon well characterized tissue toxicity data. Miften et al [57] had previously published a similar study using lung functional data to modify treatment plans. Van Der Wel et al [60] undertook a modelling study of the toxicity effects of the introduction of PET into stage N2-N3M0 NSCLC treatments. The group defined a therapeutic ratio and again demonstrated that in a group of patients tumour doses could be escalated at no detriment to normal tissues. The model was based rigidly upon ICRU guidelines for treatment volume definitions [28] with a rigid protocol for the introduction of PET data. All volumes used were based on the CT structures, with the CT defined structure removed if the PET scan was negative and CT structures added if the PET scan was deemed to show a positive uptake. No attempt was made to define volumes based on the segmentation of the PET data. It would appear essential that these methodologies are used to inform the construction of clinical trials aimed at rigorously assessing patient outcomes and the associated cost benefits of such developments.

Conclusions and future work The introduction of integrated PET/CT scanners has enabled the collection of inherently fused functional and anatomical data. The modification of the data acquisition process to permit the use of radiotherapy treatment planning couches and immobilization devices is well developed, and systems are being developed to provide larger fields of view and patient manipulation space. The integration of external laser positioning systems is essential with the integration of coordinate data into the PET/CT dataset. First generation PET/CT scanners had poorly developed standards for image formats and presented significant problems in relation to data transfer to radiotherapy planning systems and data validation. However, this has largely been solved with RTP systems capable of the receipt and direct utilization of PET and CT datasets. The process of integrated data acquisition requires the skills of therapy radiographers to ensure continuity through the treatment pathway. Experience has shown that if the PET/CT data are to be used directly in treatment planning, then the therapy radiographers will acquire a radiation dose from the close proximity work with the radioactive patient. This dose is not insignificant and is not readily avoided, unless the PET data are not acquired in the exact treatment position and introduced into the treatment planning process through software image fusion. The utilization of the PET functional data still remains a challenge. The unique nature of the PET/CT dataset still presents many opportunities to use the two data types to optimize the definition of tumour volumes as the first stage of the radiotherapy treatment plan. At present, most centres have relied on the assessment of volumes from PET images based upon simple or adaptive thresholding methods validated with reference to spherical objects. These methods are relatively crude and do not take account of respiratory and cardiac The British Journal of Radiology, Special Issue 2006

motion, which effectively modify volumes in the lung and abdomen. There is also the potential to utilize the CT data to enhance the PET image to assist in automatic segmentation. Physiological motion may have a significant impact upon the reconstructed PET image as the misalignments between the PET and CT datasets will result in an incorrect attenuation correction being applied which, in the chest in particular, will lead to incorrect tracer distributions [59]. This limitation has led to the active investigation of the use of respiratory gating in the acquisition of both the CT and PET datasets. The successful implementation of this technology has the potential to fully map the motion of tissue within the treatment fields and to provide aligned, correctly reconstructed, PET functional data. Commercial systems are already available, although their implementation is not straightforward. It should be noted that respiratory gated CT will potentially lead to a significantly increased CT dose to the patient. The gating of the PET acquisition will lead to a reduction in the signal to noise ratio in the images spanning the respiratory cycle as it will not be possible to increase the injected dose or increase the scanning times to compensate for the reduction in image acquisition time. It has been shown that this will lead to increased errors in volume determinations, which may outweigh the gains of better localization through gating. It must be remembered that the aim is to segment the lung based upon periods of equal volume and not to segment the respiratory cycle into a fixed number of time bins. Commercial systems are currently based on time or phase segmentation as the trigger systems do not produce lung volume related signals. Systems based upon chest expansion, such as a chest strain gauge, have the potential to provide a signal related to lung volume and to permit segmentation based upon phases of equal volume. For all of these methods the process must be capable of controlling the delivery of the radiation therapy to best utilize this input data. This phase of development is aimed at the more accurate definition of treatment fields linked to the ability to conform the treatment fields to these defined volumes using IMRT and associated techniques. The introduction of functional maps into the process is, in itself, a confounding variable. An accurate understanding of the meaning of the image is vital to the definition of protocols which permit the routine and reproducible use of the data. Limited sensitivity and specificity for the PET signal will mean that some PET tracers will have very limited use in radiotherapy planning for some tumour types in some organ systems. It must also be remembered that the accuracy of PET interpretation, as with CT data, will be observer dependent and will almost certainly require the formation of teams of specialists to adequately and reliably incorporate PET data into the planning process. The integration of functional data still presents the potential to better target external beam therapies. Dose escalation to hypermetabolic or radio-resistant areas may also be guided by specific functional markers. Evidence of the effectiveness of these methods in patients is still limited and at least one publication suggests that extreme caution is required if outcomes are not to be detrimental to the patient. There remains an urgent need S33


to develop well validated protocols for the interpretation and use of PET functional images in the specification of radiotherapy treatment plans. These should be tested in well constructed clinical trials which themselves have been assessed by the modelling of radiation toxicity effects associated with these technology developments.

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Localization: conventional and CT simulation G R BAKER,

BSc, MSc, MIPEM

Kent Oncology Centre, Maidstone Hospital, Maidstone, Kent ME16 9QQ, UK ABSTRACT. Recent developments in imaging and computer power have led to the ability to acquire large three dimensional data sets for target localization and complex treatment planning for radiation therapy. Conventional simulation implies the use of a machine capable of the same mechanical movements as treatment units. Images obtained from these machines are essentially two dimensional with the facility to acquire a limited number of axial slices to provide patient contours and tissue density information. The recent implementation of cone beam imaging on simulators has transformed them into three dimensional imaging devices able to produce the data required for complex treatment planning. The introduction of computed axial tomography (CT) in the 1970s was a step-change in imaging and its potential use in radiotherapy was quickly realised. However, it remained a predominantly diagnostic tool until modifications were introduced to meet the needs of radiotherapy and software was developed to perform the simulation function. The comparability of conventional and virtual simulation has been the subject of a number of studies at different disease sites. The development of different cross sectional imaging modalities such as MRI and positron emission tomography has provided additional information that can be incorporated into the simulation software by image fusion and has been shown to aid in the delineation of tumours. Challenges still remain, particularly in localizing moving structures. Fast multislice scanning protocols freeze patient and organ motion in time and space, which may lead to inaccuracy in both target delineation and the choice of margins in three dimensions. Breath holding and gated respiration techniques have been demonstrated to produce four-dimensional data sets that can be used to reduce margins or to minimize dose to normal tissue or organs at risk. Image guided radiotherapy is being developed to address the interfraction movement of both target volumes and critical normal structures. Whichever method of localization and simulation is adopted, the role of quality control is important for the overall accuracy of the patient’s treatment and must be adapted to reflect the networked nature of the process.

The development of the delivery of radiation therapy is closely related to the accuracy with which the target tumour can be located with respect to surrounding anatomical structures. In recent years, the increase in computing power and the development of refined computer graphics have resulted in the ability to perform complex treatment planning in three dimensions and to manipulate images in real time. Early simulators were machines capable of the same mechanical movements as treatment units and were used to confirm treatment set up rather than for localization [1, 2]. Simulators that were developed commercially in the 1960s had the addition of fluoroscopy that was used to set the isocentre with the aid of remotely controlled movements of the couch. Field portals adequate to encompass the target volume to be treated could also be set by remote adjustments to the field defining wires. The introduction of computed axial tomography (CT) scanning in the 1970s was a step change in the ability to define tumours in relation to normal anatomy, and over the ensuing years has been widely adopted in tumour localization. Today it may be used in conjunction with complex graphics software as a virtual simulator. However, the conventional simulator still retains its place in many radiotherapy departments S36

Received 30 June 2005 Revised 28 February 2006 Accepted 1 March 2006 DOI: 10.1259/bjr/17748030 ’ 2006 The British Institute of Radiology

for localization of some tumour sites, either as a result of lack of sufficient access to a CT scanner or for relatively simple techniques not requiring the production of a dose plan. The conventional simulator is also frequently used to verify the more complex treatment plans, producing an image corresponding to a beam’s eye view (BEV) from the treatment planning system (TPS) or by verifying the isocentre location from orthogonal films.

Brief history Mould [3] describes the development of simulation, from the use of diagnostic radiographs and skin marks in the 1950s to the introduction of virtual simulation in the 1980s. In 1973, Hounsfield and Ambrose [4, 5] published their work on computerized transverse axial tomography and the potential uses of CT in radiotherapy were quickly recognized [6]. However, access to a CT scanner was often very limited, and in many cases the scanner was not even in the same hospital as the treatment facilities. In addition, a CT scanner was principally a diagnostic tool with limitations for treatment planning imposed by the small aperture and the design of the The British Journal of Radiology, Special Issue 2006

Localization: conventional and CT simulation

couch, which frequently prevented the patient from being scanned in the treatment position. Harrison and Farmer [7] recognized the usefulness of being able to acquire a cross-sectional image of the patient in the treatment position using a simulator as a CT scanner and went on to describe the implementation of their idea using a fluorescent screen and an Isocon camera [8]. A number of other adaptations of the simulator to produce cross-sectional images were also proposed at this time [9–12]. This functionality was called Sim-CT and became standard on simulators in the 1990s, but the system had its limitations: 1. The heat capacity of the X-ray tube generally meant that only a few slices could be scanned; 2. The time taken to scan was limited to approximately one revolution per minute, which introduced motion artefacts resulting in images that were of a poorer quality than those produced on a diagnostic scanner; 3. The uncertainty in the Hounsfield units (HU), which depends on the field of view and the phantom/ patient size, a result of the beam hardening in the unfiltered X-ray beam from the simulator CT. However, the uncertainty in HU is translated into dose variation not exceeding 3% for photon beams in the range 6–18 MV [13]; 4. The relatively high dose to the patient which was shown to be approximately 10 times that delivered with a diagnostic scanner under similar conditions [14]. In spite of its limitations, the Sim-CT was a useful tool for planning in a department with limited access to a diagnostic scanner. It was a more accurate way of producing a patient outline than manual methods using callipers and flexicurves and enabled CT numbers to be converted to relative electron densities for tissue inhomogeneity corrections to be applied to a single CT slice in dose calculations. The dose distributions and monitor unit calculations showed good agreement with those obtained with diagnostic scan data [14]. In 1998, Cho et al [15] described the application of digital technology to a radiotherapy simulator in which the imaging system was replaced by a digital spot imager (DSI). The DSI consisted of an image intensifier, digital image processing, display and data transfer facilities. The images were stored during acquisition for later archiving or transfer to workstations. Simulator manufacturers now offer digital capabilities on their machines and conventional image intensifiers have been replaced by flat panel amorphous silicon (aSi) detectors. Their longevity in this application has to be proved and it is possible that the need for regular replacement may have significant revenue consequences. The most recent simulators include anatomical protocol selection, automatic correction for image distortion, last image hold, multileaf collimator (MLC) verification, a variety of image viewing and manipulation tools with annotation, image printing to film or paper, Digital Image Communications in Medicine (DICOM) export to TPS, electronic portal imaging device (EPID), record and verify, and patient management systems. The image manipulation tools enable adjustments to be made to field parameters and image quality on the last-held The British Journal of Radiology, Special Issue 2006

image, which reduces the screening time and hence patient dose compared with non-digital systems. A wide aperture (typically 90 cm) CT option is available. However, because of the restriction on gantry rotation speed, acquisition times are still slow and reconstruction time does not match that of a diagnostic scanner. In an attempt to overcome this, volume or cone beam CT (CBCT) has been developed. A number of authors describe cone beam reconstructions, based on Feldkamp’s original back projection algorithm [16], for the acquisition of volumetric data [17–19]. When first proposed, the size of the detector was a severe limitation on the reconstruction volume and, although promising results were obtained, its use in treatment planning was not realised until aSi flat panel detectors of a reasonable size became available. Commercial systems are now available. For example, the Acuity (Varian, Palo Alto, CA) with cone beam option gives a cone of 17 cm at the isocentre but with added penumbra of 1.9 cm at either end regardless of the scan length. It is therefore not appropriate to acquire a single narrow slice. A single slice takes 45 s and 675 images are acquired per rotation. Early reports (private communications, A Vinall, K Venables, 2005) suggest that the geometric performance and image quality are adequate for radiotherapy planning purposes although the images are not of diagnostic quality. The rotation time of 45 s does, however, result in significant movement artefacts. Figure 1a shows the streaking that results from the movement of bowel gas during the acquisition of a CBCT scan compared with a CT planning scan.

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(b) Figure 1. (a) Movement artefacts on an axial slice of a CBCT scan as a result of movement of bowel gas. (b) An axial slice from a planning CT of the pelvis for comparison. (Courtesy of Varian Medical Systems, Palo Alto, CA and Memorial SloanKettering Cancer Centre).

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Figure 2. Movement artefacts on an axial slice from a CBCT acquired during normal breathing. (Courtesy of Varian Medical Systems and Hirslanden Klinik, Aarau).

Figure 2 shows similar streaking in the soft tissue around lungs in a CBCT taken during normal breathing. As with the single slice option on the simulator, there seem to be problems with the HU values both in accuracy compared with the calibration and reproducibility on a day-to-day basis. Slice thicknesses of 1–5 mm are available. Reconstruction times vary with the slice thickness and are in the order of 90 s. There is no standard way of quoting doses for these scans. Computed tomography dose index (CTDIw) is a measure of the dose from a CT scan, weighted between the centre and the surface to give an average value across the section. A CTDIw/ 810 mAs value of 15 mGy has been measured for a 10 cm scan length collimated to 13.8 cm (15 pulse s21, pulse length 15 ms, 80 mA, 125 kV, 45 s rotation). Setting the scan length to 1 cm in clinical mode gave 54 mGy/ 810 mAs with the same exposure factors. This compares with the national reference dose of 20 mGy for a multislice scanner [20].

CT simulation The alternative to using the simulator and CBCT to acquire a volume data set of the patient in the treatment position was to modify CT scanners to meet the needs of radiotherapy and add software to perform the simulation function. With the rapid development of computer technology, enabling fast reconstruction of images in three dimensions, the true value of the enormous quantity of data acquired by a CT scanner and its use in radiotherapy planning was recognized. The development of the concept of the beam’s eye view (BEV) into the transmission image from CT scans that would result from any beam orientation paved the way to producing images from CT data that correspond to conventional simulator films [21–23]. These digitally reconstructed radiographs (DRRs) could be overlaid with the outlines of anatomic structures, field shapes S38

and cross wires, and hence could display images similar to simulator radiographs. However, the spatial resolution of DRRs is limited by the voxel size of the CT scans and cannot match that of a simulator radiograph taken with a small focal spot and a short exposure. Even in the early implementation of this process the reconstruction time of the DRRs was reasonable, being in the region of 10 s for a 50 slice study. However, studies were limited by the specification of the CT scanner. The acquisition of a single slice might take 2–3 s with a delay between scans required for repositioning of the scanner and tubes with low heat capacity needed cooling time during the scan [24]. Early critical analysis of the CT simulation process highlighted the areas for improvement [25]. These included the limitations imposed on both treatment technique and the size of the patient by the aperture of the scanner (normally 70 cm), the time required for CT data acquisition and transfer from the scanner to the planning system, time required for outlining and contouring target volume and critical structures and the inconsistent accuracy of portal marking on the patient’s skin. Complete field ports were marked on the patient’s skin in most cases and novel devices for doing this constituted an important part of the virtual simulation process reported. [26, 27]. These drawbacks have now largely been overcome. Multislice helical scanning, with high heat capacity CT tubes, has reduced the time required to acquire a CT data set of 100 slices to a matter of seconds. Wide bore scanners have removed most of the constraints of patient size and technique. Increased computing capacity and speed allows for real time reconstruction of the slice images at the scanner and real time manipulation of images in the virtual simulation software. In addition, the DICOM protocol facilitates fast transfer of image data between systems.

Current practice Conformal radiotherapy (CRT) is now accepted best practice for a number of treatment sites, having the advantages of sparing normal tissue and providing the opportunity for dose escalation. Intensity-modulated radiotherapy (IMRT) is the ultimate expression of this, but successful implementation of CRT and IMRT cannot be achieved without three-dimensional information on the location and extent of the target volume and the position of adjacent organs at risk (OAR). The threedimensionality of virtual simulation is essential to visualize the coverage of the target volume and the avoidance of OARs in the highly complex treatment plans required for CRT and IMRT. For some sites, such as the lung where the relative position of the target and OARs varies with time, this fourth dimension needs to be taken into account. Sherouse et al [28] introduced the term virtual simulation in 1987 to describe the process of using computer aided design and digitally reconstructed radiographs to replace the process of physical simulation. The process of virtual simulation has been described in detail by Aird and Conway [29] who also gave examples of its application to a number of different sites. The British Journal of Radiology, Special Issue 2006


The specification of a CT simulator The fundamental requirements of a CT simulator are a CT scanner with a flat couch, positioning lasers and virtual simulation software.

fixed or move in a vertical plane. The sagittal laser must be able to move laterally to account for lack of lateral movement on the CT couch. These lasers move under computer control to define the isocentre for the treatment plan in terms of shifts from the reference marks.

CT scanner

Virtual simulation software

Advances in the design and capabilities of CT scanners have modified the specifications given by Aird and Conway [29]. Multislice scanners enable very fast scanning times, even for the large studies, with narrow slice thicknesses required for the production of good DRRs. High heat capacity anodes are required for the large datasets that are frequently required for treatment planning applications. One manufacturer (Siemens Medical, Erlangen, Germany) has introduced a new design of directly cooled anode that should eliminate delays due to anode heating and enable fast acquisition of scans with the large number of narrow slices required for good DRRs. Three manufacturers now produce wide aperture (85 cm) scanners designed for radiotherapy applications. In two, the scanned field of view (SFOV) is 60 cm with an extended reconstructed FOV of 85 cm. It should be noted that in the extended reconstructed FOV the HU numbers may not be consistent with the SFOV. In reality, it is unlikely that the uncertainty in HU translates into a dose discrepancy of more that 1–2% in the target. The third manufacturer claims a true SFOV of 85 cm.

The virtual simulation software may either be part of a treatment planning system or may be a stand-alone system. If the latter is chosen, it is essential that connectivity is easily established with the treatment planning system for dose calculation. Since the introduction of DICOM-RT this connectivity is more readily achievable, but the user must be aware that not all manufacturers interpret the standard in the same way and there are frequently hidden licensing issues associated with the connectivity. Essential features of virtual simulation software include automatic contouring of body outlines and semi-automatic contouring of other structures and critical organs such as spinal cord, kidneys and lungs. Particular attention should be paid to treatment of bifurcating structures. Contouring tools should be simple to use and interpolation between nonadjacent slices, with correction as necessary, should be provided to speed the contouring process. The ability to contour in three dimensions, i.e. in sagittal and coronal as well as axial sections, is particularly helpful. Figure 3 shows how three single contours in orthogonal planes produce a three dimensional structure. This functionality can considerably reduce the time taken to outline structures. The shape of the contours can be modified on any slice as necessary. Similar interpolation tools should be available for target volume delineation and true three-dimensional volume margin growth with different margin widths in different directions. Threedimensional display systems are an essential feature of

Positioning lasers A system of three lasers for the accurate positioning and alignment of the patient is required. The lateral lasers may be wall or frame mounted, and may be either

Figure 3. A single contour in axial sagittal and coronal planes defines a three dimensional target in Prosoma. (Courtesy of Oncology Systems Limited, Shrewsbury, UK and Medcom, Darmstadt, Germany). The British Journal of Radiology, Special Issue 2006

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any virtual simulation software. It should be possible to display axial, sagittal and coronal sections on the same screen and relate each section to the others, and to visualize the DRRs in the same window. An Observer’s Eye View, with the patient on the couch and the floor and gantry angles depicted, is an aid to patient setup, as is a light-field displayed on the patient’s skin related to skin marks or tattoos. Anti-collision software avoids planning a treatment field which it is physically impossible to reproduce in the treatment room. There are many different ways of rendering the target volume and OARs, but they should be unambiguous and should be rendered in three-dimensions so that coverage can be checked from all aspects. Optimization of MLC leaf positions and collimator angle should be available but adjustable by the planner. For treatment planning where a full dose distribution will not be calculated, a particularly useful feature is the calculation of the equivalent square of an irregular field, the parameter required for simple dose calculations. Increasingly, oncologists are using a number of other imaging modalities such as MRI (see Khoo and Joon in this issue) and positron emission tomography (PET) (see Jarritt et al in this issue) to help in determining target volumes. Most virtual simulation packages include an image fusion function enabling registration of two datasets of the same or different modalities, CT/CT, CT/MRI, CT/PET. Image registration and fusion may be achieved in a number of different ways, both manual and automatic (see Kessler in this issue). Irrespective of the algorithm, there is a variety of display modes to assist in performing and viewing the fusion, some of which are shown in Figure 4. Figure 4a shows the two data sets (MR and CT) fused with information from both sets displayed in the same window. The image can be ‘‘faded’’ between the two showing 100% of the primary data set (CT in this case) through to 100% of the secondary data set (MRI in this example). Figure 4b shows a split screen, with two quadrants displaying the CT data and two showing the MRI data. The point of intersection can be moved around the image to display the intersection at any position on the image. This will assist in delineating the structures using information from both data sets. Figure 4c shows a split screen with the secondary data set fused with the primary in the centre of the image and the primary image on either side. Contours outlining the target or OARs can be drawn on either data set or on the fused images in any of these display modes. These three screens show the fused images in the top three windows and the secondary data set in the lower windows. Figure 4d shows the region of discrepancy between the two fused data sets, in this case two CT studies, as areas of enhancement on the image. Improved localization of a brain tumour when CT and MRI data sets are fused compared with localization on CT alone for treatment planning is demonstrated in Figure 5.

Comparison of conventional and virtual simulation Conventional and virtual simulation approach the task of localizing the target volume for treatment planning in very different ways, which may result in significantly S40

different treatments. Realisation of the steps performed to provide the data to a treatment planning system is compared for the two modalities in Table 1. In comparing the two methods of simulation, the first question that arises is whether the two are comparable in terms of accuracy of the treatment set up. There are a number of studies addressing this question for different treatment sites. Bollet et al [30] showed that in a series of 20 patients who were CT scanned and had conventional simulation, the precision of set up evaluations using DRRs was similar to that using simulator films in conformal prostate treatments. They also considered whether errors were introduced at the simulation stage and found a statistically significant systematic error between DRRs and simulator, in both the craniocaudal direction and the anteroposterior direction. In another study of prostate patients Valicenti et al [31] showed that there was no statistically significant reduction in treatment setup error if patients have physical simulation following virtual simulation and concluded that physical simulation may be omitted if virtual simulation is available. In a study of 86 patients undergoing palliative radiotherapy for lung cancer using parallel opposed fields, McJury et al [32] found that setup errors were comparable between the group planned by virtual simulation and that planned using conventional simulation. Similar results are reported at different treatment sites [33–35]. In a detailed study of setup errors in 39 patients undergoing CT planned radiotherapy for lung cancer, de Boer et al [36] concluded that the setup errors introduced at simulation, which become systematic errors if the simulator film is used as the reference image, were comparable with systematic errors at the treatment unit. Hence, omission of the simulation stage would reduce systematic errors on treatment. This conclusion supported a similar result for prostate patients [37]. In comparing the two methods of simulation, studies have shown that the target volumes and field sizes are smaller for virtual than conventional simulation in lung cancer with the associated reduction in irradiation of normal tissue [32, 38]. Smaller field sizes have also been reported for maxillary cancer with a corresponding reduction in long-term side effects [39]. One of the perceived advantages of virtual simulation is the improved coverage of the gross tumour volume (GTV) and the avoidance of OARs as a result of better visualization of soft tissue structures on a CT scan compared with a simulator image, particularly if shielded by bone. This is aided by software functions that remove overlying structures, giving better definition of the region of interest. A study comparing conventional and virtual simulation in the treatment planning of malignant lymphoma showed incomplete coverage of the spleen and spleen hilus in 5 of 15 and 6 of 15 patients, respectively, on conventional simulation and incomplete coverage of the right and left hilus in 4 of 15 and 1 of 15 patients, respectively. In addition, the left kidney was inadequately shielded in 6 of 15 of the conventionally planned patients [40]. Similar improvements in target coverage and OAR avoidance are reported for other anatomical sites [41–44]. Improved visualization of soft tissue structures may bring to light hitherto unsuspected pathology. Mehta The British Journal of Radiology, Special Issue 2006


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(b) Figure 4. (a) Fusion of MRI and CT data sets, fused images in the top windows and MRI images below. (b) A split screen showing fusion between CT and MRI data sets in quadrants. (Continued)


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Figure 4. (Cont.) (c) An alternative split screen representation of fusion between CT and MRI data sets. (d) Areas of mismatch between two CT data sets displayed as image enhancement. (Courtesy of OSL and Medcom).

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Figure 5. Improved localization of brain tumour using fused CT and MRI data sets. (Courtesy of OSL and Medcom).

and Goffinet [45] reported 17 unsuspected abnormalities in 153 scans (11%) obtained for treatment planning for patients referred for irradiation of the breast or chest wall. Of these, four represented disease that altered the treatment plan.

Working practices The introduction of CT simulation has had a considerable impact on working practices in radiotherapy departments.

Oncologist attendance The most notable change is that an oncologist is not required to be present during the scanning process. This releases the planning schedule from reliance on the oncologist’s timetable, and the oncologists are free to undertake volume definition at a time convenient to them.

Time A number of centres have reported on the different time allocation between conventional and virtual simulation [25, 28, 35]. Experience at the Kent Oncology Centre has shown that there is little difference in the total time needed for localization between the two modalities for the planning radiographers. With three radiographers in the scanning suite, 20 min appointments are adequate for most patients. Patients undergoing planning for breast radiotherapy are usually allocated 30 min because of the complex immobilization and positioning required with a narrow aperture scanner. These times are shorter than conventional simulation (30 min and 45 min, respectively), but more time is spent in manipulating the acquired data in the virtual simulation software. This includes the registering of reference marks and the production of DRRs for palliative patients, and outlining of target volumes and OARs for radical patients. Reduced simulation time for the patient leads to improved patient compliance, resulting in fewer problems from movement during scanning.

Table 1. Comparison of localization with CT and conventional simulation Function

Conventional simulation

Virtual simulation

Patient alignment Reference point definition Localization Definition of target and organs at risk

Room lasers Skin markers Fluoroscopy Drawing on plane films

Isocentre Field definition Patient outline Isocentre compared with reference point Treatment verification

From simulator scales or film From simulator scales or film Manual/optical/single slice on Sim CT Shifts measured on film Plane films

Room lasers Skin markers CT scan Contouring on original or reconstructed slices DRR from CT Virtual Sim Axial slice Calculated from Virtual Sim data DRRs

DRR, digitally reconstructed radiograph.


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Reference marks In conventional simulation, using fluoroscopy for localization of the target volume, the isocentre can usually be established and marked at the time of simulation. In CT simulation, a reference point is chosen at the scanning session and the eventual isocentre is defined by movements of the couch from the reference point. If virtual simulation of palliative patients is undertaken with the patient remaining on the couch, the isocentre can be marked immediately from the couch movements indicated.

Verification It has already been shown that to verify a plan on a conventional simulator after virtual simulation is not only unnecessary, but it could also be a source of systematic errors. However, treatment verification is still required and is of greater importance because of the use of reference marks. Verification takes place on the treatment unit with the electronic portal imaging system. The portal images acquired are then compared with the DRRs produced by the TPS or the virtual simulation software. For complex plans, this may require an extra treatment slot to allow time for the detailed comparison of portal images and DRRs before treatment.

Advantages and disadvantages of conventional and CT simulation The advantages and disadvantages of conventional and CT simulation are summarized in Tables 2 and 3. The availability of a three-dimensional dataset for all patients has some unexpected benefits. The increased information available may demonstrate previously unsuspected disease that may influence patient management. In palliative patients the extent of bone destruction from osteolytic lesions is easier to visualize on a CT scan than on a simulator film (Figure 6) and the use of software functions to remove overlying structures and display images optimized for different tissue types enables quicker localization of the disease. In breast planning, cardiac and lung volumes are more clearly

demonstrated and therefore the fields can be adjusted or shielding employed accordingly. One disadvantage of CT simulation is the increased patient dose. Doses for CT scanners are quoted as CTDIw with values in the region of 20 mGy. This dose is delivered to regions of normal healthy tissue as well as the tumour volume. Manufacturers of CT scanners provide various methods to reduce the total dose to the patient, taking account of the different dimensions of the patient at different levels and modulating the exposure in response to the detector measurements. Some challenges still remain. Respiratory motion can affect the position of lung tumours and their relationship to OARs. Fast scanning protocols freeze patient and organ motion giving a snapshot view in time and space which may lead to inaccuracy in target delineation and choice of margins in three dimensions. Imaging techniques to overcome this drawback are an area of active investigation. The conventional method of treatment planning for lung tumours is to use fluoroscopic imaging to determine the maximum migration of the tumour during respiration and adopt large margins around the CTV to ensure that the target remains in the high dose region throughout the breathing cycle. A similar philosophy can be adopted by performing scans at deep inhale and deep exhale [46]. However, a number of other techniques have been suggested involving breath holding and respiratory gating techniques [47]. Deep inspiration breath hold (DIBH) increases the lung volume relative to normal breathing and hence the total volume of lung irradiated will be reduced using this technique [48]. In some patients, DIBH may displace the tumour away from OARs [49], which has the potential for dose escalation to the target for the same level of toxicity to OARs. Gated respiration techniques may either be active or passive. In active breathing control (ABC), the patient is prevented from breathing at a given part of the respiratory cycle during which the scan is performed and subsequent treatment takes place. By acquiring a number of scans at different parts of the breathing cycle, motion of the organ in three-dimensions can be demonstrated. Passive techniques allow the patient to breathe normally and a surrogate for the respiratory induced motion, such as the movement of the anterior chest wall, is monitored. Images obtained from CT scans are sorted according to respiratory phase to produce a 4D CT data set [50–52].

Table 2. Advantages and disadvantages of CT simulation Advantages

Disadvantages

Three-dimensional dataset available, resulting in better visualization of tumour and nodal involvement, leads to reduction in side effects Reduced simulation time leads to improved patient compliance

Organ motion not visualized

One fewer patient visit during planning Oncologist not required during scanning Reduced transfer inaccuracies by omitting conventional simulator verification Can simulate non-coplanar fields

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Repeat scan required for changes in patient set-up/shape/size during treatment Palliative patients may spend longer in department between scanning and treatment Transfer of verification to treatment unit may require extra treatment slot Some patients/techniques may not be suitable for small aperture scanners (availability of wide aperture scanners should eliminate this problem) Data storage Higher patient doses


Localization: conventional and CT simulation Table 3. Advantages and disadvantages of conventional simulation Advantages

Disadvantages

Fluoroscopy gives idea of organ motion

Difficult to visualize some tumours, especially if overlaid by bone (e.g. mediastinal lesions) Limited three-dimensional information, even with CT option. Therefore cannot plan conformal or IMRT (cone beam may improve this) Two patient appointments required, localization and verification Difficult or impossible to simulate non-coplanar treatment fields

High spatial resolution

Field visualization on patients skin IMRT, intensity-modulated radiotherapy.

Breath hold and ABC techniques both require the cooperation of the patient and are therefore not appropriate for all patients. Some verbal or visual coaching helps to maintain regular breathing. An alternative approach to the problem of organ motion is suggested by Murphy [53] who describes the real-time tracking of moving organs. Tracking respiratory motion is a complex procedure as it involves fast movement of organs relative to each other. For real-time tracking to be successful, the system must be able to locate the target, predict the motion to account for any time delays in repositioning the beam and adapt the treatment plan to allow for the change in relative positions of target and OARs. Although respiratory motion appears fairly regular, there are changes in amplitude and period from one cycle to the next which make prediction complicated. Murphy discusses two ways of predicting respiratory movement, by developing a mathematical model and by using an empirical algorithm that is based on measurements of previous breathing cycles. The technical challenges of fast response times to organ motion in continuous real time tracking are presented, but Murphy suggests that in the future it should be possible to treat lung tumours in some patients during free breathing, without needing to include movement margins in the treatment plan. Respiratory correlation techniques developed to minimize motion artefacts in axial and helical scanning are

Figure 6. Osteolytic lesion of the spine. The British Journal of Radiology, Special Issue 2006

not applicable to CBCT and different techniques have been developed for the CB application. Sonke et al [54] describe a method for sorting the projections into different phases of the breathing cycle to produce a 4D CBCT scan. Sillanpaa et al describe a method of acquiring megavoltage cone beam CT projection images at the same phase of breathing at all acquisition angles, giving a three-dimensional reconstruction at a single breathing phase [55]. It must be emphasised that gated respiration techniques must be employed at both the localization stage and during treatment.

Quality assurance The accuracy of both conventional and CT simulation has a crucial effect on the overall accuracy of the patient’s treatment. Whereas the accuracy of conventional simulation relies mainly on geometric features such as gantry and collimator angles and field defining wire positions, that of CT simulation depends on the image obtained by the scanner and the faithful transfer to the virtual simulation software. This connectivity should be part of any quality assurance (QA) programme. A detailed description of quality control tests in conventional simulation and their recommended frequency is given by Tuohy [56]. Virtual simulation forms part of the network of the radiotherapy department, the end result of which is the treatment of the patient. The QA of this network should be seen as a process to which the various components of the hardware and software contribute. Guidance for the QA of a networked radiotherapy department is due to be published soon [57]. A QA programme should be established that reflects the importance of the contribution of each component of the system to the accuracy of the patient’s treatment. Some components will be checked daily, such as the alignment of the lasers, the accuracy of positioning of any moving lasers and the HU accuracy for water. Others may be checked monthly, annually or after significant upgrades to the system. Special phantoms have been designed to assist with various aspects of QA [58, 59]. The Kent Oncology Centre has produced its own phantom that incorporates checks for a number of parameters in one scan study. These include spatial resolution, HU number, slice thickness, alignment and geometric accuracy. Mutic et al [60] provide a comprehensive guide to the QA of CT simulators. They stress the need for audit and review of the process and flexibility in the programme as CT simulation evolves. S45

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Figure 7. Fusion of positron emission tomography (PET) and CT images from a CT/PET scanner to localize a left lung tumour.

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The future The aim of radiotherapy is to deliver a tumoricidal dose of radiation to the clinical target volume (CTV) whilst sparing normal tissue and critical organs as far as possible. Localization is aimed at answering the question ‘‘where is the target?’’ The gross tumour volume (GTV) is neither a simple line nor an unchanging volume. It is an oncological concept and will vary according to the imaging technique or techniques used, any additional clinical data available and the judgement of the clinician. Each imaging modality displays different information about the GTV. Traditionally, delineation of the GTV has been associated with an anatomical abnormality that is imaged by plane radiography, CT or in some cases MRI. This gives structural, not functional information. However, molecular and physiological imaging techniques are now available which give an indication of the functional state of the tissues. This information can potentially be used in addition to CT and MRI to assist in defining clinically relevant targets more accurately [61]. Ling et al [62] proposed treating a biological target volume defined from anatomical, physiological and/or molecular images. For example, increased glycolysis is a function of a tumour and fluorine-18 fluorodeoxyglucose positron emission tomography (18FDG-PET) studies have been used as an addition to CT for planning patients with poorly defined non-small cell lung cancer (NSCLC) [63, 64], head and neck cancers [65] and malignant gliomas [66] (see Jarritt et al in this issue). Figure 7 shows the fused images from 18FDG-PET and CT acquired in a single session on a PET/CT scanner. The lesion in the left lung is clearly demonstrated in both modalities in this example. Other PET agents may be used to identify areas of hypoxia within a tumour that may benefit from higher doses of radiation such as can be delivered by IMRT. Similar inhomogeneous dose distributions may be applied to regions of the prostate demonstrating a high choline:citrate ratio, indicating a region of active tumour, as demonstrated on MR spectroscopy [67] (see Payne and Leach in this issue). Modalities such as functional MRI (fMRI) and single photon emission computed tomography (SPECT) may also be used to assist in GTV and OAR delineation. SPECT perfusion studies for NSCLC can be used in treatment planning to provide information on normal lung tissue and help to reduce the volume of normal lung irradiated [68]. Imaging techniques are continually evolving and as they are refined they will reveal more information about the disease to be treated. Collaboration between radiologists and oncologists will be essential if the information contained within these new images is to be maximized for the benefit of the patient. No consideration of the future of radiation therapy would be complete without mention of image guided radiotherapy (IGRT). IGRT aims to address the interfraction movement of tumours and their relationship to OARs. Of the linear accelerator manufacturers, both Elekta (Crawley, Sussex, UK) and Varian (Palo Alto, CA) provide kilovoltage cone beam CT (CBCT) on the gantry and Siemens (Erlangen, Germany) have installed a CT scanner on rails in the treatment room (see Moore et al and Thieke et al, respectively, in this issue). The British Journal of Radiology, Special Issue 2006

These imaging devices provide the ability to localize the tumour immediately prior to treatment and to reposition the patient to correct for interfraction variation in tumour position. Wong et al [69] describe the use of daily scans in the treatment room to reposition prostate patients for the final phase of their treatment. 46% required no isocentre adjustment in the anterior–posterior direction, but 44% required a shift of greater than 5 mm. In the superoinferior direction, 25% required a shift greater than 5 mm and in left-right direction 24% required a shift greater than 5 mm. The shifts were associated with significant changes in the dosimetry. Other authors describe the implementation of CBCT for IGRT [54, 70, 71]. IGRT is a rapidly evolving field and will undoubtedly have implications for treatment planning.

Conclusion Both conventional and virtual simulation have developed in line with the changes in imaging techniques over recent years. The anticipated advantages of virtual simulation have been realised to a great extent and have changed the work flow in treatment planning. The availability of wide bore scanners enables most treatment techniques to be imaged. Fast computer graphics that have reduced image reconstruction times enable the acquisition of large data sets that can be manipulated for respiratory correlated techniques. The rapid development of biological imaging holds the prospect of multimodality localization, which is already being realised for some disease sites such as lung and prostate. The addition of cone beam CT to conventional simulators may add flexibility to departments with both a scanner and a simulator. However localization is achieved, it must be considered as part of the overall process that leads to treatment. The accuracy of the data acquisition and transfer is vital to this process and a comprehensive QA programme is essential. I would like to thank Dr Ruth Beddows for the design of the Kent Oncology Centre phantom, Ms Alison Vinall and Ms Karen Venables for reports on cone beam CT, David Hill for assistance with image processing and colleagues for discussions during the preparation of this manuscript.

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Developments in electronic portal imaging systems M C KIRBY,

PhD, CSci, MIPEM

and A G GLENDINNING,

PhD, MIPEM

North Western Medical Physics, Radiotherapy Department, Rosemere Cancer Centre, Royal Preston Hospital, Preston, UK ABSTRACT. Verification of geometric accuracy at the time of treatment delivery has always been a necessary part of the radiotherapy process. Since the introduction of conformal and intensity-modulated radiotherapy, the consequences of patient positioning errors are more serious. Portal imaging has played a large part in fulfilling the need for improved geometric accuracy. This review examines how portal imaging has progressed through the development and evolution of electronic portal imaging devices (EPIDs). Changes in technology, including the current commercial systems, and how image quality has changed are presented. The clinical usage of EPIDs and the technological innovations being devised for further improvements in image quality and systems are considered.

Technological developments in radiotherapy have been driven by the fundamental requirement to maximize dose to the intended target volume whilst minimizing dose to healthy surrounding tissue. Advances in collimation have enabled accurate, precise and more reproducible beam shaping and placement. Continuing accelerator and treatment planning system (TPS) development has further improved dose conformality around the target volume through intensitymodulated radiation therapy (IMRT). In addition, improvements in diagnostic imaging and their combination through different modalities are showing clinicians more precisely the anatomical areas that must define the high dose volume. However, such efforts can be made ineffective by poor patient immobilization, set-up and verification at the time of treatment delivery [1]. The ‘‘chain’’ of the radiotherapy process must include a link for treatment verification. The overall effectiveness of radiotherapy is dependent upon all links in the chain. Portal imaging, or taking images at the time of treatment delivery, is a key element of that verification. Its main use has been for geometric verification of field placement, but its dosimetric applications are also being developed. However, verifying the position of a soft tissue target volume with megavoltage energy X-ray beams is not trivial. Bony landmarks have to act as a surrogate for the target volume because, at megavoltage energies, the inherent subject contrast is naturally poor [2]. Soft tissue markers can be used for certain tumour sites with the advent of more advanced imaging detectors [3, 4]. Kilovoltage imaging could be used to improve the contrast of soft tissue and bony detail, and is the subject of current developments in image-guided radiation therapy (IGRT) [5]. See also the paper by Moore et al in this issue. This review paper will examine the developments in portal imaging systems over the last 30 years, and is divided into three main areas: (1) how the technology and methods have evolved; (2) clinical developments S50

Received 18 August 2005 Revised 6 March 2006 Accepted 9 March 2006 DOI: 10.1259/bjr/21517185 ’ 2006 The British Institute of Radiology

over that time; and (3) technical innovations being researched for future generation of EPIDs.

Technical developments Portal film Radiographic film was the traditional medium for acquiring images for treatment verification and is still used alongside EPIDs. Early images were acquired with industrial, direct exposure films, usually in cardboard holders, and required special processing techniques [6, 7]. In 1974, Kodak introduced film in ‘‘ready-packs’’, which could be processed in automatic, rapid processors [7]. In conventional film for high energy X-rays, the media is sandwiched between metal plates (usually Cu about 1 mm thick) [6–9] or the rear plate may be made of a thermoplastic [7, 10]. The front metal plate acts as a build-up layer producing high-energy electrons that in turn expose the film. It will also filter incident scattered radiation. The rear plate produces back-scattered electrons and helps maintain a tightly packed cassette. For the more recent, advanced film systems, additional phosphors are used either side of the film. For example, Kodak have a commercial film, known as enhanced contrast localization (EC-L), which provides much improved image quality through the use of a fine grain, high gamma film combined with two phosphor screens and a 1 mm Cu front metal screen [7, 11]. Electrons produced by the front plate interact with both phosphor screens, producing optical photons that expose the film. The detection efficiency of the X-rays is increased by a factor of 2 [10]. The total noise within the system is reduced due to the smaller size and distribution of the very fine grain microcubic crystals. The film also has a higher gamma (around 6) [11] giving higher display contrast – ideal for the inherent low contrast of objects imaged at megavoltage X-ray energies [2, 10, 12]. The British Journal of Radiology, Special Issue 2006

Developments in EPI systems

Advantages include high quality imaging, using simple, relatively lightweight cassettes, allowing for single/multiple exposures and with good, well-defined dosimetric properties. Disadvantages include the required processing, a fixed dynamic range and digitization required to enhance or manipulate the images. This makes film impractical for on-line imaging. They are also consumables and require significant storage solutions.

Computed radiography systems (photostimulable phosphors) Computed radiography (CR) systems were initially developed for diagnostic imaging [13] and used for radiotherapy from the 1980s [1, 6, 14, 15]. Film is replaced by a flexible plate, about 1 mm thick, coated with europium-activated fluorohalide compounds in crystalline formation embedded in an organic binding material. The photostimulable phosphor acts as an energy trap when exposed to ionizing radiation, producing a latent image. When scanned with a red (633 nm) laser, the energy is released as visible light, which is then converted to electrical signals via a photomultiplier tube [1, 14]. Commercial systems are available from Kodak (Figure 1), Agfa (www.agfa.com/healthcare) and Fuji (www.fujimed.com). These systems are used for diagnostic and therapy applications (e.g. by reversing the imaging cassette). Image quality is good and compares well with that from modern EPIDs (Figure 2). They have very similar advantages and disadvantages to those of film. However, they are reusable and the readout mechanism (effectively equivalent to film

processing) produces immediate digital images in multiple formats.

First generation EPIDs Electronic systems for portal imaging were first proposed in the late 1950s. Reviews by Boyer, Langmack and Antonuk chart their progress very well [2, 6, 10]. A variety of technologies have emerged over the years which may be classed broadly into (i) camera based systems, (ii) ionization chamber matrix devices and (iii) scanning array and other systems. Devices in classes (i) and (ii) were successfully commercialized. The motivation for electronic devices was simple: to obtain high quality images rapidly, with the versatility and flexibility afforded by computer digitization. With this premise, systems were designed to acquire quick digital images, making on-line enhancement, analysis and set-up correction a real possibility.

Camera based EPIDs The first electronic systems developed were camera based systems [2, 6, 10], three of which became commercially viable products for manufacturers Siemens (‘‘Beamview’’) [16], Infimed (‘‘Theraview’’) [17] and Philips/Elekta (SRI-100/’’iView’’) [18]. Two systems are shown in Figure 3. The principle of detection is similar to that described previously for film. The X-ray conversion stage is a metal plate (typically about 1–2 mm copper, steel or brass) bonded to a rare-earth phosphor (usually gadolinium

Figure 1. The Kodak CR System for radiotherapy verification. (Images from Kodak website www.kodak.com/go/oncology). The British Journal of Radiology, Special Issue 2006

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M C Kirby and A G Glendinning

Figure 2. Simple comparison of images acquired of an anthropomorphic type phantom using (a) an amorphous silicon (a-Si) imaging system (Elekta iViewGT) and (b) a CR system (Kodak 2000RT). Both images acquired using 6 MV X-rays. The a-Si image was acquired with 100 MU and the CR with 20 MU. (CR image courtesy of Steve Weston, Cookridge Hospital, Leeds, UK).

oxysulphide) of various thickness (100–400 mg cm21). Xrays incident on the metal plate produce high-energy electrons. The metal plate also filters some of the lowenergy scattered radiation (X-rays and electrons), which would otherwise reduce the contrast of an image [2]. The phosphor converts the high-energy electron energy (and also that of residual high-energy X-rays) into light. The

light diffusing through the phosphor on the camera side is reflected off a series of mirrors into a lens and camera. This enables the camera, and its radiation sensitive components, to be placed outside the primary X-ray beam, usually behind lead shielding [18, 19]. The lens and camera capture only a very small proportion of the light, which is one of the inefficiencies of such devices.

Figure 3. Two commercial camera-based electronic portal imaging devices (EPIDs). (a) The Elekta iView system attached to the gantry of an SL Linac. (b) The Infimed (Cablon) Theraview system in its retracted position. Also shown in (a) is a typical portal film cassette (Kodak EC-L system) in its portal film holder.

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The camera converts the image into a video signal that is fed into computer hardware for digitization, processing, enhancement and display. The spatial frequency dependent detective quantum efficiency (DQE(f)) is a widely used index of X-ray imaging performance [2, 10, 20]. It characterizes the efficiency of information transfer from the input to the output stages of a system. It can also be expressed as the ratio of the squared output signal to noise ratio (SNR) to the squared input SNR. The detectability of an object is dependent upon SNR [2, 12]. DQE depends upon the spatial resolution, gain and noise transfer properties of the system [2, 10, 20]. It is maximized by increasing the efficiency of detection of incident radiation, improving spatial resolution and decreasing system noise. For camera-based systems, the noise components are numerous [21], but the major inefficiency is light output and collection. Increasing phosphor thickness improves light output, but at the cost of spatial resolution [17, 22]. Generally, only about 10–20% of the light emitted by the phosphor escapes. More important limitations are the poor light collection efficiency and the electronic noise within the camera chain [21, 22]. Light is emitted from the phosphor isotropically but only photons emitted within a small cone subtended by the camera lens generate a signal in the camera. Between 0.1% and 0.01% of the light emitted reaches the camera [10]. If this light subsequently produces a small signal, then it may be swamped by noise generated in the camera and associated frame processing system [21, 22]. Attempts to improve light collection have included modifying the phosphor screen [23, 24] and using large aperture lenses. However, the latter introduce image distortions and aberrations, among other effects, which can reduce spatial resolution and cause non-uniformities across the image [22]. Camera and electronic noise can be reduced by different integration strategies and using cooled, low noise target chips on CCD type cameras [10]. The maximum DQE achievable has been about 1%. The main advantages of camera-based systems are that the whole image is viewed simultaneously, imaging is very fast (video rate), there is good spatial resolution and the system is relatively cheap to service and maintain. A practical downside is their ‘‘bulkiness’’ when in use, making it difficult to work around them for patient set-up.

Liquid ionization chamber matrix EPID This device was designed and first used at the Netherlands Cancer Institute in the 1980s and commercialized rapidly by Varian Medical Systems in the early 1990s [25, 26]. Its operational design and characteristics are well reviewed in the literature [2, 6, 10, 22]. Briefly, the ionization chamber is formed by two planes of electrodes, separated by a 0.8 mm gap. The gap is filled with an organic fluid (iso-octane or 2,2,4trimethylpentane), which acts as the ionization medium during exposure. Ions produced in the liquid are collected by the electrodes, of which there are 256 in each plane spaced 1.27 mm apart. The planes are oriented orthogonally forming a 2562 matrix. Primary X-rays incident on the front surface of a 1 mm thick The British Journal of Radiology, Special Issue 2006

plastoferrite plate produce high-energy electrons, which (in addition to non-interacting X-rays) produce ionization within the liquid medium. High voltage is applied to each electrode individually, but multiplexing electronics are used so that the image is readout sequentially row upon row. The detector and peripheral electronics can be tightly compartmentalized so that the EPID can retract fully (under motor control) into the gantry. A full resolution image is acquired and processed in just over 5 s; faster acquisition was initially made possible by reading pairs of electrodes, but at the cost of a lower spatial resolution. The most recent systems use a higher voltage and a shorter readout time per electrode, bringing full resolution image acquisition down to about 1.25 s [22]. For the commercial system, its greatest advantages were its compact, practical design making it easy to use, its lack of geometric distortion, but also its associated software [2, 22, 27]. Its main disadvantages were the sensitive nature of the control electronics surrounding the active detection area and, most significantly, the under-utilization of incident X-ray quanta compared with true area detector EPIDs. Whilst about 1.5% of the incident X-rays interact in the plastoferrite plate and liquid ionization medium, generating measurable signal [22], the DQE of the system is only about 0.5% due to signal loss in sampling [10]. This means that the dose for image formation is higher than true area detectors.

Other EPIDs used in the development of portal imaging Two other designs of EPID have greatly helped to progress the clinical developments of electronic portal imaging [2, 10]. The first was a variation of the camera based EPID [28], which used a construction of fibre-optic bundles to ‘‘pipe’’ the light directly from phosphor to camera. This made the device less bulky and more practical than the standard camera-based construction. However, image distortion and non-uniformity was a problem because of minor irregularities in alignment and shape of the fibre bundles at the output stage, and variations in light transmission through the bundles [2, 28]. The second was a scanning system that used a linear array of scintillation crystals optically coupled to photodiode detectors [2, 10, 29]. Zinc tungstate, bismuth germanate and caesium iodide were all used as scintillators in variants of the device. High quality images were produced, but scanning times were relatively long, especially compared with video based EPIDs. However, many clinical studies were performed using the device, as well as dosimetric and megavoltage CT investigations [6, 10, 30, 31].

Active matrix flat-panel imagers (AMFPIs) Initially conceived in 1987 by research scientists at the University of Michigan, Xerox PARC and elsewhere [10, 32–34], indirect detection active matrix flat panel imagers (AMFPIs) now form the main focus of developments in portal imaging. All manufacturers offer this technology S53


Figure 4. Active matrix flat panel imagers (AMFPIs) from two of the major radiotherapy equipment manufacturers; (a) Elekta and (b) Siemens. Electronic portal imaging devices (EPIDs) of this type are highly practical devices, being fully motorized and retractable to the gantry of the linear accelerator when not required for use. The AMFPI from Varian uses similar gantry mounting technology.

and, because of their enhanced image quality and potential for dosimetric uses, they constitute an effective second generation of EPIDs. These devices became commercially available in 2000; two devices are shown in Figure 4.

AMFPI technology and process The technical details for such EPIDs are covered in recent review papers [2, 6, 10, 34]. In brief, AMFPIs consist of a number of subsystems. These are (i) a large area, pixelated array, (ii) an X-ray converter, (iii) a control and acquisition system and (iv) a computer for controlling acquisition and processing, enhancing and archiving the images. The array consists of a ,1 mm thick glass substrate upon which the electronic circuits reside. Each pixel in the active matrix-imaging array incorporates a thin-film switch connected to a sensor. The pixels are arranged into a two-dimensional grid with the conductivity of each switch regulated by varying the voltage of control lines, where each control line is connected to all the switches in a particular row. During irradiation, the switches are kept non-conducting so that radiation generates electrical signals integrated in the capacitative element of each pixel. At the time of readout, the switches are made conducting, one row at a time. At this point, the charge in each pixel is transported to external electronics by means of the data lines. Each data line is connected to all the pixels in a given column, so a line of data is read [10]. There are two general methods for the initial conversion of incident X-ray energy into charge stored in each of the capacitative elements of the AMFPI. These are S54

Indirect or Direct detection methods [10, 34]. For Indirect detection, a metal plate/phosphor screen combination is used as the X-ray converter. High-energy X-rays and electrons produce light in the phosphor or scintillator, which is positioned directly over the photosensor of the array. Optical photons are converted into electron-hole pairs in the photosensors, which also act as the capacitative element in each pixel until readout takes place using the pixel switches [6, 10]. For Direct detection, the X-ray converter consists of a metal plate and a photoconductor. The photoconductor is electrically coupled to a separate capacitor built into each pixel. Electron-hole pairs are produced in the photoconductor, which are then stored in each capacitor before readout [10, 35, 36]. At present, all commercial AMFPIs use the indirect detection method [10]. Image acquisition can be fast (up to 10 frames per second), with dynamic ranges up to 16 bits deep and pixel resolutions up to 10242 in matrices up to 41 cm641 cm at the detector [10, 37]. Readout is usually synchronised to occur between linac pulses. Image correction and processing is similar for all the commercial devices, consisting of the use of gain (flood field) and offset correction as well as a method of filtering for reducing or eliminating the influence of bad pixels or line defects [10, 38–40]. Being a flat panel design, there is no spatial distortion present in the resulting images [41].

Image quality – qualitative comparisons Resultant image quality is very good, but this is to be expected since (i) there is good, close contact between the phosphor and the photodiode array (improving The British Journal of Radiology, Special Issue 2006


differences of up to 2–3 mm in identifying the edges of structures [41]. Improvements in SNR are easily seen and, since the quantity of noise is greatly reduced, image processing can be applied to the images with greater benefit. Figures 6 and 7 show examples of this comparing the Varian as500 AMFPI and liquid ionization chamber matrix EPID. Figure 6 shows an enhanced image for a lateral pelvic treatment. It illustrates how enhancement for the ionization chamber EPID image does show bony detail like AMFPI images, but that the noise within the image is equally enhanced by the processing. Figure 7 shows how the reduced noise characteristics of the AMFPI permits the use of more aggressive image processing techniques.

dramatically the light collection efficiency), (ii) the areas of the photodiodes pixels and the phosphor are similar, (iii) there is a high conversion efficiency for optical photons into electron-hole pairs and (iv) the readout of signals from the pixels is also highly efficient. The noise characteristics are considerably improved over previous systems, which means that both the SNRs and contrast to noise ratios (CNRs) are better than all the first generation EPIDs. This results in a higher DQE for the flat panel imagers, and shows that the devices are now quantum noise limited. Pixel size for all commercial systems is smaller than previous EPIDs, with most devices having arrays which are 10242. Spatial resolution is better although this depends upon the phosphor thickness and not solely on the pixel size [10, 34, 37]. The improvements can be seen qualitatively in Figure 5. These images have been acquired with Elekta equipment, but the comparisons are similar for all commercial systems. They compare first generation technology (iView, a camera-based EPID) [42] with a newer AMFPI (iViewGT) [10, 37]. Images of an anthropomorphic phantom are ideal for assessing comparative image quality between different EPID technologies. Images of the head and neck region (particularly of cervical vertebrae) have always been an excellent visual indicator of the improvements in EPID image quality. Bony detail in the chest (such as the lumbar spine and ribs) and in the pelvis was visible with the earliest EPIDs. However, fine detail, such as the posterior spinous processes in the cervical vertebrae, have always been hard to identify. The expected gains in terms of reduced noise and improved spatial resolution are clearly visible. The graininess is much reduced for the AMFPI images and there is a considerable increase in the visibility of low contrast, high spatial frequency bony detail, which improves with exposure (Figure 5). A simple comparison of structures outlined on the 2 MU exposure images from the two different technologies shows that there can be

(a)

Image quality – quantitative comparisons Quantitatively, the improvements in spatial resolution and SNR are clearly demonstrable. Using contrast detail type-phantoms (such as those described in Kirby et al, Hermann et al, Dong et al, Harrison et al and Low et al [19, 22, 43–45]), AMFPIs demonstrate improved detectability of small-scale (1 mm) low contrast (1.2%) objects at 6 MV [41]. This is similar to that of EC-L film [6]. These thresholds are exposure dependent (as expected) and show that phantoms like the Las Vegas [19, 22] are no longer suitable for assessing AMFPIs; for certain exposures and energies, all objects are now visible [41]. Phantoms like the PIPS QC3V (www.standardimaging.com) can give quantitative analysis of spatial resolution and noise characteristics (such as CNR) [46]. This phantom has been used in studies to examine and compare the image quality of AMFPIs from different manufacturers, the same manufacturer (over a protracted period of time) and also compared with first generation EPIDs [37, 41, 47]. Clements et al [37] examined indices of CNR, 50% (f50) and 30% (f30) relative modulation transfer function

(b)

(c)

Figure 5. Images comparing, qualitatively, the image quality obtainable with a camera based first generation EPID (iView) and a new AMFPI system (iViewGT). The subject is the neck region of an anthropomorphic type (‘‘Rando’’) phantom. All images have been acquired with 6 MV X-rays. (a) iView, 2 MU, (b) iViewGT 2 MU, (c) iViewGT 100 MU. From Kirby et al 2002 [41]. The British Journal of Radiology, Special Issue 2006

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(a)

(b)

Figure 6. Images comparing, qualitatively, the image quality obtainable with (a) an ionization chamber matrix electronic portal imaging device (EPID) (PortalVision) and (b) a new active matrix flat panel imager (AMFPI) system (as500). Both images show bony detail for lateral pelvic treatment fields after image enhancement. (Images courtesy of Margaret Bidmead, Royal Marsden Hospital, London, UK).

(RMTF) for three early examples of commercial AMFPIs. Their results indicated that some systems demonstrate better spatial resolution whilst one showed better noise characteristics. All examples were considerably better than first generation EPIDs. These results were also observed in a multicentre study (Table 1), which examined 10 iViewGT systems

and 7 iView systems from four radiotherapy centres in the UK, over a 12–18 month period (from first installation for the iViewGT systems) [47]. The data show that spatial resolution can be improved by over 50% compared with camera based EPIDs. The CNR is also improved at the longer exposures (by a factor of 4), as expected.

(a)

(b)

Figure 7. Images showing the visibility of bony detail (a) before and (b) after image processing and enhancement. Image acquired with the Varian as500 AMFPI. (Images courtesy of Margaret Bidmead, Royal Marsden Hospital, London, UK). S56


Developments in EPI systems Table 1. Data from the QC3V phantom examining the contrast to noise ratio (CNR), f50 and f30 values of 10 active matrix flat panel imagers (AMFPIs) (iViewGT systems) and 7 camera based electronic portal imaging devices (EPIDs) (iView systems). f50 and f30 values are in lp mm21. Data are a ‘‘snap-shot’’ of values taken in 2003 across the 4 UK radiotherapy centres for various exposures. All data are for 6 MV X-rays, 15 cm615 cm collimator setting, with the phantom in contact with the EPID (approx. 153 cm focus to surface distance) System

f50

Short exposure (2–10 MU) iView Mean 0.265 Range (0.229–0.301) iViewGT Mean 0.421 Range (0.355–0.523)

f50

f30

f30

CNR

Long exposure (100 MU) Mean 0.272 Range (0.232–0.308) Mean 0.432 Range (0.400–0.537)

Short exposure (2–10 MU) –

Long exposure (100 MU) –

Mean 0.724 Range (0.627–0.791)

Mean 0.741 Range (0.685–0.779)

Short exposure Long exposure (2–10 MU) (100 MU) Range Range (40–65) (100–200) Range Range (65–90) (450–550)

The data in Table 1 are a single ‘‘snap-shot’’ of the data collected from the EPIDs across the radiotherapy centres during the early part of 2003 [47]. They do not necessarily reflect results expected from newly installed AMFPIs, or represent the best values obtainable. The study showed clearly that there is greater variation in results between examples of the same type of AMFPI compared with older technology. Some of the AMFPI EPIDs (those between 12 months and 18 months old) showed that the image quality was fairly stable; most failures in these early examples were due to wholesale changes (e.g. failures in the acquisition electronics) rather than degradation in the active detection matrices. Note that the examples shown for the camera-based EPIDs are not as good as the most recent systems. Improvements have been made in terms of light collection efficiency [10, 22] and noise characteristics of the camera [10, 48–50].

Problems with commercially available AMFPIs As the use of AMFPIs grows, data are being accumulated showing current imaging problems and long-term stability of the devices. Tolerance to radiation damage was proven in the early prototype versions, with the active imaging area able to withstand doses of about 104 Gy without significant changes in performance [34]. In the absence of radiation, pixels accumulate signal due to ‘‘dark current’’ effects resulting from array and sensor leakage currents [34, 38]. These currents are small, which is a requirement if they are not to contribute significantly to the additive noise within the system as a whole. The dark current is stable over long periods of time [34], although there is a dependence upon temperature which may be accounted for dynamically [40]. Ghosting or image lag is another recognized problem [38, 39, 51–53]. Here, charge trapping within the photodiodes (post-irradiation) manifests itself as a latent image on subsequent frames, or alters the gain (or sensitivity) of the a-Si layer itself [38, 39, 51]. The magnitude of the effects is dependent upon exposure and acquisition time (frame rate) [52, 53]. A number of other artefacts have been noted in the clinical setting. These are all types of structural (or fixedpattern) noise since they are produced in a non-random manner [21]. They significantly affect the visual image quality of the clinical images; more so than random noise components (such as X-ray quantum noise and electronic The British Journal of Radiology, Special Issue 2006

CNR

noise). Some examples are shown in Figure 8. They include:

N N N N N

Artefacts caused by start-up characteristics of the linear accelerator (coupled with readout methods on the AMFPI) [41] Structural noise contributions from image correction strategies (such as offset images) [41] Incorrect synchronization of image acquisition and readout with the pulses of the linear accelerator [6, 12, 54] Ring artefacts and other non-uniformities produced by inappropriate weighting of image and correction data or the selection of correction images for the wrong energy [41, 55] The influence of bad pixels over time [40, 47]

The most common failure modes appear to be associated with damage to the external acquisition and readout electronics around the array [34, 40, 54]. Once affected, whole sections of the panel may cease to function, or rows of individual pixels begin to accumulate no data at all [40, 54], as shown in Figure 8d. The timescale for these types of failure is still quite variable, ranging from a few months to years. The failure may be catastrophic (e.g. losing a whole panel section so there is no image data), or more gradual (Figure 8d). For the latter, the artefact becomes more predominant over time, eventually making it impossible to adequately assess clinical images. Panels have been changed for this type of problem after about 2–3 years. One difficulty with AMFPIs, at present, is their maintenance/replacement costs. Currently, the best first generation EPIDs (such as cooled charge coupled device (CCD) camera systems) can produce images as good as the AMFPIs, for certain applications. Whilst this is the case, the user is faced with the dilemma of procuring a mature, well-established system with minimal running costs or a system with more advanced technology and improved imaging characteristics, but with much higher maintenance costs and a variable lifetime.

Clinical developments Clinical implementation and image acquisition A number of published guidelines are available on clinical implementation [22, 56, 57]. Changes in working S57


(a)

(b)

(c)

(d)

Figure 8. Examples of some typical artefacts that have been observed with commercial active matrix flat panel imager (AMFPI) electronic portal imaging devices (EPIDs). (a) Component of the offset correction image present in clinical images at low exposures (linear, streaking artefact which has been ringed). (b,c) Examples of line artefacts arising from pulsing of the linac. (d) Example of loss of signal in rows of pixels in a complete segment caused by radiation damage to the associated electronics (ringed area).

practices may be extensive and depend upon the verification strategy being used. Strategies for image acquisition, analysis and set-up correction are varied and can depend on anatomical site [22, 57]. Implementation is best achieved using a multidisciplinary approach and the full range of technologies and strategies available [12]. EPID acquisition modes are varied and image quality is now better than film for certain applications. It is in the flexible use of digital information and availability of processing and analysis software that EPIDs have shown their real value [6, 58, 59]. Good images are possible from both short exposures (for on-line applications) and also from the entire duration of a treatment field exposure (Figure 9). Double exposures or single open field images may be used where treatment fields alone show insufficient anatomy for analysis. This type of acquisition requires high quality imaging with very short exposures, which is achievable with both film and EPIDs. However, even with the latest AMFPIs, the image quality may be S58

dependent upon exposure (Figure 9). EPIDs also enable modes of image acquisition that are only possible with digital technology [60–62]. These include movie loop acquisition and display. Individual frames within the loop could be acquired during a single exposure or may be images from different fractions. Images may be acquired from the individual segments of a complex technique, summed or manipulated digitally and then displayed.

Evaluating field placement errors The main goal of portal imaging is verification of the geometric placement of the irradiated high dose volume [6, 22]. Deviations between the actual and planned geometric positions are termed field placement errors (FPEs) or set-up errors. The FPE will often vary between fractions (interfraction variation) and sometimes during a treatment fraction/field (intrafraction variation). When analysed over the course of treatment, it may reveal The British Journal of Radiology, Special Issue 2006


(a)

(b)

Figure 9. Images acquired with an active matrix flat panel imager (AMFPI) for a lateral pelvic treatment (6 MV X-rays). (a) An image acquired with a reticule present to indicate central axis position, orientation and verify scaling. A 4 MU exposure was used. (b) An image acquired with the full treatment exposure (approximately 80 MU). Distinct differences can be seen in image quality between the two, with (a) having a poorer signal to noise ratio (primarily due to the increased presence of both structured and random noise components).

systematic and random variations [31, 57, 63]. These may vary between different techniques and anatomical sites [31]. There are various strategies that could be used to evaluate and correct for these FPEs, for example, on an individual image basis in which case both random and systematic components are eliminated. Alternatively, a series of images over a number of fractions may be acquired, analysed and a protocol used to evaluate and correct for the systematic component. Analysis of the random component may be used to determine the margin used in treatment planning to account for motion (both organ motion and set-up error) [63–65]. In most cases, bony anatomy must be used for assessment [22]. However, bony landmarks may not always represent the true position of the PTV. This may be due to relative movement between the respective tissues, normal physiological functions (such as bladder and rectal filling), respiratory movement etc. For certain sites, radio-opaque seed markers may be introduced into the soft tissue within the planning target volume (PTV). Acceptable images are possible with EPIDs (especially AMFPIs) showing the position of the markers, and studies have shown that there can be differences between FPE evaluated through bony anatomy compared with implanted markers [57]. For any evaluation of FPE, some reference image must be available. This may be a verified simulator radiograph or a digitally reconstructed radiograph (DRR) from the original CT planning data. Bony anatomy is compared The British Journal of Radiology, Special Issue 2006

with the linac coordinate system for both reference and treatment images. The linac coordinate system may be identified using field edges [6, 22, 57] or through an indication of the position of the isocentre (achieved during EPID calibration [66] or with a reticule, as shown in Figure 9a). The reticule is easy to use, but has practical disadvantages in that operators need to enter the treatment room to fit it, and it limits the exposure. However, a simple imaging strategy helps minimize the disruption. For example, two orthogonal images could be acquired with the reticule, and then it may be removed prior to treatment. Alternatively for a treatment that includes, say, an anterior and lateral treatment fields, imaging may be performed at the end of one field and at the beginning of the next. If auto (assisted) set-up is used, then the operator need only enter the room twice rather than four times. Identifying the field edges by automatic methods was developed in first generation EPID software. It may also be used for double exposures, whereby contrast is improved for the open field portion by simply overlaying the field edges of the treatment field [6, 57]. FPE evaluation may be performed using simple measurement (using digital rulers etc.). Alternatively individual points or anatomical outlines may be used (outlined manually or semi-automatically) and software used to register images providing a quantitative index of FPE. Much work has been done to automate this [6, 22, 59, 67], but fully automated analysis is still not possible for all anatomical sites. When quantitative analysis is S59


performed on multiple images (usually orthogonal) then both 2D and 3D information may be obtained for the overall set-up error [6, 59, 68]. Once the FPE has been identified, corrections to setup may be required. If the imaging, evaluation and analysis are performed using short exposures at the start of a single fraction, then it is possible to correct patient set-up prior to delivery of the full treatment. This on-line strategy addresses both systematic and random error components simultaneously, although possibly at the cost of patient throughput [56]. In addition, it needs to be performed every fraction in order to be effective (for patients with a significant systematic error), with possibly higher concomitant dose to normal tissues. More common is analysis of a series of images from different treatment fractions and a protocol used to identify the systematic and random errors. This is an offline strategy [57]. Correction of patient set-up is made for the systematic component for the rest of the treatment course with or without further imaging. Data from a population of patients for a particular anatomical site helps to validate the treatment technique itself and quantify the margins required for treatment planning [63, 65]. For both on-line and off-line strategies, integration of the EPID analysis with the record and verify or network system is highly desirable [69]. This verifies the set-up correction itself and also makes full analysis of individual and population data easier. A number of commercial developments are available for this type of networking. Effective clinical use can only be maintained through appropriate quality assurance for the EPID and its clinical use [70, 71]. This includes quality control for the hardware (to maintain optimum image quality) and also the software (for ensuring that image evaluation and analysis is correct) [57, 72, 73]. The clinical implementation and results should also be periodically reviewed [31, 56, 57].

Verification of patient dose in vivo The verification of patient dose in vivo is one of the main dosimetric applications [30, 50, 83–89]. One approach is to predict the dose at the plane of the EPID from the treatment plan and compare it with EPID measurement. Another is to use the twodimensional information within the portal image to determine the dose in the patient at a point or within a plane. By combining this information with the pretreatment CT dataset, a volumetric model of the dose distribution can be produced. Ultimately combining EPID dosimetry with treatment time volume imaging will produce a model of the actual delivered dose distribution [76]. To date, the accuracy achieved using first generation EPIDs is between 2% and 5% [22]. Initial studies suggest that AMFPIs are better suited to dosimetry (the calibration and modelling is simpler) and accuracy might improve [39, 40, 81].

IMRT verification Verification of IMRT delivery is achievable with the EPID in a dosimetric mode [83]. Images of individual segments (for step and shoot IMRT) or full dynamic delivery may be acquired, and MLC leaf positions computed and compared with those predicted [90–92]. Alternatively, the data may be used to compare predicted with actual delivered fluence in one or two dimensions [6, 22, 82–84, 88, 93–95]. One group has used the compact nature of AMFPIs in an ingenious way to perform in-phantom dosimetry for IMRT verification [96]. Recently, measured pre-treatment fluence data have been used as the input to a TPS to compute dose distributions [97]. Some groups are also examining the extraction of geometric information for patient set-up during the actual delivery of IMRT [95, 98].

Technical innovations – the future for EPIDs

Volume imaging is achievable through various methods; in-room CT systems [74], kilovoltage cone-beam CT (CBCT) equipment [4, 5, 75] and megavoltage CBCT techniques [76–79]. Much research is being conducted into all these methods, their clinical implementation and the issues of concomitant dose which they deliver [80]. See also papers by Thieke et al, Moore et al and Chen et al in this issue.

A number of advances are being made to improve the performance of EPIDs for both planar and volume imaging. Ultimately, imaging of both kinds requires optimum image quality whilst minimizing concomitant radiation doses to the patient. There are three main areas of development. These are (i) looking at further improvements to current EPID technology, (ii) examining ways in which the radiation source itself could make a difference to image quality and (iii) examining completely new technologies for EPIDs.

Non-geometric clinical uses

EPID modifications for improving image quality

Although the primary use of portal imaging is to verify geometric field placement, dosimetric verification is also possible under special circumstances. For any dosimetric application the EPID response must be a well-defined, quantitative function of the dose delivered, it must be stable and reproducible [53, 55, 81, 82]. All commercial EPIDs have been investigated and a number of applications are possible [83].

All current improvements in EPID design focus on enhancing the quantum efficiency of the detector.

Volume imaging

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Changes to camera based EPIDS For camera based EPIDs, attempts are being made to increase efficiency by changing the phosphor, adapting The British Journal of Radiology, Special Issue 2006


the light collection mechanism and improving the camera. Light collection efficiency can be improved with large aperture lenses (although possibly with consequent spherical aberrations [10, 22]), and larger format camera target chips. Noise within the system can be reduced by using Peltier cooling mechanisms for CCD based cameras [48, 94, 95]. Changes in the phosphor include adaptations of current materials, but in different geometrical configurations (such as ‘‘grooved’’ screens [23]). Conventional phosphors are rare earth based materials (such as gadolinium oxysulphide Gd2O2S), but their quantum efficiency is only a few percent for 6 MV X-rays. This can be increased 5–6 fold by using CsI crystals. MoslehShirazi et al [77] have used segmented 3 mm63 mm610 mm CsI(Tl) crystals, changing the Xray quantum detection efficiency from 2–3% to about 18%. Each crystal is isolated optically by introducing titanium dioxide powder into the 0.3 mm gap between them. A 3 mm thick aluminium plate acts as the initial Xray converter, bonded to the crystal layer. Sawant et al [99] have used a large-area array of CsI in a novel geometric arrangement. Here, incident X-rays pass through the mirror before striking a transparent lead-glass plate (about 1 cm thick), which is coupled (source side) to a transparent CsI(Tl) scintillator layer 1.25 cm thick. Thus, the camera-lens-mirror mechanism is located source side of the X-ray converter and scintillator. For this system the X-ray quantum detection efficiency is about 25%.

Changes to AMFPIs Indirect detection AMFPIs are also being investigated with different phosphors or phosphor configurations. Seppi et al [78] have used individual 0.38 mm60.38 mm68 mm CsI(Tl) crystals. Each crystal was coated (on five sides) with a reflective powder/ expoxy resin mixture, with the uncoated end in contact with the flat panel sensor array. Sawant et al [20] have investigated a design using a two-dimensional matrix of cells dimensionally matched to the pixels of the active array. Each cell is optically isolated from its neighbours and the inside of each cell is reflective to maximize the light reaching the photodiodes. The cells are packed with Gd2O2S:Tb powder, with 3–7 mm grains. The results are promising, but not as good (in terms of DQE) as expected. This is attributed to the strongly depth-dependent light escape efficiency of the phosphor – clear scintillator materials promise further gains. High QE AMFPI detectors using the direct-detection mechanism are also being investigated [10, 36, 100]. For these devices, a continuous amorphous selenium (a-Se) layer is deposited over the active array. The thick a-Se layer, coupled with a build-up layer, converts X-rays into secondary electrons which are directly converted into electron-hole pairs in the energy sensitive a-Se layer. The holes are collected on the pixel electrodes, stored on the pixel capacitors and read out through the matrix. Even for these devices, attempts are being made to couple their high efficiency with improved spatial The British Journal of Radiology, Special Issue 2006

resolution. One group is proposing a design which uses a photolithographic process to create a large number of packed, aligned and focused micro-structured plates [101]. The cavities between the plates are filled with an ionization medium (such as a-Se) surrounded by highdensity materials.

Linac beam-line modifications Some research groups are examining changes in the linac conditions that would favour verification imaging. For example, using a low Z target with low MV energy X-ray or electron beams gives the imaging beam a greater component of kV energy X-rays [102–104]. Low Z materials have been used in place of primary electron scattering foils [105], in place of the main X-ray target itself [102, 105] and also in the secondary electron scattering foil holder [104]. Some configurations remove the X-ray flattening filter to improve X-ray output and maintain a softer beam spectrum [105]. Some clinical results are now available (using the new beam lines with radiographic film) for head and neck patients showing better overall contrast for anatomical features than with film [106].

New designs for EPIDs One of these novel approaches is the use of kinestatic charge detection methods [10, 107]. These were originally designed for diagnostic applications and use a scanning technique. The detector has separate volumes for detecting the incoming X-rays and collecting the subsequent signal. The detection volume is continuous, but the collection volume consists of a linear array of detectors. A narrow, fan beam of X-rays is used so that when the X-rays interact, a line of ionization charge is produced. An electric field is applied across the volume which drives the charge towards the linear array. Scanning the detector in the opposite direction to that of the driven charge, at a velocity equal but opposite to the mean signal charge drift velocity and perpendicular to the X-ray beam, produces a 2D image. Using a gas (such as Xenon) as the ionization and detection media gives a predicted X-ray quantum detection efficiency of about 36% [107]. Another device is also based upon a gaseous detection medium. It uses gas electron multipliers (GEMs) with the potential for dual energy imaging [10, 108]. A single gas converter overlying a GEM might be used for diagnostic X-rays and a combination of solid converters and GEMs used to increase the detection efficiency for megavoltage imaging.

Expected gains from modified and new EPIDs All the innovations described aim to increase the efficient use of the X-rays incident upon the detector. The absolute contrast of anatomical objects will inherently be limited whilst megavoltage energy X-ray beams are used. However, improved X-ray quantum detection efficiency and noise reduction throughout the system S61


will increase SNR and thus the overall DQE. This would enable improvements in image quality for low dose imaging. Improvements in SNR can help in other areas; most notably in allowing the use of more aggressive image processing techniques, making automatic feature detection and extraction algorithms more achievable and improving the detection of implanted markers. It would also improve image quality for volume imaging, reducing the concomitant dose still further and thereby allowing their more frequent use.

Conclusions Great advances have been made in the technical development and clinical use of electronic portal imaging systems. Image quality has improved and, with AMFPIs, promises to progress still further. Cone beam CT techniques have become a practical reality and pave the way for more precise IGRT. Although artefacts, longevity and replacement costs are still being addressed, AMFPI is the technology of the future for EPIDs. The best first generation EPIDs are probably close to their optimum image quality achievable, whereas flat panel technology is still in its development. However, there is still a role for film and other commercial EPIDs and verification devices within a modern radiotherapy department. Clinical implementation and use of EPIDs has reached maturity and protocols show the way to full use. The advent of complex radiotherapy techniques has demonstrated portal imaging is vital. The systems and software now available make it a real, integrated part of the verification process for improving patient positioning and perhaps, in the future, dosimetrically verifying and fine-tuning the actual treatment plan.

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Developments in and experience of kilovoltage X-ray cone beam image-guided radiotherapy 1

C J MOORE, PhD, 1A AMER, PhD, 1T MARCHANT, PhD, 1J R SYKES, MSc, 2J DAVIES, DCR(T), 2 J STRATFORD, DCR(T), 2C MCCARTHY, PgDip, 3C MACBAIN, MD, 3A HENRY, MD, 3P PRICE, MD and 1 P C WILLIAMS, PhD 1

North Western Medical Physics Department, 2Radiotherapy Research Facility and 3Academic Department of Radiation Oncology, Christie Hospital, Manchester M20 4BX, UK

ABSTRACT. This paper offers a realistic review of kilovoltage X-ray cone beam tomography integrated with the treatment machine for image-guided radiotherapy in the light of experience taking a commercial system from prototype development into clinical use. It shows that key practicalities cannot be ignored, in particular the regular characterization of mechanical flex during gantry rotation, the mapping of defects in flat panel image transducers and their response to X-ray exposure. The number of X-ray projections and the doses required for clinically useful cone beam reconstruction at different therapy sites are considered in the context of imaging that is fit for purpose. Three roles for cone beam tomography in radiotherapy are identified: patient setup in three dimensions (3D), where even low dose cone beam tissue detail is superior to megavoltage imaging; disease targeting where, despite wide field scatter and slow scanning, it is possible to generate images that are suitable for tumour delineation even at challenging sites; adaptive treatment planning, where calibrated cone beam images have been shown to provide sufficient target detail to support ‘‘plan of the day’’ selection and have the potential for planning with bulk corrections. With frequent use in mind, the need to limit patient dose during setup, yet maximize much needed image quality in the target zone, is considered. Finally, it is noted that the development of cone beam tomography for radiotherapy is far from complete, with X-ray source, image transducer, reconstruction algorithms and techniques for image profile collection still being researched.

Historical background and introduction In 1984, industrial researchers Feldkamp, Davis and Kress (FDK) [1] reported a circular scanning algorithm for three-dimensional (3D) X-ray cone beam tomographic (CBT) reconstruction. However, even two-dimensional (2D) reconstructive imaging was slow to appear in the radiotherapy treatment room, although in 1992 Lewis et al [2] described 2D, megavoltage X-ray, computerized tomographic (CT) reconstruction. This delay can be understood in the context of the operating tolerances for many radiotherapy techniques, which remained relatively generous until the mid-1990s when the impact of the recommendations set out in ICRU report 50 [3] and the development of computerized multileaf collimation (MLC) for conformal radiotherapy began to be felt. Until then, it had been sufficient to base patient setup on skin marks and verification on the bony detail visible in wide field, megavoltage verification films. Megavoltage electronic portal imaging devices (EPIDs) appeared in the late 1980s [4], offering the advantage of integration with the gantry of the treatment machine and instant availability of a digital image without the need for film processing. The authors gratefully acknowledge funding support from the Christie Hospital Centenary Appeal Fund, Elekta Oncology Systems, Cancer Research UK and The Royal College of Radiologists.

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Received 1 July 2005 Revised 10 October 2005 Accepted 25 January 2006 DOI: 10.1259/bjr/68255935 ’ 2006 The British Institute of Radiology

Alignment and use with the treatment source provided beam’s eye views (BEVs) at arbitrary gantry angles. Early EPIDs offered adequate spatial resolutions of a few millimetres in the imaging plane and repeated use for time lapse imaging of tumour motion, notably in the lung. The obvious alternative to the EPID, a CT scanner introduced into the confined space of an existing treatment room, had none of these particular advantages for setup and monitoring. CT had been carefully optimized for temporally and spatially discontinuous, sequential transaxial scanning, with images accurately calibrated in Hounsfield units. These were ideally suited to treatment planning, but not to BEV production. Antonuk’s work on X-ray image transducer panels [5], employing selenium and amorphous silicon (aSi) technologies, appeared to promise a significant improvement in the application of the EPID, despite the underlying physical limitations constraining megavoltage imaging to physical density mapping (see Chen et al in this issue). However, as well as improving portal imaging these efficient, high resolution flat panel transducers provided the basis for the practical implementation of wide-field CBT imaging for radiotherapy using kilovoltage X-rays. Conformal radiotherapy, implemented with ICRU guidelines in mind [3], requires image grey level detail to be reduced to a set of explicit structural boundaries measured within the treatment room frame of reference. The British Journal of Radiology, Special Issue 2006

kV X-ray cone beam image-guided radiotherapy

Indeed, spatial measurement from CT image sequences, effectively an image volume, has become a distinct activity in the modern radiotherapy treatment planning room. Clearly ‘‘image assisted’’ planning has the potential to develop still further as functional imaging, in the form of MRI and PET, becomes more widely available. In the last decade, this geometric conformation of irradiation fields to projected tumour shape has rapidly evolved into more precise 3D dose conformation, the most demanding application of this being intensitymodulated radiotherapy (IMRT). In this technique, MLC technology is used to sequentially superimpose differently sized and offset beam segments to form the equivalent of one intensity-modulated beam. In the treatment room, this technically complex procedure has propelled radiotherapy from 2D setup imaging into 3D image-guided radiotherapy (IGRT). Setup imaging has been supported by the assumption that the correlation between readily identifiable and stable features, usually rigid bony anatomy, and the tumour target is maintained within certain margins for each fraction of the treatment. Orthogonal portal images showing bony detail, if not the target itself, are then sufficient for setup against digitally reconstructed radiographs (DRRs). A further assumption is that patient rotations can be identified unambiguously from a pair of superposition images. IGRT aims to actively inform setup and disease targeting by imaging bone and soft tissues, showing tumour and organs at risk in 3D, as close to moment of therapeutic irradiation as practicable. The adequacy of targeting, and even the pretreatment plan itself, can then be assessed quantitatively and corrective action taken as appropriate. This definition of IGRT indicates a measurement process in the treatment room on a scale that rivals image assisted pretreatment planning. Consequently, IGRT has stimulated the rapid and continuing development of complementary, commercial CBT systems, the subject of this paper. This is a case of X-ray imaging development being driven by a need that is not diagnostic and it signals the first widespread use and commercial availability of low energy X-ray imaging in the treatment room.

Practicalities and the evolution of clinical cone beam systems for radiotherapy When CT scanners became available in the 1970s, it became possible to gather 2D axial image sequences with enough soft-tissue contrast for tumour delineation and to use image processing to extract body contours and organ boundaries for dosimetric modelling in radiotherapy planning (RTP). In addition, pixel by pixel dose computation became an alternative to planning based on bulk inhomogeneities. These attributes have influenced the specifications underlying the recent commercial development of CBT 3D-imaging for radiotherapy by the major manufacturers. Despite ongoing research seeking to generate reliable sources of low energy X-rays from the linear accelerator (linac) itself, all the major treatment machine manufacturers have opted to offer kilovoltage X-ray tubes mounted on the treatment gantry for CBT. They are using flat panel transducers to acquire a rotation fluoroscopy image sequence for input to 3D filtered back-projection algorithms. The British Journal of Radiology, Special Issue 2006

At least for now, the algorithms are mostly extensions of the approach originally described by FDK in 1984, who had in mind the needs of the industrial non-destructive testing communities. This is a world apart from radiotherapy where professionals might have unduly high expectations of CBT because of their acquaintance with diagnostic CT. FDK CBT is based on the circular rotation of a cone about a fixed isocentre. This has a central plane with projection profile that appears to be from a 2D object. Hence, in a precalculated manner, it is filtered prior to reverse projection along the paths of the original X-rays, so taking out backprojection blurring to produce a clear reconstruction [14]. The process is iterated for non-central planes, even though the projection data are insufficient for exact reconstruction. In order to complete the projection data, one would need to move the scanning cone of X-rays away from the purely circular trajectory in some manner, or to move the patient. The limitations of the FDK approach result in blurring in the axial direction wherever there are inhomogeneities [6]. High density bone can produce streaking in surrounding soft tissues, the magnitude of which, for the Elekta Synergy system (Elekta, Crawley, UK), can be judged by inspecting the figures appearing in this paper. In so far as FDK artefacts alter grey-level gradients, this adds some uncertainty to boundary identification and contouring. However, this has not been found to affect cross modality matching for 3D setup purposes, even in the head and neck regions. By comparison, the partial volume effects seen in conventional CT can be large with slice widths of a few millimetres, although this is often not fully appreciated. Blurred, low contrast reconstruction due to anatomical motion during the slow process of profile acquisition is a significant problem facing CBT in radiotherapy, especially when that motion deviates from periodic behaviour. Add to this the effects of cone beam scatter and it can be seen that the practical limitations to CBT are at least as important as those arising from inexact reconstruction. FDK CBT is a practical proposition in external beam radiotherapy where rotating gantries suitable for circular scanning are the norm. The integration of CBT and medical linac technology has clearly been influenced by the work of Mosleh-Shirazi et al [7] and Jaffray et al [8]. Elekta’s Synergy clearly reflects the latter and has the cone beam X-ray volume imaging (XVI) system mounted on a linac gantry at 90˚ to the treatment beam with extending, twin pillar supports for the X-ray tube and an extending fork mount for the folding panel imager. Varian’s On Board Imager (OBI; Varian, Palo Alto, CA) also has its cone beam system perpendicular to the treatment beam, but uses folding robotic arms for both the X-ray tube and imager. In contrast, Siemens’ Artiste (Siemens, Erlangen, Germany) comes closest to providing a BEV by using a 180 ˚ design based on a kilovoltage source firing directly towards the treatment head in front of which the panel imager is deployed. Given the similarities of the integrated systems, the practicalities of CBT in the treatment room inevitably revolve around common factors; mechanical aspects, image processing and doses.

The mechanics of CBT The lessons learned implementing CT remain invaluable, even though at first sight the refined, purpose built S67

C J Moore, A Amer, T Marchant et al

machinery in the form of the CT ring mounting bares no comparison with a bulky linac gantry to which a radiographic unit has now been added for CBT. The IEC requirements for safety limit full linac gantry rotation are asked to be a minimum of 60 s [9], which makes CBT in the treatment room at least as slow as the first clinical CT scanners. Flex in the combined structures of linac and radiographic imaging system means that the central rays from the kilovoltgage source do not cross at a single point in space during the rotation fluoroscopy that feeds CBT reconstruction. X-ray projections of an optimally centred ball-bearing test object trace out a track on the aSi transducer plane that can be resolved into orthogonal curves termed ‘‘flex maps’’ that are quite specific to the machine being investigated [10]. The flex map pairs allow kilovoltage X-ray image projections to be re-aligned prior to filtered back-projection in CBT. For the Synergy system installed at the Christie Hospital, Manchester, extended monitoring has shown that the pattern and range of mechanical flexing in a large IGRT system is reproducible and stable over months of clinical use. Flex is periodic with peak-totrough amplitudes that are usually less than two pixels (,1.5 mm) in the plane of the flat panel imaging transducer. The kV and MV isocentres and their systematic difference can be accurately set to , 0.5 mm, which is now an action level for re-evaluating flex maps. Maintenance operations also trigger reevaluation. The reason why is illustrated by Figure 1a, which shows a flex map pair produced in July 2004 and then again in September 2004 following work on the positioning micro-switches of the imaging panel. Although the flex maps measured along the panel’s Vaxis (gun target direction) are the same shape, they are clearly shifted by nearly a pixel. After re-evaluation, flex maps are much more reproducible and repeatable than shown in this example. Figure 1b shows the isocentre stability, which is assessed weekly. Over a period of approximately 1 year the isocentre reached the 0.5 mm action threshold only twice. These and two mechanical maintenance events prompted flex map re-evaluations. For 90 ˚ mounts, the retraction of the kilovoltage imaging system is necessary for easy access to the patient on the treatment couch. Clinical use requires repeated deployment of the X-ray tube to a single, fixed position. On our system, the unfolding and offsetting of the transducer panel for CBT imaging is to one of three fixed positions, determined by the choice of a nominal 26 cm, 41 cm or 51 cm reconstruction field. Isocentre tests allow the reproducibility of deployment to be monitored, since any drift in panel translation will be revealed as a shift in the isocentre position on the imaging transducer array. Ideally, the normal to the plane of the imaging panel also remains normal to the gantry rotation axis. In addition the rows of transducer elements in the aSi array should remain parallel to the scanning axis throughout gantry rotation. Any deviations from this ideal need to be measured and taken into account during CBT reconstruction. Figure 2 shows the effect of 0.6 ˚ panel skew in the plane of the imaging panel. This was discovered by simply imaging a suspended engineering-standard steel ruler and noting that the projected edges were stepped rather than smooth, due to diagonal crossing of the rows of discrete S68

Figure 1. (a) Flex map pairs taken 2 months apart. U and V describe the projection of a centred ball bearing resolved in the directions of the major orthogonal axes of the imaging panel. Deviations are shown in pixels, at a scale of 0.5 mm pixel21. Flex maps peaking with the two highest offsets (0.0 and 0.7 pixels) were measured on 19 July 2004 and the remaining two were measured on 15 September 2004 (lower right hand key). (b) Isocentre calibration results for a 1 year period. Variation to the threshold level occurred twice (right hand vertical lines) triggering flex map re-evaluation. Mechanical maintenance triggered precautionary flex map re-evaluation on a further two occasions (left hand vertical lines).

transducer elements [11]. Comprehensive geometric calibration of cone beam systems and the development of test phantoms have been described by Smekal et al [12] and Cho et al [13]. Skew produces an obvious loss in apparent bone density, the blurring of otherwise fine detail and the possibility of double edging in bony anatomy. The principle of using image processing algorithms to automatically identify the degree of skew and optimally correct for its effects has been described by Amer et al [11]. However, skew has only been encountered once in the 3 years that Synergy has been installed at the Christie Hospital.

Image data processing in CBT The aSi panel detectors in commercial CBT are indirect detection systems. Light produced in a thin layer of luminescent material is captured efficiently by contact The British Journal of Radiology, Special Issue 2006


Figure 2. (a) The result of 0.6 ˚ in-plane, image transducer skew on cone beam tomography (CBT). (b) Results after skew correction. Note the improved contrast and fine detail.

interfacing to a silicon diode array. When deployed the arrays must be sufficiently large to image typical treatment volumes. Hence, it is not uncommon to find compound aSi panels that are made from two transducer arrays. In conventional CT the behaviour of a few thousand, high efficiency, solid state crystal detectors is charted. In CBT, flat panels have 2D arrays typically providing 102461024 transducer elements that must be calibrated for detector flaws, offset and gain. The manufacture of the imaging panel is an imperfect process and so redundancy is inbuilt. Nevertheless, unresponsive and unreliable, or ‘‘bad’’, transducer elements remain that must be identified and entered into a bad-pixel map in order to avoid their use in CBT back projection, where they would produce ring artefacts. The aSi panels used at the Christie are Perkin Elmer RID16401,048,576 gadolinium oxysulphide transducers (Perkin Elmer, Santa Clara, USA). They provide a compound, rectangular transducer area of 41 cm641 cm operating at a sampling rate of 3 Hz. The gadolinium oxysulphide panel has a measured bad pixel population of several hundred unresponsive pixels and several thousand unreliable pixels, locally termed ‘‘transients’’. Figure 3 illustrates their distribution and instability. More efficient caesium iodide panels, working at 6 Hz, are available. Not surprisingly reconstruction by filtered back-projection is sensitive to small variations in the signals emerging The British Journal of Radiology, Special Issue 2006

from the flat panel transducer array, whether they are due to real variations in exposure in an image profile or defects. A 1% local variation is enough to produce obvious effects in CBT reconstruction. Hence, in addition to bad pixels, it is necessary to correct for the inherent structure generated by the panel imager whilst acquiring X-ray image profiles. In this context, the panel dark-field, an offset image seen in the absence of X-rays, introduces a great deal of unwanted irregular structure. Figure 4a shows an example. In practice, it is now known that sampling the offset image close to the time of clinical profile acquisition greatly improves CBT image quality. Flood field (gain) calibration imaging is the other side of profile correction. It allows the anisotropic, non-linear exposure responses of pixels across the flat panel to be taken into account. Figure 4b is an example of the unwanted, regular panel structure seen in high exposure gain images. Rather reassuringly, the authors have found offset and gain calibration of X-ray image profiles to be eminently practical and effective. Bow-tie filters can be used to physically equalize exposure across the volume of view, though these are not used in the examples presented in this paper. A useful side-effect is the reduction of low energy X-rays and reduction in skin dose. CBT is essentially a process of carefully calibrated rotation fluoroscopy followed by reconstruction to produce an X-ray image volume, most commonly using S69


Figure 4. (a) Mean dark field (offset image) computed from Figure 3. Bad pixels, based on their variability, on an amorphous silicon imaging panel. (a) 7000 entries with standard deviation.3 on 1 day. (b) 2000 entries with standard deviation.3 over 5 days.

the FDK filtered back projection algorithm. The image volume must be viewed with at least two very basic, but often ignored, rules from diagnostic radiology in mind. The first is to interrogate the CBT volume image in the most appropriate plane, which may not be the transaxial plane that most radiotherapy observers are conditioned to select by virtue of their experience with axial RTP-CT images in treatment planning systems. Coronal and sagittal views of RTP-CT image data can be highly discontinuous because axial slice thicknesses of several millimetres have been the de-facto standard for decades. CBT image volumes are not inherently slice based and S70

50 sample frames for a gadolinium oxysulphide aSi imaging panel. The axes indicate that the panel can provide up to 100061000 pixels per image frame. (b) Mean flood field (gain image) computed from 50 sample frames for a gadolinium oxysulphide aSi imaging panel and high exposure.

the spatial resolution is the same along all three Cartesian axes. Hence, CBT coronal and sagittal sections can provide exquisitely resolved detail that is ideal for structure delineation almost free from significant partial volume effects. The top two rows of Figure 5 amply illustrate the point. The second rule is to compare like with like when matching features visible in RTP-CT and CBT in IGRT. A natural tendency is to compare axial CBT images, in their raw state, perhaps equivalent to 1 mm slice thickness, with RTP-CT axial slices that might be several times thicker. Averaging CBT data to produce The British Journal of Radiology, Special Issue 2006


Figure 5. (a) Radiotherapy planning (RTP)-CT scan, 5 mm slice thickness. (b) Cone beam tomography (CBT) using 383 image profiles, 1 mm slice thickness. The reconstruction is consistent with the theoretical minimum number required for exact reconstruction. (c) CBT using 148 image profiles, 1 mm slice thickness. (d) CBT using 77 image profiles, 1 mm slice thickness. Note the profile aliasing crossing the CBT axial.

an equivalent slice width to that found in the RTP-CT scan will increase soft tissue contrast–noise ratios and produce equivalent partial volume effects. Figure 6 illustrates the process for pelvic scans. CBT reconstruction is underpinned by rotation fluoroscopy, which when replayed offers useful insight into patient dynamics, much like the conventional simulator. Although the fluoroscopy trajectory is prescribed for CBT, there is a particular advantage over conventional simulation. Where the gross tumour volume (GTV) or organ at risk (OAR) structure, surgical clips or other suitable landmarks are clear, then the fluoroscopy sequence can be interrogated to assess the amplitude and direction of movements, since the average spatial location (and hence depth), from any particular gantry angle, is known from the CBT reconstruction itself. The British Journal of Radiology, Special Issue 2006

Since the rotation fluoroscopic sequences and CBT reconstructions are both clinically useful, their storage, communication and archiving is an issue that must be adequately addressed. In Manchester, this has been done by installing additional dual storage PCs on the clinical network. Both have terabyte disk capacity and gigabyte memory for multiple image volume display and manipulation. A multi-tape archiving system with matching capacity has also been purchased.

Doses in CBT With high patient exposures and a large number of X-ray image profiles for reconstruction, it is possible to generate high quality CBT images. However, a principle S71


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(c) Figure 6. (a) Radiotherapy planning (RTP)-CT images, 5 mm thick should be compared with (b) the equivalent cone beam tomography (CBT) images, 5 mm equivalent slice thickness by processing (c) the raw CBT data, 1 mm thick gathered at treatment time. In this example, note the changes in rectal and bladder fill.

that supports the practice of diagnostic radiology in the UK is that images produced using ionizing radiation need only be fit for purpose. In particular, the IRMER legislation [15] embodies the principle of doses being ‘‘as low as reasonably practical’’ (ALARP) in the context of risks and benefits to the patient. For radiotherapy there are two key applications of serial CBT; patient setup, which can be achieved using low dose imaging, and GTV/OAR imaging, which at first sight demands higher doses than those for patient setup in order to provide the necessary contrast detail. RTP-CT scanning outside the treatment room sets a benchmark weighted dose (two-thirds of the surface dose plus one third central dose in a cross section, used throughout this paper unless otherwise stated) of approximately 1–2 cGy, which is in line with megavoltage portal imaging in the treatment room at 1–2 cGy per image (maximum dose). Wide field imaging at these dose levels has been acceptable in radiotherapy because its use has been restricted to the early stages of the treatment S72

planning and delivery process, where the correction of systematic setup errors, at least in terms of bony anatomy if not the tumour itself, is likely to have maximum benefit assuming the absence of time trends. This message is reinforced by radiobiological modelling studies [16]. Nevertheless, setup based on wide field orthogonal portal images for only 5 treatment days can lead to a cumulative dose of 10–20 cGy. With the appearance of CBT in the treatment room it is practical to acquire images for bony setup at every fraction with the added potential benefit of GTV/OAR visualization (see the next section). Serial use of cone beam reconstruction techniques delivering approximately 5 cGy per image can result in patient doses that approach the 1 Gy level [17, 18]. In the context of IRMER, these are approaching cautionary levels where deterministic effects of low energy kilovoltage X-ray on the skin might become visible across large areas. They have already added fuel to the debate on induced secondary cancers [19, 20]. The British Journal of Radiology, Special Issue 2006


The distinction between CBT and conventional CT is purely technical; patient dose is due to kilovoltage X-ray interactions with tissue, and so it is appropriate to make use of the existing framework for describing dose levels that is already established in diagnostic radiology. The CT dose index (CTDI) is the 2:1 weighted measure of skin and central dose that forms the basis of the European dose reference levels (EDRL) for different anatomical sites [21]. EDRL values range from 30 mGy to 60 mGy moving from the head to the pelvis. Amer et al [22] have reported typical CBT dose levels for head and neck (1.6 mGy using 100 kVp, 25 mA, 4 ms profiles), lung (6.0 mGy using 120 kVp, 40 mA, 10 ms profiles), breast (3 mGy using 120 kVp, 40 mA, 5 ms profiles) and pelvis (25 mGy using 130 kVp, 40 mA, 30 ms profiles). These are based on the Christie Hospital’s approach of using a minimum set of approximately 400 profiles for exact CBT reconstruction to a 512 cubic image matrix. Serial CBT over 20 fractions would then result in total doses of between 2 cGy and 50 cGy, depending on the anatomical site. Given the benefits of target visualization these might be justifiable. However, the technique of zonal CBT reported by Moore et al [23] has the potential to simultaneously improve image contrast in the target zone whilst halving doses outside the target zone, making serial CBT throughout treatment no more expensive in terms of dose than short term EPID use.

The clinical role of cone beam in radiotherapy The parallels between CBT and conventional CT are by now very clear. However, there are some differences that reflect the intended application of CBT in the radiotherapy treatment room. For example, the Christie Synergy is an IGRT facility intended for patient setup and tumour targeting rather than the production of images with grey scale values calibrated for use in treatment planning. Nevertheless, there is clear potential for adaptive planning with bulk inhomogeneity corrections and the long scanning periods are reminiscent of the scanning protocols used in the early days of RTP, when averaged patient motion was viewed positively. Varian’s cone beam promotional material (http://www.varian.com/orad/ pdf/radonc.pdf) illustrates how adaptive planning of patients might be performed using cone beam volume images taken on the day of treatment. The clinical applications of CBT in radiotherapy deserve closer scrutiny since the expectation is that patient setup, disease targeting and possibly treatment planning can be used frequently, if not daily, over the course of a treatment. The liberal use of the wide field CBT exposures could drive concomitant patient doses to levels associated with North American medical practice. This will be avoided by advancing clinical practice in the context of the European legislation that protects the interests of the patient and embodied as IRMER in the UK [15].

Patient setup Kilovoltage CBT is an obvious replacement for megavoltage EPIDs in patient setup, where 2D portal images are matched against digitally reconstructed radiographs The British Journal of Radiology, Special Issue 2006

(DRRs) derived from RTP-CT scans. The latter amounts to the matching of spatially continuous, low contrast physical density images to superposition images derived from sequential axial sections, which have high contrast by virtue of their photoelectric origins. Since CBT produces high contrast 3D image volumes with equally high spatial resolution along all axes, it can support translational and, given appropriate couch technology, full rotational correction of patient setup by direct comparison with RTP-CT scan volumes. RTP-CT protocols have evolved to match the resolutions of MLCs in wide field conformal therapy, which started with leaf projections of order 10 mm at isocentre. Hence, they focused on the production of matched CT sections spaced at intervals of 10 mm. More recently this has become 3–5 mm along the key superior–inferior (SI) axis parallel to the portal image plane. In turn, the crude DRRs generated from these sections for setup verification purposes were consistent with the relatively poor performance of early EPIDs in the treatment room, some of which could only offer 4 mm resolution. CBT in the treatment room can now offer better than 1 mm spatial resolution in the SI direction. Hence one would expect this to stimulate the increased use of matching, 1 mm resolution multi-slice and spiral CT in conformal treatment planning. This will improve not only field shaping but also setup verification. CBT ushers in an era where the manual matching of selected features throughout two almost contiguous image volumes is far from easy. In addition, unambiguous 3D visualization through solid data is beyond current display technologies. However, computer assisted volume matching has already reached a degree of maturity. Using the Philips Syntegra package running on a Pinnacle treatment planning system manual (Philips Pinnacle, Milpitas, USA) and automatic matching of CBT to RTP-CT have been compared by Marchant et al [24]. This work concentrated on the pelvis, showing the equivalence of manual matching using bony detail and automated local correlation matching based on all the available CBT grey scale data. This result is not entirely unexpected, given the qualitative similarities of the human visual response function and the mathematics of correlation, which are greatly influenced by discontinuities in image data. However, it does highlight the fact that, so far as radiotherapy setup is concerned, high integrity grey scale reproduction in CBT may not be required and the need for many hundreds of X-ray image profiles to support reconstruction is questionable. An indication of the minimum number of image profiles needed for optimum CBT reconstruction is readily established by considering the circular scanning trajectory and approximately cylindrical reconstruction region of interest (RROI) in commercial systems. For an RROI with base diameter and height N pixels, there are pN3/4 volume elements to be characterized from a set of square X-ray image profiles, each representing N2 measurements. Simple division shows that pN/4 (0.785N) image profiles are required to solve exactly for the attenuation at every point in the RROI. For N5512, this suggests the acquisition of at least 402 X-ray image profiles or roughly one per degree for a full gantry rotation. To achieve this in a minimum 60 s scan requires the image transducer to sample at approximately 7 Hz, S73


which is well within the performance of systems now offered by manufacturers. More profiles and more patient dose will improve image quality. However, guided by the ALARP principle, Sykes et al [25] at the Christie Hospital reported head phantom studies in which CBT reconstruction based on fewer than 100 image profiles was sufficient for matching to a high quality RTP-CT scan with sub-degree, sub-millimetre accuracy. Figure 5 illustrates the effect of reduced profile CBT reconstruction for a patient. Soft-tissue detail remains visible down to very low profile counts, despite the appearance of aliasing streaks crossing the RROI. Reduced profile CBT reconstruction for setup offers the prospect of containing or even reducing patient dose over wide fields of view. Reducing profile collection in turn suggests that faster scanning is a possibility and revisiting gantry design with a view to influencing IEC rotation limits could be beneficial.

Disease targeting Targeting requires sufficient soft-tissue contrast detail at the disease site in order to define the position and shape of what can be termed the GTV of the day. The same is true for adjacent OAR, which may be part of a locally interacting configuration of diseased and healthy anatomy. This suggests high CBT exposure levels and large numbers of image profiles must be collected to improve the signal to noise ratios in reconstructions. Where there is a moving target, as in the lung, the possibility of motion reduction through phase selection and reconstruction using 4D CT techniques, as demonstrated by Sonke et al [26] for CBT on a Synergy system, can increase the need for profiles still further. The result is significantly increased X-ray tube loading. Such considerations have influenced the production versions of commercial IGRT systems, including Elekta’s Synergy, which has an increased capacity X-ray tube compared with the prototype. Improving soft tissue contrast in CBT is complicated by the appreciable scatter field produced by a wide field cone beam as it passes through the bulk of the patient. Anti-scatter grids at the image transducer plane have had some success. However, ‘‘coning down’’ potentially restricts the use of CBT in patient setup. This has prompted the use of the EPID dual exposure approach by Letourneau et al [17] for CBT, where a wide RROI CBT for setup purposes and a coned down ‘‘local tomography’’ RROI are superimposed to improve target visualization. Technique development is clearly influencing the time needed for CBT investigations and increasing the dose to the patient. However, having already noted that reduced profile, reduced dose CBT can support accurate setup, Moore et al [27] have reported a simple ‘‘zonal CBT’’ technique to reduce scatter at source in the patient, improve image contrast–noise ratios in the target zone and simultaneously reduce patient dose compared with a full field technique. They use machined aluminium transmission diaphragms attached to the kilovoltage X-ray tube of an Elekta Synergy. These operate independently of the usual lead diaphragms, which are left fully open. During CBT, dose to the tissues lying beneath the S74

transmission diaphragms is reduced by a factor of two or more, according to the thickness and composition of the diaphragms. In turn, the scatter from these tissues is correspondingly reduced. Between the transmission diaphragms, an unattenuated X-ray beam is directed towards the target zone, which now receives reduced scatter and hence lower dose from the extrazonal tissues. The early indications are that isocentre dose levels are readily halved by this technique. The result is improved soft-tissue contrast where it is needed, at the target, and sufficient edge detail elsewhere for wide field setup using Syntegra. Figure 7 shows a bladder patient scanned using zonal CBT with contrast enhanced target detail and useful soft tissue contrast well outside the enhanced target zone.

Treatment planning EPID targeting is based on bony setup comparison with the RTP-CT scan, where the systematic errors are likely to be trend free over several days, if not weeks. With CBT it is at once obvious that the GTV, OARs and impacting soft-tissues, such as bowel, execute movements that are not necessarily random or small scale. In addition to shape change, for example in bladder filling, the target can show positional time trends, as in rectal cases. Unexpectedly large internal displacements have been observed using CBT in cervix cases. Given CBT volumes of sufficient quality, a method of rapidly delineating targets and computer assisted movement of the treatment couch, there is the potential to perform fast adaptive planning. Armed with knowledge from CBT studies, an alternative has been examined for bladder patients at the Christie Hospital [28]. In this approach setup errors are addressed and then three quantized margins are added in turn to the bladder CTV seen on the RTP-CT scan to generate a set of three PTVs. The PTVs are designed to encompass all the bladder movements that might be seen in a population and are planned in advance of treatment. Since the greatest changes to the bladder are usually seen in the superior direction, the superior margin ranges from 5 mm to 15 mm in 5 mm increments. The anterior margin is set at 15 mm and all other margins at 10 mm. During treatment a CBT scan is used for edge based registration to the RTP-CT scan and the best margin fit to the bladder seen in the CBT is chosen from the three available plans. However, CBT has shown that there are occasions when unexpected deformations thwart even the best laid plans. Figure 8 shows a CBT coronal section of a bladder patient displayed against the primary planning data within the Pinnacle RTP system. Point by point dosimetric modelling requires cone beam reconstructions with grey levels that are a close approximation to the CT numbers found in conventional CT planning scans, despite minute long scanning times. This is particularly demanding given the scatter from a wide field, minimally collimated, kilovoltage X-ray cone beam. From the data given previously for the Synergy system at the Christie Hospital it can be seen that the cone beam imaging angles range from 15 ˚ to 30 ˚, which is much larger than in diagnostic multislice CT scanners. Furthermore, the transducers in the multislice scanners The British Journal of Radiology, Special Issue 2006


Figure 7. Zonal cone beam tomography (CBT) of a bladder patient reduces scatter from tissues outside the target zone, thus improving target contrast whilst simultaneously preserving edge structures for setup and reducing patient dose. (a) CBT coronal section showing improved target zone contrast (arrow). Note the preserved soft tissue detail above and below. (b) 1 mm axial image detail within the target zone. (c) 1 mm axial image detail outside the target zone.

differ markedly from the thin, indirect detection panels commercially available for radiotherapy. Clearly the development of superior image transducers will help.

Cone beam in transition The Christie Hospital is using kilovoltage CBCT for patients with disease of the central nervous system, head and neck, breast, lung, upper gastrointestinal tract, rectum, bladder, prostate, cervix and peripheral soft tissue sarcomas. At the time of writing approximately

1000 CBT scans have been completed on a single preproduction machine. CBT is being deployed to support IMRT, including individuals selected from the large pool of breast patients seen at the Christie Hospital each year. The Synergy treatment head has a touch guard that is 43 cm from the isocentre, which is closer than the CBT panel imager at nearly 54 cm. Hence, the equivalent ‘‘bore’’ for CBT scanning is 86 cm, which is comparable with widest bore CT-simulators. This provides sufficient space to acquire whole body scans for most treatment techniques. Figure 9 shows a wide field lung scan without motion gating. However, the geometric limits to cone beam scanning are reached with breast patients who lie supine on a breast board that can be inclined to 20 ˚ with the patient’s arm raised and immobilized. For these patients it is not possible to perform a full rotation scan to generate a 50 cm, whole body RROI. Instead, a half scan is performed with the patient shifted vertically from the treatment centre. This produces a field of view optimally encompassing breast, heart, lungs and lateral body surface. The shift is recorded and transferred to the hospital’s Pinnacle planning system where cross modality comparison with planning scans is performed, assisted by common colour overlays of radiotherapy structures, as in Figure 10.

The future

Figure 8. Coronal plane cone beam tomography (CBT) of a bladder, registered with the radiotherapy planning (RTP)-CT scan (inset) showing the displacement and deformation of organs in the target zone due to unexpectedly large rectal distension. The central graphic defines the clinical target volume (CTV) and the outer graphic the planning target volume (PTV). The British Journal of Radiology, Special Issue 2006

A particular concern of the authors has been that cone beam imaging in radiotherapy looked set to become an additional process executed before or after treatment. Hence, the configuration of target and OARs during the process of beam delivery itself would not be known. To address this problem, we have collected X-ray projections during the (segmented) gantry rotations between the beam deliveries of a single treatment fraction, and assembled these partial profile sequences into a contiguous data set for CBT reconstruction. The results have been encouraging. Figure 11 demonstrates the comparable quality of CBT reconstructions generated from S75


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Figure 9. Matched radiotherapy planning (RTP)-CT and cone beam tomography (CBT) image volumes on the Pinnacle treatment planning system. (a) RTP-CT axial and coronal, 5 mm sections with inner clinical target volume (CTV) and planning target volume (PTV). (b) CBT axial and coronal, 1 mm sections with registered CTV and PTV. (c) CBT axial and coronal, 5 mm equivalent sections for comparison.

continuous and between beam acquisitions. An added bonus comes in the form of the potential for quantification of patient shift during treatment using measurements from the reconstructed bony detail in the cone beam image volume itself. Algebraic reconstruction algorithms [29] have already been researched as alternatives to FDK filtered back projection. Their versatility in the face of scatter and their potential to produce reconstructions with CT-number accuracy makes their practical development attractive. Furthermore, they marry well with the out of plane collection of projection data that improves the information content beyond the central reconstruction plane. Nutation of the X-ray source and image transducer during the execution of an otherwise circular scanning trajectory would allow the extra data to be gathered.

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The value of taking projections from trajectories other than a strict circular rotation suggests that integration of the kilovoltage imaging system with the treatment gantry may be too restrictive. Indeed, as long as it is possible to coregister the cone beam imaging device with the local treatment machine coordinates there is no absolute reason for gantry integration. Furthermore, freed from the restriction on speed of therapy gantry rotation, a separate cone beam device could offer fast scanning that would improve work flow and reduce motion artefacts in the reconstructed image volumes. It would also avoid treatment machine down time caused by maintenance of the imaging system. Solberg et al [30] have reported their assessment of such a device from Siemens (Erlangen, Germany). This technology is a mobile, kilovoltage C-arm device that could be shared across a clinical department.

(b)

(c)

Figure 10. Comparison of breast simulator-CT and cone beam tomography (CBT) in the treatment room. (a) Transverse sim-CT section with breast and lung contours. (b) Treatment room CBT axial (1 mm equivalent slice thickness) overlaid with simulation contours. (c) CBT coronal section (1 mm equivalence slice thickness). S76



Figure 11. (a) Cone beam tomography (CBT) coronal image generated from profiles gathered between treatment beam deliveries. (b) CBT coronal image for the same patient generated from image profiles gathered in a single continuous posttreatment scan.

Conclusions X-ray CBT for IGRT looks certain to expand into mainstream clinical use because it allows 3D setup and soft tissue visualization that is both practical and superior to the alternative megavoltage modalities. Imaging technology in the treatment room is now at least as capable as that deployed for treatment planning, taking into account the large 3D contiguous image volumes that can be produced. Frequent deployment of CBT in support of patient setup, disease targeting and adaptive treatment planning is unlikely to be limited by technology or dose considerations. It appears to be only a matter of time before the multiway flow of reconstructed volumetric images between planning and treatment departments will require a supporting infrastructure every bit as impressive as that already found in diagnostic radiology.

References 1. Feldkamp LA, Davis LC, Kress JW. Practical cone beam algorithm. J Opt Soc Am 1984;1:612–9. 2. Lewis DG, Swindell W, Morton EJ, Evans PM, Xiao ZR. A megavoltage CT scanner for radiotherapy verification. Phys Med Biol 1992;37:1985–99. 3. ICRU Report 50. Prescribing, recording and reporting photon beam therapy. Washington, DC: International Committee on Radiation Units and Measurements, 1993. 4. Antonuk LE. Electronic portal imaging devices: a review and historical perspective of contemporary technologies and research. Phys Med Biol 2003;47:R31–R65. 5. Antonuk LE, El-Mohri Y, Huang W, Jee KW, Siewerdsen JH, Maolinbay M, et al. Initial performance evaluation of an indirect-detection, active matrix flat-panel imager (AMFPI) prototype for megavoltage imaging. Int J Radiat Oncol Biol Phys 1998;42:437–54.


6. Turbell H, ‘Cone Beam Reconstruction using Filtered Back Projection’, Dissertation 672, February 2001, Linko¨ping Studies in Science and Technology, Department of Electrical Engineering, Linko¨ping University, Sweden. 7. Mosleh-Shirazi MA, Evans PM, Swindell W, Webb S, Partridge M. A cone-beam megavoltage CT scanner for treatment verification in conformal therapy. Radiother Oncol 1998;48:319–28. 8. Jaffray DA, Siewerdsen JH, Wong JW, Martinez AA. Flatpanel cone-beam computed tomography for image-guided radiation therapy. Int J Radiat Oncol Biol Phys 2002;53:1337–49. 9. European Standard EN 60601-2-1 Medical electrical equipment Part 2-1: Particular requirements for the safety of electron accelerators in the range of 1MeV to 50 MeV (IEC 60601-2-1:1998). 10. Fahrig R, Holdsworth DW. Three-dimensional computed tomographic reconstruction using a C-arm mounted XRII: image-based correction of gantry motion nonidealities. Med Phys 2000;27:30–8. 11. Amer A, Sykes J, Moore CJ. Imaging panel skew correction and autofocussing in radiotherapy cone beam imaging. Procs IEEE Computer Society, Procs VIIIth Information Visualisation, 2004 July: 92-97. 12. Von Smekal L, Katchelriess M, Stepina E, Kalender WA. Geometric misalignment and calibration in cone-beam tomography. Med Phys 2004;31:3242–65. 13. Cho Y, Moseley DJ, Siewerdsen JH, Jaffray DA. Accurate technique for complete geometric calibration of cone-beam computed tomography systems. Med Phys 2005;32:968–83. 14. Hsieh J. Computed tomography; principles design, artifacts and recent advances. Washington, USA: SPIE Press, 2003. 15. The Ionising Radiation (Medical Exposure) Regulations 2000 UK. Statutory Instrument SI 2000, No 1059. London: HMSO, 2000. 16. Amer A, Marchant T, Sykes J, Henry A, Mcbain C, Moore CJ. The potential of X-ray volume imaging in adaptive

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radiotherapy to reduce the small bowel high dose volume in bladder cancer treatment. ESTRO-23; 2004 24-28 October; Amsterdam. Radiother Oncol 2004;73(Suppl-1): Abstr-513, S230. Letourneau D, Wong J, Oldham M, Gulam M, Watt L, Jaffray D, et al. Cone-beam-CT guided radiation therapy: technical implementation. Radiother Oncol 2005;75: 279–86. Yamada Y. Image guided intensity modulated radiotherapy with the on board imager. ASTRO 46th Annual Meeting; 2004 October 3-7; Atlanta, USA. http://www.varian.com/ orad/prd183.html# [Accessed 28 April 2006]. Aird EGA. Second cancer risk, concomitant exposures and IRMER(2000). Br J Radiol 2004;77:983–5. Waddington SP, MacKenzie AL. Assessment of effective dose from concomitant exposures required in verification of the target volume in radiotherapy. Br J Radiol 2004;77:557–61. EC Quality Criteria for Computed Tomography. EC Working Document EUR 1626. Brussels EU, 1998. Amer A, Marchant T, Sykes S, Czajka J, Davies J, Stratford J, et al. Doses from cone beam CT integrated to a radiotherapy treatment machine. United Kingdom Radiation Oncology Conference; 2005 11-13 April; York, UK. Moore CJ, Marchant T, Amer A, Sykes J, Henry A. Zonal Xray volume imaging. ESTRO-23; 2004 24-28 October; Amsterdam, The Netherlands. Radiother Oncol 2004;73(Suppl-1): Abstr-111, S51.

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24. Marchant T, Sykes J, Amer A, Moore CJ, Stratford J, Mccarthy C, et al. Measurement of setup error using image registration of X-ray volume images (XVI) and treatment planning CT. ESTRO-23; 2004 24-28 October; Amsterdam, The Netherlands. Radiother Oncol 2004;73(Suppl-1): Abstr510, S229. 25. Sykes JR, Amer A, Czajka J, Moore C. A feasibility study for image guided radiotherapy using low dose, high speed, cone beam Xray volumetric imaging. Radiother Oncol 2005;77:45–52. 26. Sonke JJ, Zijp L, Remeijer P, van Herk M. Respiratory correlated cone beam CT. Med Phys 2005;32:1176–86. 27. Moore CJ, Marchant T, Amer A, Sykes J, Henry A. Zonal Xray volume imaging. ESTRO-23; 2004 24-28 October; Amsterdam, The Netherlands. Radiother Oncol 2004;73(Suppl-1): Abstr-111, S51. 28. Amer A, Marchant T, Sykes J, Henry A, Mcbain C, Moore CJ. The potential of X-ray volume imaging in adaptive radiotherapy to reduce the small bowel high dose volume in bladder cancer treatment. Radiother Oncol 2004;73(Suppl-1): Abstr-513, S230. 29. Mueller K, Yagel R, Wheller JJ. Anti-aliased 3D cone-beam reconstruction of low contrast objects with algebraic methods. IEEE Trans Med Imaging 1999;18:519–37. 30. Solberg TD, Chow PE, Sorensen S, Wink N, Mitschke MM, Ritter D, et al. Image guided radiotherapy using a mobile kilo-voltage X-ray device. ASTRO 46th Annual Meeting; Abstr 2459; 2004 October 3-7; Atlanta, USA.



Kilovoltage CT using a linac-CT scanner combination 1

C THIEKE, MD, PhD, 2U MALSCH, MS, 2W SCHLEGEL, R BENDL, PhD and 1C THILMANN, MD, MS

PhD,

1

J DEBUS,

MD, PhD,

1

P HUBER,

MD, PhD,

2

1

Department of Radiation Oncology, University Hospital of Heidelberg and German Cancer Research Center, Heidelberg and 2Department of Medical Physics, German Cancer Research Center, Heidelberg, Germany ABSTRACT. Modern radiotherapy techniques such as intensity modulation are capable of generating complex dose distributions whose high dose areas tightly conform to the tumour target volume, sparing critical organs even when they are located in close proximity. This potential can only be exploited to its full extent when the accumulated dose actually delivered over the complete treatment course is sufficiently close to the dose computed on the initial CT scan used for treatment planning. Exact patient repositioning is mandatory, but also other sources of error, e.g. changes of the patient’s anatomy under therapy, should be taken into account. At the German Cancer Research Center, we use a combination of a linear accelerator and a CT scanner installed in one room and sharing the same couch. It allows the quantification and correction of interfractional variations between planning and treatment delivery. In this paper, we describe treatments of prostate, paraspinal and head and neck tumours. All patients were immobilized by customized fixation devices and treated in a stereotactic setup. For each patient, frequent CT scans were taken during the treatment course. Each scan was compared with the original planning CT using manual checks and automatic rigid matching algorithms. Depending on the individual case, the adaptation to variations was carried out offline after several fractions or in real-time between the CT scan and linac irradiation. We discuss the techniques for detecting and correcting interfractional errors and outline the procedural steps of a linac-CT scanner-supported radiation treatment course.

CT has been used in radiotherapy for decades and plays a decisive role in the development of conformal radiation treatment. Computer science, physics and engineering have had a tremendous impact on radiotherapy. The most obvious progress has been achieved in treatment planning and treatment delivery. Here, we are close to the limits of what can be reached with high energy photons. Modern irradiation techniques such as stereotactic radiotherapy and intensity-modulated radiotherapy (IMRT) are capable of generating complex dose distributions whose high dose areas tightly conform to the tumour target volume. Exact repositioning of the target is mandatory to ensure that the dose actually delivered at treatment time is as close as possible to the dose computed on the initial CT scan used for treatment planning. Despite great efforts, this is the weak point of high precision radiotherapy especially in the trunk region. Available fixation devices reduce interfraction and intrafraction motion. For example, our in-house developed combination of a body cast and head mask system in a rigid stereotactic body frame ensures noninvasive patient fixation for fractionated extracranial stereotactic radiotherapy and IMRT [1]. It provides precise and reliable positioning with regard to bony structures [2]. Nevertheless, internal soft tissue structures may deviate considerably from what we expect. When treating internal targets in the body region, the target position cannot be guaranteed with external The British Journal of Radiology, Special Issue 2006


immobilization devices. No further improvement of repositioning can be achieved by more rigid fixation. Additionally, a CT examination for treatment planning is a snapshot of anatomical structures and is gathered several days before treatment. Especially in high conformal radiotherapy with dose escalation and conformal avoidance of critical structures requiring a high degree of positional accuracy, the daily position of the target needs to be confirmed before irradiation by a reliable imaging modality. There are different approaches to image guidance aimed at the reduction of uncertainties of target position. For acquisition of 3D anatomic information in the treatment position, megavoltage CT [3] and cone beam kilovoltage CT fixed at the gantry [4] or a separate inroom CT scanner [5–7] has been tested so far. When using an in-room CT, the system consists of a linear accelerator (linac) and a conventional CT scanner connected via a conventional treatment couch. The CT scanner is mounted on rails. The obvious advantage of this system is that all components are separately established for clinical application. The accuracy has been demonstrated by Cheng and coworkers [8]. We would like to point out three particularities of the system. (1) The patient is immobilized in the usual way on the treatment couch and CT scanning is performed for confirmation of the correct isocentre and patient (target) position. By using a gantry mounted on rails, scanning is S79

C Thieke, U Malsch, W Schlegel et al

performed without either couch or patient movement, eliminating reductions in accuracy. After performing the scan, the treatment couch is rotated to the linac side for irradiation. (2) The kilovoltage technique allows an image quality with high soft-tissue contrast. This is a requirement for detecting deviations of soft tissue targets from their position at the initial CT scan used for treatment planning, even when bony landmarks are repositioned correctly. (3) The CT data set can be used without any transformation for treatment planning and dose calculation. The basic input for radiotherapy planning systems considering tissue heterogeneities is the relationship between CT Hounsfield units and electron densities, which is determined using common CT-calibration methods. Thus the in-room CT is well suited for target point verification, correction of setup errors and interfraction target deviations due to organ motion, as well as for recalculation of the dose actually given. The simplest correction is to correct the target point without changing the treatment plan. Beyond this, systematic changes of the patient’s anatomy under therapy (e.g. weight loss, tumour mass reduction) can be considered and, if necessary, the radiation treatment plan can be re-optimized until the next fraction. The highest level of adaptation achievable with a linac-CT combination is the adaptation/optimization of the treatment plan to the actual given situation, i.e. between the CT scan and the linac irradiation. Since a CT scan is not possible in treatment position, the described system cannot be used to detect intrafraction motion. The purpose of this study is to demonstrate the potential of an integrated linac-CT scanner system for different tumour sites and to derive a workflow which is reasonable for clinical routine with target point correction and re-optimization of the dose distribution.

Materials and methods Treatments at DKFZ with the linac-CT scanner Siemens Primatom Since October 2002, the Siemens Primatom (Siemens OCS, Concorde, CA) has been in clinical use at the German Cancer Research Center (DKFZ) in Heidelberg, Germany. It consists of a 6 MV Siemens Primus linear accelerator and a single-slice spiral CT scanner Siemens Emotion, which are set 90 ˚ apart and share the same couch (Figure 1). Depending on the treatment room, a 180 ˚ design can also be used [6, 8]. Almost all fractionated treatments at DKFZ are carried out as IMRT. The indicated treatments are prostate cancer, head and neck cancer, paraspinal tumours, base of skull tumours, breast cancer and, more recently, pleural mesothelioma, oesophageal cancer and pancreatic cancer. Radiosurgery is used for arteriovenous malformations, lung tumours and liver metastases.

Figure 1. Treatment room with the Siemens Primatom at DKFZ Heidelberg. The linear accelerator and the CT scanner on rails share the same couch, which is rotated by 90 ˚ to switch between CT scanning and treatment position.

therefore cast is produced [9]. For intracranial radiosurgery we use either an invasive ring or the Scotch (3M, St Paul, MN) cast mask. For extracranial targets, we use either a wrap-around body cast [2] made from the same material as the head mask or a vacuum pillow [2]. Both extracranial fixation devices are complemented by a head mask to eliminate head rotations, which might translate into movements of the spine, and for single-fractionated treatment of lung and liver an additional abdominal pressure plate reduces the respiratory motion.

Treatment For all single-fractionated treatments, a control CT scan is performed directly prior to irradiation. By comparing the control CT scan to the CT scan made for treatment planning, any translational displacement is detected and the target point corrected accordingly. For fractionated treatment of intracranial targets, the repositioning accuracy of the patient’s head with the customized head mask was found to be better than 2 mm in all three dimensions [9]. We usually therefore include a 2 mm safety margin in the planning target volume (PTV), and apart from the verification of the target point on the first fraction, we do not perform further control CT scans. Fractionated treatment of extracranial targets is accompanied by frequent CT control scans during the treatment course, since even with a rigid fixation significant interfractional discrepancies are observed (see results below). The scan frequency depends on the individual case and is between once weekly for unproblematic cases, up to daily CT scans for patients with large repositioning errors and critical proximity of target structures to organs at risk.

Patient fixation Every patient at DKFZ is treated in an individually customized fixation device. For fractionated treatment of intracranial targets, a head mask made of Scotch S80

Analysis of CT scans All CT scans generated with the Siemens Primatom are performed with stereotactic localizers attached to the The British Journal of Radiology, Special Issue 2006

kV CT using a linac-CT scanner combination

fixation frame in exactly the same way as they were attached in the treatment planning CT scan. This defines a frame-based coordinate system independent of the patient’s actual position, which is a prerequisite for any further analysis. For the first fraction, lead ball bearings (BB) were additionally put onto the laser adjustment lines for target point verification, so no further portal imaging was necessary. When we started using the Siemens Primatom, the correct positioning of the patient could only be verified by manually highlighting representative anatomical landmarks on selected CT slices. For example, about three bony landmarks on the isocentre plane were predefined on the planning CT in the treatment planning system and then located on the control CT using the CT scanner console. The average coordinate difference between the scans then gave an estimation of the average displacement error. Although technically feasible, this procedure was quite time consuming and movements visible on only some CT slices could be missed. We therefore developed a new workflow to both reduce the time requirements and improve the accuracy of the analysis, and integrated it into our in-house developed radiotherapy software environment. The control CT scan is transferred via a network to our treatment planning system VIRTUOS. Inside the treatment planning system, the localizers of the control CT and the planning CT are stereotactically correlated, making the coordinates in both cubes directly comparable. The region of the control CT cube containing the target volume is then automatically matched onto the planning CT. The transformation determined in this step immediately gives the current displacement error. At the moment, we use two rigid matching algorithms: Rigid correlation matching (RCM) based on bony anatomy considers only translational movements and provides the target point correction vector, and mutual information matching (MIM) for bony and soft structures considers translational and rotational movements. Both matching algorithms derive the transformation vector T between planning volume A and transformed control volume BT by searching for maximal similarity in the region of interest V, measured by the correlation coefficient CC and the mutual information MI, respectively. The correlation coefficient between A and BT is calculated by: CC ðA,BT Þ~

X A ! x i {A BT ! x i {BT ! x i [VA,B

ð1Þ

¯ is the mean value of A and BT the mean value whereby A of BT. A three-dimensional Fourier transformation of A and BT into the frequency space using a fast Fourier transformation (FFT) transforms the convolution in Equation (1) into a simple multiplication. After inverse transformation, we obtain the correlation volume that holds the CC values for all possible 3D translation vectors, and the global maximum can easily be found. Since we use discrete Fourier transformation, the precision is restricted to the voxel dimension. The RCM analysis of a typical CT dataset (5126512 points, 30 The British Journal of Radiology, Special Issue 2006

slices) by a standard PC takes less than 30 s. RCM matching was performed for the complete region around the target volume and separately for the upper und lower quarter. Significant differences between the three results indicate that the deviation is more complex than a pure translation. The similarity between volume A and volume BT measured by mutual information is given by: MI ðA,BT Þ~

XX a

b

pABT ða,bÞ: log

pABT ða,bÞ ð2Þ pA ðaÞ:pBT ðbÞ

where a and b are the Hounsfield units of A and B, respectively, and pA(a) the probability density of a and pAB T(a,b) the combined density of A and BT [10, 11]. We do not use all points of A and B to determine each p, but rather a sample of approximately 200 000 values, which is about 2.5% of the complete dataset. This significantly speeds up the calculation without deterioration in the result. In an iterative optimization procedure, BT is translated and rotated until MI reaches a maximum. Therefore an optimization of six variables (tx, ty, tz, and rx, ry, rz) must be performed. In our implementation, we use Powell’s approach [12] to determine the transformation where MI reaches its maximum. This approach provides subvoxel precision. The MIM analysis of a typical CT dataset (5126512 points, 30 slices) is calculated by a standard PC in about 3 min. RCM is significantly faster than MIM, always finds the global optimum and gives a transformation that can be corrected for by a simple shift of the target point. We perform RCM matching of the upper and lower quarter of the region of interest and the MIM matching only to check for deviations that cannot be corrected for by translation only. In any case, the user has to visually check that the resulting match is correct. With standard tools, such as red-green overlay and checkerboard provided by VIRTUOS, this manual check usually takes less than 1 min.

Adaptive radiotherapy In adaptive radiotherapy, deviations of patient repositioning and anatomy from the initial planning CT scan are detected and, if necessary, corrected for. This is in contrast to conventional static radiotherapy where the plan is based on one single CT scan and delivered to the patient throughout all fractions. Different levels of adaptation can be defined: In level 0, the treatment plan is based on several CT scans rather than a single one, and the information about statistical movements of the target and organs at risk are integrated into the treatment plan. This can be accomplished by defining patient-specific or at least sitespecific safety margins [13] or by including statistical methods into the inverse optimization process [14, 15]. Level 1 corrects for errors offline, i.e. the correction is done after several CT scans are gathered. Systematic interfractional errors can be detected and corrected. If the main deviation is a translation, usually it is sufficient to shift the target point and leave the plan unchanged (level 1A). If the error is more complex, the contours have to be S81


adapted to the new geometry, and a new plan has to be generated by re-running the optimization of the inverse planning program (level 1B). Level 2 uses the same methods as level 1, but here the correction is placed in between the CT scan and the directly-following linac irradiation. By analogy with level 1 definitions, level 2A stands for shifting the target point and 2B for generating a new plan. This way not only the systematic error, but also daily random interfractional errors can be corrected. Since the patient remains in the fixation device during the adaptation procedure, the time requirements for this step are more critical than for the lower adaptation levels. Level 3 eventually also takes into account intrafractional errors. This task requires an imaging device operating during the irradiation itself and cannot be accomplished by a linac-CT scanner combination. In the following discussion we will focus on our experiences with Primatom-based radiotherapy for prostate, paraspinal and head and neck cases. However, the general aspects and conclusion apply to other indications as well.

leads to good repositioning of the prostate itself, as is indicated by the results of Beard et al [16] who reported deviations of the prostate above 1 cm in 3% of all cases after bony matching, whereas Wong et al [7] reported deviations above 1 cm in 15% of all cases without bony matching. Manually checking the soft tissues also showed that in our data the displacements of the prostate itself were bigger than the bony variations. In a separate study, the positions of the prostate, bladder and rectum were statistically analysed and taken into account by adding a margin to the gross tumour volume (GTV) and clinical target volume (CTV) to obtain the planning target volume (PTV) (adaptation level 0) [13, 17]. To reduce these margins and further improve the accuracy of prostate treatments, we are currently working on matching algorithms that will enable the automatic correction of interfractional displacements of the prostate itself on adaptation level 2. Deformation of the shape of the prostate and seminal vesicles is reported to be small relative to organ motion [18], so simply shifting the target point (level 2A) will probably be sufficient.

Paraspinal tumours

Results and discussion Dose The additional dose delivered to the patient by frequent CT scans should be as low as possible. Since we do not need the best possible image quality for localization, we use a low mAs product (which introduces some noise to the image) and only scan slices containing the target volume plus approximately 2 cm in cranial and caudal direction. This way, the measured CT dose e.g. per prostate scan was 0.003 Gy at the isocentre and 0.0053 Gy at 1 cm depth. Considering the overall benefit of higher treatment accuracy, we do not find this dose clinically relevant.

Prostate cancer Patients with prostate cancer treated at our institution are immobilized by a wrap-around body cast and a head mask. Usually they receive weekly control CT scans during the treatment course. Figure 2 shows the displacements of the bony structures around the prostate in the control scans relative to their position in the planning CT scan for 10 prostate patients (P1–P10) as they were calculated by rigid correlation matching. The symbols (m, , &) indicate the mean displacement values for each coordinate in x, y and z direction, with the range represented by the standard deviation of all values observed. Checks with RCM for the upper and lower quarter of the target volume separately showed differences below 1 mm, and checks with mutual information matching confirmed that the rotational error was negligibly small with only around 0.5 ˚ (maximum 1 ˚) for all three rotational axes. As can be seen in Figure 2, even without corrections of the target point the repositioning accuracy of the bony anatomy was very good, with displacements below 3 mm in almost all cases. The correct positioning of the bony anatomy alone already

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Paraspinal tumours of different histological origins are challenging because of their close proximity to the spinal cord as a critical organ at risk. For treatment we immobilize these patients by a wrap-around body cast and a head mask and perform regular control CT scans. Since the tumours are very closely related to, or even inside the bony spine, the evaluation of bony landmarks is an exact measure of the tumour position. Figure 3 shows the accuracy and precision of repositioning for seven patients with paraspinal tumours treated at our institution. Also for the paraspinal tumours, rotational errors were negligible as we verified visually, by piecewise RCM and by MIM. The overall repositioning accuracy was good. The accuracy in ventro-dorsal (y: ) direction was equal to or better than 5 mm. In lateral (x: m) and craniocaudal (z: &) direction the typical displacement was around 3 mm. However, in these directions maximal errors above 10 mm were observed. Lateral displacements especially can become critical for these patients, as it is exemplarily shown in Figure 4 for the paraspinal case P7 (a patient with a chordoma of the lumbal spine). The left panel shows the planning CT scan, the right panel shows the stereotactically matched control CT scan of fraction 20. Note the slightly lower image quality of the control CT scan due to the low mAs product we used for scanning. The contours of target, boost and spinal cord are based on the planning CT scan and, by comparing the relation of these contours to the underlying anatomy of the control scan, one immediately sees that the lateral displacement shifts the spinal cord into the high dose area intended for the boost, and that parts of the boost volume are moving out of the high dose area. Paraspinal targets are not significantly affected by intrafractional organ motion, e.g. due to breathing [19], did not show a relevant rotational component and did not change their shape during the treatment course. Therefore we could correct for setup errors by simply shifting the target point. As can be seen in Figure 3, the interfractional error had a pronounced systematic

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Figure 2. Interfractional setup errors in x, y, z direction of 10 prostate patients, based on bony anatomy calculated by the rigid correlation matcher. For each patient, the mean translational error and its standard deviation is plotted.

component. When several (3–4) CT scans revealed such a systematic error, we corrected the target point accordingly and performed further control CT scans (adaption level 1A). Two cases had a highly critical proximity of tumour to the spinal cord (e.g. patient 7 in Figure 4) and showed a stronger random error component, so we performed daily CT scans and corrected the target point directly prior to irradiation (level 2A).

Head and neck tumours The target volume for head and neck tumours regularly includes the base of skull and extends to the upper thoracic aperture. The patients are fixated with a head mask and a vacuum pillow. The cranial part inside the head mask is very accurately repositioned during the

whole treatment course. In contrast, the lower extracranial part shows more variations. The result is a complex deformation of the target volume that cannot be described by a translation and cannot be corrected easily by shifting the target point without changing the treatment plan. Here we show an exemplary case of a patient who was treated for teratocarcinosarcoma of the paranasal sinuses. During the treatment course frequent control CT scans were performed, and while the intracranial part was accurately positioned throughout all fractions, the lower, extracranial part of the body was systematically shifted approximately 1.5 cm along the y-axis from the middle of the treatment course onwards, see Figure 5 comparing the planning CT with the control CT of fraction 20. Note that the contours in Figure 5 refer to the planning CT and do not fit to the anatomical situation at

Figure 3. Interfractional setup errors in x, y, z direction of seven paraspinal patients, based on bony anatomy calculated by the rigid correlation matcher. For each patient, the mean translational error and its standard deviation is plotted. The British Journal of Radiology, Special Issue 2006

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Figure 4. Planning CT scan and control CT scan at fraction 20 of the paraspinal case #7. The contours of the target volume, the boost volume and the spinal cord are also shown.

fraction 20. We therefore re-drew the organ contours for fraction 20 and re-calculated the dose of the original plan on the CT of fraction 20. The results are shown in the form of dose–volume histograms (DVHs) in Figure 6. Figure 6a is the original plan based on the planning CT scan and Figure 6b is the same plan applied to the situation at fraction 20 with updated organ contours. One can see that especially the coverage of the lower target volume has significantly deteriorated. The shift of the extracranial part of the body was systematically seen on three successive control CTs, so we decided to adapt the plan to the new geometry by re-running the optimization in our inverse planning program (KonRad by Siemens OCS) for the control CT scan with the new contours, but without changing the original setting of dose constraints and weighting factors. The new plan was calculated in approximately 2 min, and as shown in Figure 6c the adapted plan resembles the original plan much better than the uncorrected one. Because the contours had to be re-drawn manually, at the moment this procedure takes too much time to fit in between the CT scan and the directly-following irradiation to correct for random interfractional setup errors. However, it allows for a very good adaptation to complex systematic variations occurring during the treatment course (adaptation level 1B).

General aspects The Siemens Primatom is a combination of a linear accelerator and a CT scanner in one room, sharing the same couch. It is built upon standard components which are clinically proven and which work reliably. The patient remains immobilized between the CT scan and the following irradiation. At the moment, we use the Primatom regularly for all patients with extracranial targets. Currently, correction of detected errors in clinical S84

practice is by target point correction either after several CT scans or for each fraction separately (level 1A and 2A) and plan re-optimization for systematic, complex setup errors (level 1B). Real time plan re-optimization (level 2B) is not practical at the moment because of the time constraints. Patients with paraspinal tumours can greatly benefit from the higher accuracy of the treatment. Target point correction seems sufficient; elastic deformation or plan re-optimization was not necessary for the patients we treated. For prostate patients, the bony anatomy is already quite precisely repositioned due to our rigid immobilization device. Here we expect further improvements by elastic matching algorithms that automatically detect the position of the prostate itself (based on Primatom CT scans) and adapt the plan either by target point correction or re-optimization. Also, for patients with head and neck tumours the quality of the radiotherapy could be significantly improved. For these cases, plan re-optimization appeared to be more important than for other tumour sites due to the complex nature of the interfractional deformations. The linac-CT scanner combination already meets all hardware requirements to completely eliminate interfractional setup errors from the treatment course. However, a fast and robust workflow is necessary for widespread use in clinical practice, and to accomplish this further development of algorithms and software tools (e.g. automatic elastic matching) is needed. Realtime plan adaptation to elastic deformations in particular is under current investigation [20] and not in clinical practice yet. Concerning the documentation of an adapted radiation treatment course, the most desirable final record would be a treatment plan where the doses to each volume element delivered throughout the course are superimposed, resulting in concise dose statistics for each structure and a single DVH for the complete The British Journal of Radiology, Special Issue 2006


Figure 5. Exemplary head and neck case – comparison of planning CT and control CT at fraction 20. In the first row the good repositioning in the upper part of the target can be seen, and the second row shows the error of approximately 1.5 cm along the y-axis in the lower part of the target volume.

treatment. However, this requires the tracking of each voxel throughout all CT scans. Simply averaging the DVHs of the single fractions to obtain a final DVH would lead to erroneous results since, in a DVH, the spatial information is lost. Techniques for tracking the voxels are under current development; at the moment, each modified (adapted) treatment plan is documented separately (as in Figure 6a,c). Our patient immobilization device with customized wrap-around body casts alone leads to quite high repositioning accuracy, but is also quite labour and time intensive to build. When using adaptive radiotherapy

strategies, the fixation can probably be made less sophisticated, lowering the overall workload and further strengthening the role of a combined imaging/treatment device such as the linac-CT scanner combination.

Conclusions The linac-CT scanner combination is a device for adaptive radiotherapy that delivers all information necessary to eliminate interfractional setup errors from a fractionated treatment course. Currently available matching algorithms

Figure 6. Same case as in Figure 5. Dose–volume histograms of (a) original plan, applied to planning CT, (b) original plan, applied to fraction 20 with updated contours and (c) re-optimized plan. For better comparison, the 80% dose/90% volume point is highlighted. The structures are 1-upper target volume, 2-lower target volume, 3-left parotid gland, 4-spinal cord, 5-right parotid gland. The British Journal of Radiology, Special Issue 2006

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and software components make it well suited for real-time and off-line target point corrections, and for off-line reoptimization. Further investigation needs to be carried out until it can be used clinically for real-time re-optimization. In principle the device is not suited for adaptation to intrafractional variations.

References 1. Lohr F, et al. Noninvasive patient fixation for extracranial stereotactic radiotherapy. Int J Radiat Oncol Biol Phys 1999;45:521–7. 2. Herfarth KK, et al. Extracranial stereotactic radiation therapy: set-up accuracy of patients treated for liver metastases. Int J Radiat Oncol Biol Phys 2000;46:329–35. 3. Groh BA, Siewerdsen JH, Drake DG, Wong JW, Jaffray DA. A performance comparison of flat-panel imager-based MV and kV cone-beam CT. Med Phys 2002;29:967–75. 4. Jaffray DA, Siewerdsen JH, Wong JW, Martinez AA. Flatpanel cone-beam computed tomography for image-guided radiation therapy. Int J Radiat Oncol Biol Phys 2002;53:1337–49. 5. Uematsu M, et al. A dual computed tomography linear accelerator unit for stereotactic radiation therapy: a new approach without cranially fixated stereotactic frames. Int J Radiat Oncol Biol Phys 1996;35:587–92. 6. Kuriyama K, et al. A new irradiation unit constructed of self-moving gantry-CT and linac. Int J Radiat Oncol Biol Phys 2003;55:428–35. 7. Wong JR, et al. Image-guided radiotherapy for prostate cancer by CT-linear accelerator combination: prostate movements and dosimetric considerations. Int J Radiat Oncol Biol Phys 2005;61:561–9. 8. Cheng CW, et al. Commissioning and clinical implementation of a sliding gantry CT scanner installed in an existing treatment room and early clinical experience for precise tumor localization. Am J Clin Oncol 2003;26:e28–e36. 9. Karger CP, Jakel O, Debus J, Kuhn S, Hartmann GH. Threedimensional accuracy and interfractional reproducibility of

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Dose-guided radiation therapy with megavoltage cone-beam CT J CHEN,

PhD,

O MORIN,

BSc,

M AUBIN,

Eng-MSc,

M K BUCCI,

MD,

C F CHUANG,

PhD

and J POULIOT,

PhD

UCSF Comprehensive Cancer Center, Department of Radiation Oncology, University of California San Francisco, 1600 Divisadero Street, Suite H1031, San Francisco, CA 94143, USA ABSTRACT. Recent advances in fractionated external beam radiation therapy have increased our ability to deliver radiation doses that conform more tightly to the tumour volume. The steeper dose gradients delivered in these treatments make it increasingly important to set precisely the positions of the patient and the internal organs. For this reason, considerable research now focuses on methods using three-dimensional images of the patient on the treatment table to adapt either the patient position or the treatment plan, to account for variable organ locations. In this article, we briefly review the different adaptive methods being explored and discuss a proposed dose-guided radiation therapy strategy that adapts the treatment for future fractions to compensate for dosimetric errors from past fractions. The main component of this strategy is a procedure to reconstruct the dose delivered to the patient based on treatment-time portal images and pre-treatment megavoltage cone-beam computed tomography (MV CBCT) images of the patient. We describe the work to date performed to develop our dose reconstruction procedure, including the implementation of a MV CBCT system for clinical use, experiments performed to calibrate MV CBCT for electron density and to use the calibrated MV CBCT for dose calculations, and the dosimetric calibration of the portal imager. We also present an example of a reconstructed patient dose using a preliminary reconstruction program and discuss the technical challenges that remain to full implementation of dose reconstruction and dose-guided therapy.

The rationale for adaptive radiation therapy and dose-guided radiation therapy Recent advances in fractionated external beam radiation therapy, such as three-dimensional conformal and intensity-modulated radiation therapy (IMRT), have increased our ability to deliver radiation doses that conform more tightly to the tumour volume. Clinical studies and simulations indicate that these more conformal, higher dose treatments can decrease both the spread of disease and normal tissue complications [1–5]. Increasing use of functional imaging will also motivate further complexity in radiation treatment plans to include concurrent boosts in regions of high cancerous growth [6, 7]. As these dose distributions conform more tightly to the patient anatomy, dose gradients necessarily become steeper inside the irradiated volume. Using IMRT, a dose gradient of 10% mm21 can be achieved easily. Thus, it is increasingly important to set precisely the positions of the patient and the internal organs. Currently, external markers and patient immobilizing masks and casts are used to reproduce the skeletal position of the patient with about 3 mm accuracy over several weeks of treatment [8]. However, the effectiveness of these alignment and immobilization techniques are limited by changes in the internal organ locations relative to bony and external markers. For example, the prostate can shift up to 1 cm relative to the pelvic bones due to variations in rectal/bladder filling. During the course of head and neck cancer treatment, the tumour This research was supported by Siemens Oncology Care Systems.



can shrink and the patient can lose significant weight, resulting in dosimetric errors as large as 40% [9, 10]. For this reason, imaging tools in the treatment room and methods of adapting treatments to match the patient anatomy on the treatment table are the keys to realising the full benefit of conformal therapy. For many decades, imaging inside the treatment room has played a role in verifying radiation therapy treatment. Portal images, projection images of the patient using the treatment aperture, are used to confirm the patient position and verify coverage of the tumour. The use of radiographic film for portal imaging has limited the frequency of this verification due to the required time and dose to the patient. However, recent implementation of electronic portal imaging devices (EPIDs) allows a digital image to be acquired in a few seconds with low doses. This has allowed the use of daily portal imaging to visualize and adjust the patient position before each treatment. For example, using implanted gold markers to locate the prostate, daily portal imaging has been used to position the prostate with 1–2 mm accuracy [11–13]. The use of portal imaging to adjust patient position before treatment is limited, however, because soft tissue cannot be visualized without implanted markers and the full three-dimensional (3D) geometry is obscured by the projection onto a two-dimensional (2D) plane. Therefore, considerable research now focuses on developing threedimensional imaging of the patient on the treatment table. Several systems have been developed including (1) a ‘‘CT on rails’’ system, requiring an additional diagnostic CT machine in the treatment room [14]; (2) a kilovoltage cone-beam CT (kV CBCT) system, consisting S87

J Chen, O Morin, M Aubin et al

of an additional kV X-ray source and detector attached to the treatment gantry [15,16] (these systems are described more fully in this issue in papers by Thieke et al and Moore et al, respectively); (3) a megavoltage cone-beam CT (MV CBCT) system using the pre-existing treatment machine and EPID for imaging [17–19]; (4) a MV CT system, using the pre-existing treatment machine with an attached arc of detectors [20]; and (5) a tomotherapy system, replacing the traditional treatment machine (beam) with a CT ring and a MV beam source [21–23]. These imaging systems continue to improve and recent results indicate that 1–2% soft-tissue contrast resolution is possible [15, 17, 18, 21] as well as accurate localization of various tumours [14, 16, 19, 20, 22, 23]. In the above examples of image-guided radiation therapy (IGRT), treatment room imaging modalities are used to translate and rotate the patient to better match the patient position used for treatment planning. Another potentially more powerful use of these images is to modify the delivered treatment fields to account for the variable patient position. This type of adaptive radiation therapy could adjust for the changing relative positions of the internal organs and the changing shape of the organs. This is particularly important for organs that move significantly during the course of treatment. For these sites, techniques under current development include gated treatments (halting irradiation when the target is out of a certain acceptable region) [24–27] or target tracking during irradiation using specially designed mobile linear accelerators [28, 29]. For some sites, however, the most important anatomical changes occur between treatment fractions. In this case, a pretreatment image may be used to adjust the treatment fields immediately before irradiation [30, 31]. Another possibility is to determine patient-specific anatomical variation using images from the first week of treatment and to tailor the treatment plan for future fractions to account for the individual’s variation [32–34]. Finally, if the dose that was delivered in previous fractions can be estimated, the treatment plan for future fractions may be re-optimized to compensate for dosimetric errors [35]. This dose-guided therapy could correct for both errors due to patient anatomical changes as well as machine delivery errors, thus providing the most accurate dose delivery. The various adaptive radiation therapy schemes are depicted in Figure 1.

The development of dosimetric verification and reconstruction Currently, few methods are used to track the dose delivered during treatment. Standard techniques involve measuring doses on the patient surface using diodes or thermoluminescent dosemeters. However, these techniques provide only point dose measurements, and the time and effort to place the dosemeters on the patient and process the data limit their clinical use. Consequently, few institutions use these methods regularly for treatment verification. A new implantable MOSFET dosemeter has also been developed [36]. This dosemeter directly measures the dose in critical internal structures, but again provides only a point measurement and is an invasive technique with limited application. What is needed to verify conformal therapies is an automated method to reconstruct the full 3D dose distribution. Several researchers have suggested methods to reconstruct the delivered patient dose during treatment. Most methods propose using on-board EPIDs to quickly and easily acquire a two-dimensional array of digitized X-ray measurements in a precisely positioned plane in the treatment exit beam. A few formulae have been derived to estimate the dose to the exit surface, midplane, or centre point of the patient based solely on EPID measurements [37–40]. To find a 3D patient dose distribution, however, requires additional information about the patient position and attenuation of the beam. For breast treatments, a simple patient contour may give sufficient information [41]. However, in general, information on tissue inhomogeneity is also necessary. Several years ago, it was suggested that the planning CT could be used for this purpose [42, 43], but this method would fail to detect dosimetric errors produced by the variable patient and organ positions and shapes. The 3D imaging modalities that are being developed for IGRT provide an obvious opportunity to simultaneously obtain the patient geometry for reconstructing dose. Currently, there is active development of dose reconstruction procedures for tomotherapy systems, and 3% accuracy in low-gradient regions has been demonstrated [44]. A pilot study using MV CBCT on a traditional treatment machine also found good relative agreement with measurements, but a systematic absolute deviation [45].

Figure 1. A general view of adaptive radiation therapy. The large grey arrow represents the conventional flow of treatment, and the small arrows indicate the possible points of feedback into the process. S88


Dose-guided radiation therapy with MV CBCT

Dose-guided radiation therapy using MV CBCT and treatment-time portal images In 2003 [46], we began developing a procedure to reconstruct the dose delivered to the patient based on treatment-time portal images and pre-treatment MV CBCT. Our procedure follows the steps described below and depicted in Figure 2. Step 1A: Prior to treatment, with the patient in the treatment setup position, acquire a MV CBCT image. This image can be used to align the patient as closely as possible to the planned position and also provides the photon attenuation information necessary to reconstruct the delivered dose. Step 1B: Convert the MV CBCT image to effective photon attenuation coefficient. Generally, this can be accomplished by calibrating the MV CBCT system using a calibration phantom composed of materials with known electron densities. However, imaging artefacts in the MV CBCT image may need to be corrected to improve the calibration accuracy.

Step 2A: During the treatment, acquire portal images of the treatment beam as it exits the patient. This portal image is acquired using the same EPID used for the CBCT imaging. Step 2B: Convert the portal images to a 2D map of treatment beam energy fluence. The acquired portal image signal is a convolution of the energy fluence incident on the detector with the detector response to radiation. Moreover, the energy fluence consists of both the primary beam and radiation scattered from the patient. To use the portal image for dose calculations, the primary energy fluence must be derived from the portal image. Step 3: Back-project the energy fluence measured at the detector plane through the CBCT of the patient, accounting for the 1/r2 falloff of radiation from a point source and attenuation through the patient. This calculation is easily accomplished if the position of the detector plane relative to the patient and source is accurately known. Step 4: Calculate the 3D dose distribution delivered to the patient using a dose calculation engine. This type of

Figure 2. Overview of proposed dose reconstruction procedure using MV CBCT imaging and treatmenttime portal imaging.


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dose calculation is the same as that performed for treatment planning purposes, and all the techniques that have been developed for treatment planning may be used. The reconstruction procedure described above provides an estimate of the 3D dose distribution deposited in the patient as represented by the MV CBCT. Several uses of the reconstructed dose distribution to guide future treatments can be envisaged. Scenario 1: The most basic use of the reconstructed dose is to provide a dosimetric verification that the treatment delivery generally provides the desired dose distribution and that no gross errors exist. This verification could be performed during the first treatment and repeated weekly throughout treatment. This simple approach would effectively reduce gross dosimetric errors, but would not otherwise increase the precision of the delivered dose. Scenario 2: If the patient dose is reconstructed for the first week of treatment, the variation in the delivered dose may also be evaluated. If the MV CBCT for each treatment is contoured to delineate the various important structures, the variation in dosimetric indices, such as the maximum dose to sensitive normal structures or the dose to 95% of the tumour volume, can be calculated. General systematic trends such as the under or over dosing of particular extremities of a structure may also be detected by examining the dose distributions over the first week. Based on this information, the treatment plan can be modified, for example, to increase or decrease margins of the tumour in particular directions. In this manner, the treatment plan can be tailored to each individual patient. Scenario 3: Finally, a complete dose-guided therapy system would be able to integrate the dose over previous fractions. This would require the ability to deform the daily MV CBCT images to map identical points in the patient before the integral dose is calculated [47]. The cumulative dose distribution can be used to adjust the treatment plan to compensate for deviations from the desired distribution, thus improving the accuracy and conformality of the overall treatment. The dose reconstruction procedure and the doseguided therapy described above continue to be developed and researched. This article summarizes the work to date and comments on the remaining challenges. First, we present a description of a MV CBCT system that has been implemented on a linear accelerator for clinical use. We then describe experiments performed to calibrate the MV CBCT for electron density and to use the calibrated MV CBCT for dose calculations. We also briefly describe the dosimetric calibration of an EPID for dose reconstruction. Finally, we present an example of a reconstructed patient dose using a preliminary reconstruction program and discuss the technical challenges that remain to full implementation of dose reconstruction and dose-guided therapy.

In the radiation oncology context, the imaging beam is produced by the conventional linear accelerator used for treatment, and the projection images are detected using on-board EPIDs. The imaging photons, therefore, are primarily in the mega-electron volt energy range. In this configuration, the patient can be positioned once on the treatment table and need not be repositioned between imaging and treatment. As the linear accelerator gantry and the EPID rotate about the patient, the EPID and beam source positions will shift from their ideal isocentric locations due to sagging of the mechanical supports. To correct for this effect, we perform a geometric calibration of the system, illustrated in Figure 3 [48, 49]. This calibration provides a unique relationship between the position of a voxel in the reconstruction volume and a pixel on the detector plane for each angle. Because the EPID used for imaging is also used to detect the exit beam fluence, the same calibration information can be employed during the dose reconstruction procedure to back-project the energy

MV cone-beam CT imaging MV cone-beam CT imaging is a 3D reconstruction procedure similar to conventional CT. A series of projection measurements, in this case 2D portal images, are acquired at many angles around the patient. The image reconstructed is a 3D image without slice artefacts. S90

Figure 3. Depiction of the geometric calibration of the linear accelerator/electronic portal imaging device (EPID) system for cone beam CT (CBCT) imaging and for dose reconstruction. The result of the calibration is a set of projection matrices (P) that map a point in space (RXYZ) to the projected point on the detector plane (Ruv).



fluence through the MV CBCT volume. This prevents any possibility of misregistration between the EPID measurements and the MV CBCT volume. The MV CBCT system installed in our clinic has been previously described [19]. Briefly, it consists of an amorphous-silicon flat panel EPID integrated with a clinical linear accelerator. The total exposure of the CBCT acquisition can be varied from 1 to 60 monitor units. Upon patient selection, a reference CT is automatically loaded into the software. The linear accelerator gantry then rotates in a continuous 200˚ arc acquiring images at 1 ˚ increments. This acquisition procedure lasts about 45 s. The image reconstruction starts immediately after the acquisition of the first portal image, and a 25662566256 reconstruction volume is completed in 110 s. The software automatically registers the MV CBCT with the reference CT and calculates table shifts for patient alignment. To date, 38 patient MV CBCT images have been acquired in our clinic. All patients have given informed consent, and the patient image acquisitions are performed in accordance with the institutional review board’s ethical standards. Depending on the frequency of the acquisitions, the dose used for MV CBCT ranges from approximately 1.5 cGy to 12 cGy delivered at the point of rotation (the isocentre). The dose at the entrance surface of the arc reaches about 160% of the isocentre dose for an imaged pelvis and 133% for the head and neck region. The dose at the exit surface falls to about 66% of the isocentre dose for a pelvis and 55% for the head and neck region. Figure 4 presents four MV CBCT images acquired weekly on the same patient to study tumour evolution. At each new acquisition, the dose was lowered. The last CBCT of the series was acquired with approximately 2.9 cGy delivered at the isocentre, still presenting enough soft-tissue information to assess the tumour size and perform patient alignment. Three-dimensional imaging of the patient in the treatment position exposes the difficulties created by distortion of patient anatomy. Figure 5 displays the fusion of a MV CBCT image (grey) with the planning CT (colour). In this case, a physician has manually registered the two sets of images by aligning the base of the skull. A considerable shift, up to 6 mm, can be observed in the positions of the spinal cord between the two image sets. This misplacement of the spinal cord could not be corrected by translating or rotating the MV CBCT image relative to the CT as it was caused by an increase in the arching of the patient’s neck. Although several fractions would be needed to assess if this misplacement occurs regularly, the new anatomy, as depicted by the MV CBCT image, could be used to study the dosimetric impact of the patient’s anatomical distortion.

MV CBCT calibration for dose calculation To use the MV CBCT image in a dose reconstruction program, the signal from each voxel must be converted to effective photon attenuation coefficient for the beam spectrum (Step 1B of our dose reconstruction procedure). To perform this conversion, the MV CBCT system can be calibrated using a CT calibration phantom (CIRS Model 062, Norfolk, VA) with tissue-equivalent inserts, as is The British Journal of Radiology, Special Issue 2006

currently done with kV CT. A table is formed mapping CT signal intensity to electron or physical density which can then be converted to photon attenuation coefficient for a known beam spectrum. Figure 6 shows the results of performing this simple calibration on our MV CBCT system using the following inserts of relative electron density with respect to water: lung inhale (0.190), lung exhale (0.489), adipose (0.952), breast (0.976), water (1), muscle (1.043), liver (1.052), trabecular bone (1.117) and dense bone (1.512). The relationship between MV CBCT signal and electron density is linear. These results are similar to previous work with MV fan-beam CT performed on a tomotherapy unit at 6 MV [50]. Although the above calibration works well for the narrow CT calibration phantom, the MV CBCT images of extended objects exhibit cupping artefacts due to the influence of scattered radiation reaching the EPID. Figure 7 illustrates this cupping effect on the MV CBCT of a large cylinder of water. If uncorrected, this cupping artefact will also appear in the image converted to photon attenuation coefficient, leading to errors in the calculated dose. However, a simulation study using the large cylinder of water pictured in Figure 7 indicates that the dosimetric errors in a homogeneous medium produced by such severe cupping artefacts remain relatively small, approximately 4% for a single open field [51]. This suggests that a crude correction of the cupping artefact in MV CBCT images may be sufficient to obtain acceptable dosimetric accuracy. To test this hypothesis, the MV CBCT of a water cylinder was used to model the spatial dependence of the cupping artefact. A spatially dependent correction function was derived from this cupping model. This correction function was then applied to the MV CBCT of an anthropomorphic head phantom as a rough correction for the cupping artefact in the image. After conversion to density using the MV CBCT calibration curve, this image was imported into a commercial treatment planning system (Philips Pinnacle, Bothell, WA). The dose calculated using the MV CBCT compared well with the dose calculated using a kV CT of the same phantom. Using a gamma index comparison with a 3% dose and 3 mm distance-toagreement criterion, 98% of calculated dose points fell within the acceptance criteria. The above example demonstrates the potential of using MV CBCT images for dose calculations. Besides using these images for dose reconstruction, using patient MV CBCT images in the treatment planning system, as performed on the head phantom described above, would also provide a useful verification. The MV CBCT provides a more accurate representation of the patient on the treatment table. Applying the treatment plan to the MV CBCT would provide a first estimate of the dose delivered to the patient during treatment. The effects of modified patient position or anatomy could be evaluated. However, the beam delivery itself could not be verified without a full dose reconstruction based on measurements of the treatment beam.

Calibration of EPIDs for exit-plane dose Besides the patient photon attenuation data, the other necessary piece of information for dose reconstruction is S91


Figure 4. Examples of megavoltage cone beam CT (MV CBCT) images at different exposure levels, from 2.9 cGy to 10 cGy.

the treatment beam energy fluence derived from the treatment-time portal images (Step 2 of our dose reconstruction procedure). An intermediate step to determining the energy fluence is to convert the EPID image to a measurable form of dose, in our case the dose in water measured in the detector plane and at a depth of 1.5 cm [52]. The advantage of first calibrating the EPID against dose in water is that it can be accomplished by experiments since the dose in a water phantom is easily measured. The calibration can then be validated by measurements as well. Moreover, the dose in water can S92

be more easily converted to energy fluence due to the great number of water dose deposition models and algorithms that have already been developed. To translate the EPID signal to dose in water, we employ convolution models of dose deposition. The lateral spread of the dose in the EPID and in the water is described by empirically derived kernels. Because the EPID consists of millions of individual pixels, the dose deposited in each pixel is also multiplied by a spatially dependent sensitivity factor that accounts for inhomogeneity in the detector response. Finally, comparisons of The British Journal of Radiology, Special Issue 2006


Bartlett, TN). Figure 8 shows a comparison between the measured dose at a depth of 1.5 cm of water and the calibrated EPID signal for a 10 cm square open field. The EPID signal matches the measured dose to within 2% (2 standard deviations) for the in-field regions (excluding the penumbra).

A dose reconstruction program

Figure 5. Registration of a patient megavoltage cone beam CT (MV CBCT) (grey) with the kV CT (colour) used for treatment planning. A large difference in the arching of the neck causes a considerable deviation in the spinal cord position.

EPID and ion chamber measurements are used to form conversion tables that translate between the EPID signal and dose in water. To test the calibration procedure, EPID images of the exit beam were acquired through a Rando anthropomorphic head phantom (The Phantom Laboratory, Salem, NY). The calibrated EPID images were compared with the dose measured using an ion chamber (Scanditronix-Wellho¨fer CC13, Bartlett, TN) scanned in a water tank (Scanditronix-Wellho¨fer blue phantom,

Figure 6. Megavoltage cone beam CT (MV CBCT) intensity as a function of electron density for tissue-equivalent inserts in a CT calibration phantom (pictured in above left). The British Journal of Radiology, Special Issue 2006

Utilizing some of the work described above, we performed a preliminary version of the dose reconstruction procedure on the treatment of a head and neck patient in our clinic. A MV CBCT image was acquired of the patient set up on the table as for treatment (Step 1A). The same day, portal images were acquired (Step 2A) during the patient’s normal course of treatment (6 MV beam, 2 opposed lateral wedged fields and an anterior– inferior oblique open field). To utilize the MV CBCT image in the dose reconstruction program, it must first be converted to effective photon attenuation coefficient (Step 1B). For this test case, the MV CBCT was converted to attenuation coefficient using a spatially dependent calibration that utilizes the kV CT patient image as a reference. This allowed us to reduce the effects of the MV CBCT calibration on the reconstructed dose, thus highlighting the dosimetric impact of the remaining steps of the procedure. To convert the portal images to energy fluence (Step 2B), the portal images were first converted to equivalent dose in water using the calibration procedure described above. To infer the energy fluence at the detector plane from the equivalent dose in water, we used an in-house dose calculation program that predicts the dose at a depth of 1.5 cm of water given the energy fluence at the water surface. This energy fluence is then iteratively corrected until the predicted dose matches the measured dose. To calculate the dose in water, we used convolution kernels published in the literature [53], derived using Monte Carlo calculations and assuming a 6 MV spectrum. The energy fluence that is derived using this method is composed of both primary beam as well as radiation scattered from the patient. For this study, the contribution of the scattered radiation was neglected. The two remaining steps to the dose reconstruction process are (Step 3) the back-projection of the energy fluence measured at the detector plane through the CBCT of the patient and (Step 4) the calculation of the 3D dose distribution delivered to the patient using a dose calculation engine. To perform the back-projection, we utilized the geometric information obtained during calibration of the MV CBCT imaging system (depicted in Figure 3). The geometric calibration of the system yields a set of projection matrices that map a point in space to a pixel in the detector plane. The projection matrix for each angle accurately accounts for all geometric factors such as sag in the detector or gantry, detector rotation, or variation in the detector to source distance. These projection matrices were used to backproject the energy fluence from the detector plane through the CBCT volume while correcting for 1/r2 fall-off and the attenuation of each intersected voxel. The final step of the reconstruction procedure is to calculate the dose deposited in the patient from the S93


Figure 7. Radial (top row) and axial (bottom row) profiles through the megavoltage cone beam CT (MV CBCT) images of a large cylinder filled with water. The unmodified CBCT (left) exhibits a large cupping artefact as a result of scattered radiation reaching the electronic portal imaging device (EPID). Using a simple 3D cupping model effectively reduces the artefact (right). The radial and axial slices of the MV CBCT images (insets) are displayed using the same windowing level.

energy fluence and the attenuation coefficient for each voxel. The total energy released in each voxel that interacts with the beam is proportional to the energy

fluence multiplied by the attenuation coefficient. The spatial distribution of the deposited energy can then be described using a kernel. The kernels we used for this

Figure 8. Comparisons of measured dose profiles (line) in water and calibrated electronic portal imaging device (EPID) profiles (circle with dot) for a 10 cm square field through a Rando head phantom.

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purpose were the same kernels used to determine the energy fluence at the detector plane from the equivalent dose in water. The application of the kernels to calculate the dose was performed using in-house software utilizing the collapsed-cone superposition method [53]. In this method, the energy deposition calculation is only performed along a set of rays emanating from each interaction voxel. Figure 9 shows the comparison between the planned dose distribution found using the patient kV CT image and a commercial treatment planning system (Philips Pinnacle, Bothell, WA) and the reconstructed dose distributions found using the MV CBCT, the treatmenttime portal images, and the in-house dose reconstruction program. There are some qualitative similarities, but also some marked differences. The reconstructed dose distribution appears to be approximately 10% higher than the dose predicted by the planning system. It is likely that this is in part due to an increase in the portal image signal from the scattered radiation that was not corrected in this preliminary version of the dose reconstruction.

There also appears to be a slight difference in the alignment of the beams detected by the portal images. The doses from the treatment planning system suggest a slight gap between the opposed lateral fields and the anterior field. In contrast, the reconstructed dose distribution has a high dose band at the intersection of the fields. Without further verification, it is not clear whether this slight difference in field alignment was a real event detected using the treatment-time portal images. Other possible causes for the differences in the two dose distributions include differences in the dose calculation engines, differences in patient position or anatomy in the two images, as well as persistent cupping artefacts in the MV CBCT. As the above example demonstrates, much research remains to be done to increase the dosimetric accuracy of our dose reconstruction program. Currently, we continue to work toward simple but effective techniques to reduce cupping artefacts in the MV CBCT images and to calibrate the MV CBCT for photon attenuation coefficient. We also continue to refine our EPID dosimetric

Figure 9. Comparisons between planned isodose contours calculated using the patient kV CT image and a commercial treatment planning system (left) and reconstructed isodose contours calculated using the megavoltage cone beam CT (MV CBCT), the treatment-time portal images, and an in-house dose reconstruction program (right). The British Journal of Radiology, Special Issue 2006

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calibration models described above and to improve the conversion of the EPID signal to primary energy fluence. One of the remaining challenges is to implement a correction for the scatter contribution in the portal images. Portal image scatter correction has been investigated by other researchers, and some good results have been reported using a scatter-to-primary ratio model and Monte Carlo-based scatter kernels [54–56]. Finally, once the individual steps of the dose reconstruction procedure have been optimized, the dosimetric accuracy of the full procedure will need to be determined using dose measurements in phantoms. As discussed below, the dosimetric accuracy achieved will affect the clinical application of the dose reconstruction procedure.

Future directions in dose-guided therapy research This article has summarized the work performed as well as the challenges remaining to develop a dose reconstruction procedure based on MV CBCT images of the patient on the treatment table and treatment-time portal images. As described earlier, the ability to reconstruct the delivered patient dose opens up the possibility of adapting the patient treatment plan to improve dose delivery. The accuracy of the dose reconstruction procedure and the availability of image processing tools will affect how treatment may be guided using this new dose information. Our initial goal is to achieve 5% accuracy for the reconstructed patient dose. With this level of accuracy, gross dosimetric errors, which have been demonstrated to be as high as 40% in cases of considerable patient weight loss [10], could be detected and corrected. Implementation of more complex dose-guidance strategies, such as scenarios 2 and 3 discussed earlier, will require increased dosimetric accuracy as well as the ability to precisely locate the dose distribution in terms of critical structures. It is here that the rapidly advancing field of 3D image processing will play a key role. Tools such as automated segmentation and 3D deformable registration increase our ability to determine under or over dosed regions as well as track the cumulative dose to various organs in the patient. By focusing on the key parameter determining radiation treatment outcomes, dose verification and doseguided therapy have the potential to considerably improve the treatment of cancer. Moreover, they offer the opportunity to increase our understanding of treatment effectiveness, improving our knowledge of the radiation doses and distributions that lead to the control of cancer or the injury of normal structures. Although this level of precision has long been a goal in radiation oncology, the continuing advances in imaging technology and in imaging processing may soon make this goal attainable.

Acknowledgments The authors would like to acknowledge the following persons for their valuable contributions, enlightening discussions and active participation on the acquisition of the clinical cone-beam images. At UCSF, Albert Chan, S96

Chris Malfatti, Amy Gillis, Ping Xia, Lynn Verhey. And at Siemens OCS, Ali Bani-Hashemi. This research was supported by Siemens Oncology Care Systems (OCS). One of the authors (OM) wishes to acknowledge a doctoral scholarship from NSERC-Canada.

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33. Rehbinder H, Forsgren C, Lof J. Adaptive radiation therapy for compensation of errors in patient setup and treatment delivery. Med Phys 2004;31:3363–71. 34. van Herk M. Errors and margins in radiotherapy. Semin Radiat Oncol 2004;14:52–64. 35. Wu C, Jeraj R, Olivera GH, Mackie TR. Re-optimization in adaptive radiotherapy. Phys Med Biol 2002;47:3181–95. 36. Scarantino CW, Ruslander DM, Rini CJ, Mann GG, Nagle HT, Black RD. An implantable radiation dosimeter for use in external beam radiation therapy. Med Phys 2004;31:2658–71. 37. Kirby MC, Williams PC. The use of an electronic portal imaging device for exit dosimetry and quality control measurements. Int J Radiat Oncol Biol Phys 1995;31:593–603. 38. Boellaard R, van Herk M, Uiterwaal H, Mijnheer B. Twodimensional exit dosimetry using a liquid-filled electronic portal imaging device and a convolution model. Radiother Oncol 1997;44:149–57. 39. Boellaard R, Essers M, van Herk M, Mijnheer BJ. New method to obtain the midplane dose using portal in vivo dosimetry. Int J Radiat Oncol Biol Phys 1998;41:465–74. 40. Chang J, Mageras GS, Chui CS, Ling CC, Lutz W. Relative profile and dose verification of intensity-modulated radiation therapy. Int J Radiat Oncol Biol Phys 2000;47:231–40. 41. Louwe RJW, Damen EMF, van Herk M, Minken AWH, Torzsok O, Mijnheer BJ. Three-dimensional dose reconstruction of breast cancer treatment using portal imaging. Med Phys 2003;30:2376–89. 42. McNutt TR, Mackie TR, Reckwerdt P, Paliwal BR. Modeling dose distributions from portal dose images using the convolution/superposition method. Med Phys 1996;23:1381–92. 43. Hansen VN, Evans PM, Swindell W. The application of transit dosimetry to precision radiotherapy. Med Phys 1996;23:713–21. 44. Kapatoes JM, Olivera GH, Balog JP, Keller H, Reckwerdt PJ, Mackie TR. On the accuracy and effectiveness of dose reconstruction for tomotherapy. Phys Med Biol 2001;46:943–66. 45. Partridge M, Ebert M, Hesse B-M. IMRT verification by three-dimensional dose reconstruction from portal beam measurements. Med Phys 2002;29:1847–58. 46. Pouliot J, Xia P, Aubin M, Verhey L, Bani-Hashemi A, Ghelmansarai F, et al. Low-dose megavoltage cone-beam CT for dose-guided radiation therapy (abstract). ASTRO 45th Annual Meeting; 2003 October 19-23; Salt Lake City, UT. Int J Radiat Oncol Biol Phys 2003;57:S183–4. 47. Schaly B, Kempe JA, Bauman GS, Battista JJ, van Dyk J. Tracking the dose distribution in radiation therapy by accounting for variable anatomy. Phys Med Biol 2004;49:791–805. 48. Wiesent K, Barth K, Navab N, Durlak P, Brunner T, Schutz O, et al. Enhanced 3-D reconstruction algorithm for C-arm systems suitable for interventional procedures. IEEE Trans Med Imaging 2000;19:391–403. 49. Morin O, Chen J, Aubin M, Pouliot J. Evaluation of the mechanical stability of a megavoltage imaging system using a new flat panel positioner. In: Flynn MJ, editor. Medical Imaging 2005: Physics of Medical Imaging. Proceedings of SPIE Medical Imaging: Physics of Medical Imaging; 2005 February 13-15; San Diego, CA. Bellingham, WA: SPIE Press, 2005:704–10. 50. Ruchala KJ, Olivera GH, Schloesser EA, Hinderer R, Mackie TR. Calibration of a tomotherapeutic MVCT system. Phys Med Biol 2000;45:N27–36. 51. Chen J, Morin O, Chen H, Aubin M, Pouliot J. The effect of MV cone-beam CT cupping artifacts on dose calculation accuracy (abstract). AAPM 47th Annual Meeting; 2005 July 24-28; Seattle, WA. Med Phys 2005;32:1936.

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Image registration and data fusion in radiation therapy M L KESSLER,

PhD

The University of Michigan, Ann Arbor, MI 48103, USA ABSTRACT. This paper provides an overview of image registration and data fusion techniques used in radiation therapy, and examples of their use. They are used at all stages of the patient management process; for initial diagnosis and staging, during treatment planning and delivery, and after therapy to help monitor the patients’ response to treatment. Most treatment planning systems now support some form of interactive or automated image registration and provide tools for mapping information, such as tissue outlines and computed dose from one imaging study to another. To complement this, modern treatment delivery systems offer means for acquiring and registering 2D and 3D image data at the treatment unit to aid patient setup. Techniques for adapting and customizing treatments during the course of therapy using 3D and 4D anatomic and functional imaging data are currently being introduced into the clinic. These techniques require sophisticated image registration and data fusion technology to accumulate properly the delivered dose and to analyse possible physiological and anatomical changes during treatment. Finally, the correlation of radiological changes after therapy with delivered dose also requires the use of image registration and fusion techniques.

Data from multiple anatomical and functional imaging studies have become important components of patient management in radiation therapy. From initial diagnosis to treatment planning and from delivery to monitoring the patient post-therapy, these data drive the decisions about how the patient is treated and help assess the progress and efficacy of therapy. While X-ray CT remains the primary imaging modality for most aspects of treatment planning and delivery, the use of data from other modalities such as MRI and MR spectroscopy (MRS) and positron/single photon emission tomography (PET/SPECT) is becoming increasingly prevalent and valuable, especially when taking advantage of highly conformal treatment techniques such as intensity-modulated radiotherapy [1–3]. These additional imaging studies provide complementary information to help elucidate the condition of the patient before, during and after treatment. The use of time-series image data to assess physiological motion for initial planning as well as anatomical and functional changes for possible treatment adaptation is becoming more widespread as diagnostic imaging devices produce quality 4D image data and as X-ray imaging systems are incorporated into the treatment room. In order to make use of the information from these multiple imaging studies in an integrated fashion, the data must be geometrically registered to a common coordinate system. This process is called image registration. Once different datasets are registered, information such as tissue boundaries, computed dose distributions and other image or image-derived information can be mapped between them and combined. This process is called data fusion. Figure 1a provides a simple example of these two processes. Numerous techniques exist for both image registration and data fusion. The choice and advantage of one The British Journal of Radiology, Special Issue 2006


technique over another depends on the particular application and types of image data involved. While exhaustive and detailed reviews of image registration algorithms have appeared in the literature [4], this paper is meant to provide a broad overview as well as examples of image registration and data fusion techniques that are employed in radiation therapy.

Image registration The basic task of image registration is to compute the geometric transformation that maps the coordinates of corresponding or homologous points between two imaging studies. While there are many different techniques used to carry this out, most approaches involve the same three basic components. The first and main component is the transformation model itself, which can range from a single global linear transformation for handling rotations and translations (six degrees of freedom; three rotations and three translations) to a completely free form deformation model where the transformation is represented by independent displacement vectors for each voxel in the image data (degrees of freedom can reach three times the number of voxels). The second component is the metric used to measure how well the images are (or are not) registered, and the third component is the optimizer and optimization scheme used to bring the imaging data into alignment. It is also worth mentioning that these general components, the transformation model, which defines the degrees of freedom or parameters, the metric or cost function used to measure the worth of the registration and the optimization engine used to reach a final solution, are completely analogous to the components required by inverse treatment planning systems. S99

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Figure 1. Schematic of the image registration and data fusion processes. (a) Anatomical information from a spin-echo MR is first registered and then fused with functional information from a 11C thymidine PET to create a synthetic MR-PET image volume. (b) General components of the registration process.

Although it is often desirable or necessary to register numerous imaging studies to each other, the process of registration is generally carried out by registering two datasets at a time. In radiation therapy, a common strategy is to register each of the imaging studies to the treatment planning CT, as it is used as the primary dataset for treatment planning and dose calculations. Transformations between studies that are not explicitly registered to each other can be easily derived by combining the appropriate transforms and inverse transforms between the different datasets and the planning CT. For the discussions that follow, the two datasets being registered are labelled Study A and Study B. Study A will be the base or reference dataset that is held fixed and Study B will be the homologous or moving dataset that is manipulated to be brought into geometric alignment with Study A. Study B’ will refer to the transformed or registered version of Study B (Figure 1b).

Transformation model The transformation model chosen to describe the mapping of coordinates between two studies depends on the clinical site, the imaging conditions and the particular application. In the ideal case, where the patient is positioned in an identical orientation during the different imaging studies and the scale and centre of the imaging coordinate systems coincide, the transformation is a simple identity transform I and xB5xA for all points in the two imaging studies. This situation most closely exists for the data produced by dual imaging modality devices such as PET-CT or SPECT-CT machines, especially if physiological motion is controlled or absent [6]. Naturally, it is common for the orientation of the patient to change between imaging studies, making more sophisticated transformations necessary. For situations involving the brain, where the position and orientation of the anatomy are defined by the rigid skull, a simple rotate-translate model can be accurately applied. In this case, a global linear transformation specified by three rotation angles (hx, hy, hz) and three translations (tx,ty,tz) can be used to map points from one image dataset to another. A more general linear transformation is an S100

affine transform, which is a composition of rotations, translations, scaling (sx,sy,sz) and shearing (shx,shy,shz). A property of affine transformations is that they preserve collinearity (‘‘parallel lines remain parallel’’). Currently, the DICOM imaging standard uses affine transformations to specify the spatial relationship between two imaging studies [7]. Most commercial treatment planning systems only support image registration using affine transformations, although support for more sophisticated transformations should appear soon. The assumption of global rigid movement of anatomy is often violated, especially for sites other than the head and large image volumes that extend to the body surface. Differences in patient setup (arms up versus arms down), organ filling and uncontrolled physiological motion confound the use of a single affine transform to register two imaging studies. In some cases where local rigid motion can be assumed, it may be possible to use a rigid or affine transformation to register sub-volumes of two imaging studies. For example, the prostate itself may be considered rigid, but it certainly moves relative to the pelvis, depending on the filling of the rectum and bladder. By considering only a limited field-of-view that includes just the region of the prostate, it is often possible to use an affine transformation to accurately register the prostate anatomy in two studies [8–10]. One or more subvolumes can be defined by simple geometric cropping or masks derived from one or more anatomical structures (Figure 2). Even with a limited field-of-view approach, there are many sites in which affine registration techniques are not sufficient to achieve acceptable alignment of anatomy. In these sites, an organ’s size and shape may change as a result of normal organ behaviour or the motion of surrounding anatomy. For example, the lungs change in both size and shape during the breathing cycle, and the shape of the liver can be affected by the filling of the stomach. When registering datasets that exhibit these kinds of changes, a non-rigid or deformable model must be used to accurately represent the transformation between studies. Deformable transformation models range in complexity from a simple extension of a global affine transformation using higher order polynomials with relatively few parameters, to a completely local or ‘‘free form’’ model The British Journal of Radiology, Special Issue 2006

Image registration and data fusion in radiation therapy

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Figure 2. Various strategies for cropping data for limited field-of-view image data. (a) Simple geometric cropping. (b) Piecewise cropping. (c) Anatomically-based cropping.

where each point or voxel in the image volume can move independently and the number of parameters may reach three times the number of voxels considered. Between these two extremes are transformation models designed to handle various degrees of semi-local deformations using a moderate number of parameters, such as splines [11]. Global polynomials have been used successfully to model and remove image distortions in MR and other image data as a pre-processing step for image registration [12], but are not typically used for modelling deformation of anatomy because of undesirable oscillations that occur as the degree of the polynomial increases. Spline-based transformations, such as Bsplines [11, 13] avoid this problem by building up the overall transformation, or deformation function, using a set of weighted basis functions defined over (or which contribute only over) a limited region. Figure 3 illustrates this approach for a one-dimensional cubic Bspline. The displacement or deformation, Dx, at a given point is computed as the weighted sum of basis functions

centred at a series of locations called knots. Changing the weight or contribution w of each basis function affects only a specific portion of the overall deformation. By increasing the density of knots, more complex and localized deformations can be modelled. Another spline based transformation, called thin-plate splines, uses a set of corresponding control points defined on both image datasets and minimizes a bending energy term to determine the transformation parameters [14–16]. Unlike B-splines, the location of each control point does have some amount of global influence, meaning that changing the position of a control point in one area will affect the entire deformation in some capacity. Using more points reduces the influence of each point but this comes at a higher computational cost than with B-splines. Finally, free-form or non-parametric transformation models are represented using vector fields of the explicit displacements for a grid of points, usually at the voxel locations or an integer sub-sample of these (Figure 4). Algorithms for solving for the displacements with

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Figure 3. B-spline deformation model. (a) 1D example of the cubic B-spline deformation model. The displacement Dx as a function of x is determined by the weighted sum of basis functions. The double arrow shows the region of the overall deformation affected by the weight factor w7. 3D deformations are constructed using 1D deformations for each dimension. (b) Multiresolution registration of lung data using B-splines. Both knot density and image resolution are varied during registration. This can help avoid local minima and decrease overall registration time.


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Figure 4. Visualization of (a) deformation computed between datasets registered using B-splines and (b) fluid flow model. The deformation or displacement is known for every voxel but only displayed for a subset of voxels for clarity ((b) image courtesy of Gustavo Olivera, University of Wisconson).

non-parametric models use some form of local driving force to register the image data. Common models include fluid flow [17, 18], optical flow (based on intensity gradients) [19, 20] and finite element methods [21].

Registration metric In most registration algorithms, the parameters of a transformation model which bring two datasets into geometric alignment are determined by maximizing or minimizing a registration metric which measures the similarity or dissimilarity of the two image datasets. Most registration metrics in use today can be classified as either geometry-based or intensity-based. Geometrybased metrics make use of features extracted from the image data, such as anatomic or artificial landmarks and organ boundaries, while intensity-based metrics use the image data directly.

Geometry-based metrics The most common geometry-based registration metrics involve the use of points [22], lines [23, 24] or surfaces [22, 25, 26]. For point matching, the coordinates of pairs of corresponding points from Study A and Study B are used to define the registration metric. These points can be anatomic landmarks or implanted or externallyplaced fiducial markers. The registration metric is defined as the sum of the squared distances between corresponding points. To compute the rotations and translations for a rigid transformation, a minimum of three pairs of points are required and for affine transformations, a minimum of four pairs of noncoplanar points are required. Using more pairs of points reduces the bias that errors in the delineation of any one pair of points has on the estimated transformation parameters. However, accurately identifying more than the minimum number of corresponding points can be difficult as different modalities often produce different tissue contrasts (a major reason why multiple modalities are used in the first place) and placing or implanting larger numbers of markers is not always possible or desirable. Alternatively, line and surface matching techniques do not require a one to one correspondence of specific S102

points, but rather try to maximize the overlap between corresponding lines and surfaces extracted from two image studies, such as the brain or skull surface or pelvic bones. These structures can be easily extracted using automated techniques and minor hand editing. As with defining pairs of points, it may be inherently difficult or time consuming to accurately delineate corresponding lines and surfaces in both imaging studies. Furthermore, since the extracted geometric features are surrogates for the entire image volume, any anatomic or machine-based distortions in the image data away from these features will not be taken into account during the registration process.

Intensity-based metrics To overcome some of the limitations of using explicit geometric features to register image data, another class of registration metric has been developed which uses the numerical greyscale information directly to measure how well two studies are registered. These metrics are also referred to as similarity measures since they determine the similarity between the distributions of corresponding voxel values from Study A and a transformed version of Study B. Several mathematical formulations are used to measure this similarity. The more common similarity measures in clinical use include: sum-of-squared differences and cross-correlation [27] for registration of data from X-ray CT studies and mutual information for registration of data from both similar and different imaging modalities [15, 28, 29]. The mutual information metric provides a measure of the information that is common between two datasets [30]. It is assumed that when two datasets are properly aligned, the mutual information of the pair is a maximum, which makes it an appropriate registration metric. It can be used for a wide range of image registration situations since there is no dependence on the absolute intensity values and it is very robust to missing or limited data. For example, a tumour might show up clearly on an MR study but be indistinct on a corresponding CT study. Over the tumour volume the mutual information is low, but no prohibitive penalties are incurred. In the surrounding healthy tissue the mutual information can be high, and this becomes the dominant factor in the registration. The British Journal of Radiology, Special Issue 2006


Optimizer and registration scheme Most image registration systems use optimization schemes such as gradient descent or problem specific adaptations of these. Registration of datasets is usually carried out in a hierarchical fashion, starting with downsized versions of the data and iteratively registering successively finer versions. The degrees of freedom of the geometric transformation can also be varied to speed the registration process. An example scheme might begin with simple translations, and then allow rotations, then low spatial frequency deformations and finally the full deformation model [12]. A hierarchical approach saves computation time and also helps avoid local minima, which become more likely as the degrees of freedom of the deformation model increase. For deformable image registration problems using a large number of degrees of freedom, some form of regularization may also be imposed to discourage ‘‘unreasonable’’ deformations such as warping of bones and folding of tissue. One approach to this problem is to filter the deformations between iterations of the optimization [31]. Another approach is to include a regularization term in the registration metric that penalizes non-physical deformations. The regularization term can even be made spatially variant using known or estimated tissue properties [32].

Data fusion The motivation for registering imaging studies is to be able to map information derived from one study to another, or to directly combine or fuse the imaging data from the studies to create displays that contain relevant features from each modality. For example, a tumour volume may be more clearly visualized using a specific MR image sequence or coronal image plane rather than the axial treatment planning CT. If the geometric transformation between the MR study and the treatment planning CT study is known, the clinician is able to outline the tumour using images from the MR study and

map these outlines to the images of the CT study. This process is called structure mapping (Figure 5). Another approach to combining information from different imaging studies is to map directly the image intensity data from one study to another so that at each voxel there are two (or more) intensity values rather than one. The goal is to create a version of Study B (Study B’) with images that match the size, location and orientation of those in Study A. These corresponding images can then be combined or fused in various ways to help elucidate the relationship between the data from the two studies. Various relevant displays are possible using this multistudy data. For example, functional information from a PET imaging study can be merged with the anatomic information from an MRI study and displayed as a colourwash overlay (Figure 6). This type of image synthesis is referred to as image fusion. A variety of techniques exist to present fused data, including the use of overlays, pseudo-colouring and modified greyscales. For example, the hard bone features of a CT imaging study can be combined with the soft tissue features of an MRI study by adding the bone extracted from the CT to the MR dataset. Another method is to display anatomic planes in a side-by-side fashion (Figure 6). Such a presentation allows structures to be defined using both images simultaneously. In addition to mapping and fusing image intensities, 3D dose distributions computed in the coordinate system of one imaging study can be mapped to another. For example, doses computed using the treatment planning CT can be displayed over an MR study acquired after the start of therapy. With these data, regions of posttreatment radiological abnormality can be readily compared with the planned doses for the regions. With the introduction of volumetric imaging on the treatment units, treatment delivery CT studies can now be acquired to determine more accurately the actual doses delivered. By acquiring these studies over the course of therapy and registering them to a common reference frame, doses for the representative treatments can be reformatted and accumulated to provide a more likely estimate of the

Figure 5. Structure mapping. A tumour volume is outlined by the clinician on an MR study and then mapped to the treatment planning CT using the computed transformation. The British Journal of Radiology, Special Issue 2006

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Figure 6. Different approaches to display data from multiple studies which have been registered and reformatted. (a) Side-byside display with linked cursor. (b) Split screen display. (c) Colourwash overlay.

Regardless of the output of any numerical technique used, which may only be a single number, it is important for the clinician to appreciate how well in three dimensions the information they define on one study is mapped to another. There are many possible visualization techniques to help to evaluate qualitatively the results of a registration. Most of these are based on data mapping and fusion display techniques. For example, paging through the images of a split screen display and moving a horizontal or vertical divider across regions where edges of structures from both studies are visible can help uncover even small areas of misregistration (Figure 7). Another interesting visual technique involves dynamically switching back and forth between corresponding images from the different studies at about once per second and focusing on particular regions of the anatomy to observe how well they are aligned. In addition to comparing how well the images from Study A and Study B’ correspond at the periphery of anatomic tissues and organs, outlines from one study can be displayed over the images of the other. Figure 8 shows a brain surface which was automatically segmented

delivered dose. This type of data can be used as input into the adaptive radiotherapy decision process.

Validation It is important to validate the results of a registration before making clinical decisions based on the results. To do this, most image registration systems provide some combination of numerical and visual verification tools. A common numerical evaluation technique is to define a set of landmarks for corresponding anatomic points on Study A and Study B and compute the distance between the actual location of the points defined on Study A and the resulting transformed locations of the points from Study B’. This calculation is similar to a ‘‘point matching’’ metric but, as discussed earlier, it may be difficult to accurately and sufficiently define the appropriate corresponding points, especially when registering multimodality data. Also, if deformations are involved, the evaluation is not valid for regions remote from the defined points.

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Figure 7. Image-image visual validation using split screen displays of native MR and reformatted CT study. S104



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Figure 8. Image-geometry visual validation structure overlay of CT defined brain outlines over MR images.

from the treatment planning CT study and mapped to the MR study. The agreement between the CT-based outlines at the different levels and planes of the MR study demonstrate the accuracy of the registration. In practice, the accuracy of the registration process depends on a number of factors. For multimodality registration of PET/CT/MR data in the brain, registration accuracy on the order of a voxel size of the imaging studies can be achieved. Outside the head, many factors confound single voxel level accuracy, such as machine induced geometric and intensity distortions as well as dramatic changes in anatomy and tissue loss or gain. Nevertheless, accuracy at the level of a few voxels is certainly possible in many situations.

Clinical applications Image registration and data fusion are useful at each step of the patient management process in radiation therapy; for initial diagnosis and staging, during treatment planning and delivery, and after therapy to help monitor the patient’s response. The overall purpose of these tools at each stage is the same; to help to integrate the information from different imaging studies in a quantitative manner to create a more complete representation of the patient. Over the past several years, treatment planning and treatment delivery systems have evolved to provide direct support for image registration and data fusion. Typical examples of how these techniques are used for treatment planning, delivery, and adaptation are described here.

Treatment planning Most modern treatment planning systems permit the use of one or more datasets in addition to the treatment planning CT for structure delineation and visualization. These are sometimes referred to as ‘‘secondary’’ datasets. In order to transfer anatomic outlines and other The British Journal of Radiology, Special Issue 2006

geometric information from these datasets to the planning CT, the transformation between the secondary dataset and the planning CT is required. Furthermore, using the inverse of this transformation, it is also possible to transfer information computed using the planning CT, such as the planned dose, to the secondary dataset. Incorporation of secondary or complementary data from MRI and nuclear medicine imaging studies is becoming increasingly common. MR provides superior soft tissue contrast relative to CT and the ability to image directly along arbitrary planes can aid in the visualization and delineation of certain anatomic structures, such as the optic nerves and chiasm. MR can also provide information on localized metabolite concentrations using spectroscopy [3, 33]. Incorporation of functional information from PET and SPECT can help remove ambiguities that might exist on the treatment planning CT between the tumour and other conditions such as atelectasis and necrosis [34]. These studies can also indicate nodal involvement and provide a map of local tissue function that can be used construct objective functions for dose optimization [3, 35]. Figure 9 illustrates the use of MR as a secondary dataset for target and normal structure delineation. An axial and coronal MR study was acquired and registered to the treatment planning CT using a geometric transformation which allowed rotations and translations, as the anatomy in this region moves in a rigid fashion. Since the image data were from different modalities, the mutual information registration metric was used. Splitscreen visualization of the registered datasets was used to validate the computed transformation which was judged to be accurate to within 1–2 mm over the image volume (Figure 7). The gross tumour volume (GTV) was defined as the region of enhancement in the post GdDPTA contrast MR studies. The clinician outlined this volume on both the axial and coronal sets of MR images. The optic nerves and chiasm were outlined on the coronal MR study. The outlines were used to generate a 3D surface description for each tissue and these were mapped to the coordinate system of the planning CT S105

M L Kessler

Figure 9. Incorporation of MR image data into the treatment planning process.

using the computed transformation. The outlines of these mapped surfaces for each CT image were derived by intersecting the transformed surfaces along the planes defined by each image (Figure 5). Because of differences in partial volume averaging between the axial and coronal MR images, the outlines derived from the axial and coronal MR data are not identical. In these cases, the clinician has the choice to use one or other outline or to generate a composite outline using a Boolean OR operation. For this example, the CT greyscale data did not contribute any information for the definition of the GTV, optic nerves, or optic chiasm. Had the physician outlined a CT-based GTV, it could have been incorporated directly into the composite GTV or compared with the MR based outlines to reconcile potential ambiguities. At this point, the outlines of the tumour and normal structures were used for treatment planning as if they were derived using only the planning CT. The final planning target volume (PTV) was created by uniformly expanding the composite GTV surface by 5 mm to account for setup uncertainty. A treatment plan and dose distribution was generated using the CT data, PTV and normal structures. The CT-based dose distribution was then mapped back to the MR study for further visualization of the dose relative to the underlying anatomy.

Treatment delivery Once a treatment plan is created, it is transferred to the treatment unit for delivery. The location and orientation of the patient on the treatment machine must be adjusted so that the centre and orientation of the coordinate system of the treatment plan coincide with that of that of the treatment unit. Image registration is typically used to carry out this process using images acquired in the treatment room and the planning CT. The most common practice is to generate a pair of orthogonal digitally reconstructed radiographs (DRRs) from the planning CT and register these simulated radiographs with actual radiographs acquired by a flat-panel imager attached to the treatment unit. It is now also possible to acquire S106

volumetric image data at the treatment unit using conebeam reconstruction of a set of projection images acquired by rotating the treatment gantry around the patient. (See papers by Kirby and Glendinning, Moore et al and Chen and Pouliot in this issue). These cone-beam data can be registered directly with the planning CT to determine how to shift (and possible rotate) the treatment table to properly position the patient for treatment [36, 37]. Figure 10 shows an example of an interface for 3D image-based alignment at the treatment unit using a cone-beam CT dataset and the planning CT. Automated image registration using successively finer data resolution and mutual information are used to determine the rotations and translations required to align the two datasets. These are then translated into machine parameters which can be automatically downloaded to the treatment unit and set-up. In this example, the accuracy of the registration is assessed using both image-image and structure-image overlay displays. These same tools and image data are also available off-line so that the clinician can track and analyse the progress of the treatments.

Treatment adaptation and customization On-line imaging has made it more convenient to acquire image data of the patient over the course of treatment. Using these data, it is possible to uncover changes in patient anatomy or treatment setup that are significant and dictate changes to the original treatment plan. Better estimates of individual treatment doses can be computed using these data and the actual machine parameters. By registering these data to the ‘‘base’’ treatment planning CT, it is possible to construct a more complete model of the accumulated dose to the patient. This information can then be used to assess if and how a treatment plan should be adapted or further customized [38–41]. Figure 11 shows an example of dose accumulation for two datasets of the patient at different points in the breathing cycle. The dose distribution displayed on the left image was computed directly using the image dataset shown [42]. The dose distribution displayed on The British Journal of Radiology, Special Issue 2006


Figure 10. Volumetric registration at the treatment unit. A cone-beam CT acquired at the time of treatment is registered to the treatment planning CT (larger dataset) to properly position the patient on the treatment table (courtesy of Peter Monroe, PhD, Varian Medical Systems).

the middle image was computed from another dataset and mapped to the image data shown using the B-spline deformation field computed by registering the two datasets using a sum-of-squares difference metric. The dose distribution on the right is the weighted sum of the two distributions. This process can be continued throughout the course of therapy to provide up-to-date information on the delivered dose.

Summary Over the past several years there has been an explosion of the use of image data from a variety of modalities to aid in treatment planning, delivery and evaluation. In order to make quantitative use of these data it is necessary to determine the transformation that relates the coordinates of the individual datasets to one another. The process of finding this transformation is referred to as image registration. Once the geometric relationship between the datasets has been determined it is possible to utilize the information they provide by mapping the image data, derived structures, and computed dose distributions between datasets using a process called data fusion. There are many different

techniques that have been and are being studied to improve the accuracy and utility of both image registration and data fusion. Both processes are now essential components for modern treatment planning and delivery. As the need, availability and diversity of image data continues to increase, they will be even more important to each part of patient management process. These tools, however, can not replace clinical judgment. Different imaging modalities image the same tissues differently and, although tools may help us to understand better and differentiate between tumour and non-tumour, they cannot yet make the ultimate decision of what to treat and what to not treat. These decisions still lie with the clinician, although they now have more sophisticated tools to help them make these decisions.

Acknowledgments Portions of the text and some of the figures presented here were published previously in Kessler ML, Roberson M. Image registration and data fusion for radiotherapy treatment planning. In: Schlegel W, Bortfeld T, Grosu A-L, editors. New technologies in radiation oncology. Springer, 2006.

Figure 11. Example of dose summation/accumulation using registered datasets (courtesy of Mihaela Rosu, The University of Michigan).


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References 1. Webb S. The physical basis of IMRT and inverse planning. Br J Radiol 2003;76:678–89. 2. Eisbruch A. Intensity-modulated radiation therapy: a clinical perspective. Semin Radiat Oncol 2002;12:197–8. 3. Ling CC, Humm J, Larson S, Amols H, Fuks Z, Leibel S, et al. Towards multidimensional radiotherapy (MD-CRT): biological imaging and biological conformality. Int J Radiat Oncol Biol Phys 2000;47:551–60. 4. Maintz JB, Viergever MA. A survey of medical image registration. Med Image Anal 1998;2:1–36. 5. Hill DL, Batchelor PG, Holden M, Hawkes DJ. Medical image registration. Phys Med Biol 2001;46:R1–R45. 6. Townsend DW, Beyer T. A combined PET/CT scanner: the path to true image fusion. Br J Radiol 2002;75:24S–30S. 7. DICOM Part 3, PS3.3 – Service Class Specifications, National Electrical Manufacturers Association, Rosslyn, Virgina, USA, 2004. 8. McLaughlin PW, Narayana V, Meriowitz A, Troyer S, Roberson PL, Gonda R Jr, et al. Vessel sparing prostate radiotherapy: dose limitation to critical erectile vascular structures (internal pudendal artery and corpus cavernosum) defined by MRI. Int J Radiat Oncol Biol Phys 2005;61:20–31. 9. McLaughlin PW, Narayana V, Kessler M, McShan D, Troyer S, Marsh L, et al. The use of mutual information in registration of CT and MRI datasets post permanent implant. Brachytherapy 2004;3:61–70. 10. Roberson PL, McLaughlin PW, Narayana V, Troyer S, Hixson GV, Kessler ML. Use and uncertainties of mutual information for CT/MR registration post permanent implant of the prostate. Med Phys 2005;32:473–82. 11. Unser M, Aldroubi A, Eden M. B-spline signal processing: part I-theory. IEEE Trans Signal Processing 1993;41:821–33. 12. Kybic J, Unser M. Fast parametric elastic image registration. IEEE Trans Image Processing 2003;75:1427–42. 13. Maurer CR, Aboutanos GB, Dawant BM, et al. Effect of geometrical distortion correction in MR on image registration accuracy. J Comput Assist Tomogr 1996;20:666–79. 14. Bookstein F. Principal warps: thin-plate splines and the decomposition of deformations. IEEE Trans Pattern Analysis Machine Intelligence 1989;567–85. 15. Meyer CR, Boes JL, Kim B, Bland PH, Zasadny KR, Kison PV, et al. Demonstration of accuracy and clinical versatility of mutual information for automatic multimodality image fusion using affine and thin-plate spline warped geometric deformations. Med Image Anal 1997;1:195–206. 16. Coselmon MM, Balter JM, McShan DL, Kessler ML. Mutual information based CT registration of the lung at exhale and inhale breathing states using thin-plate splines. Med Phys 2004;31:2942–8. 17. Lu W, Chen M, Olivera GH, Ruchala KJ, Mackie T. Fast free-form deformable registration via calculus of variations. Phys Med Biol 2004;49:3067–87. 18. Christensen GE, Carlson B, Chao KS, Yin P, Grigsby PW, Nguyen K, et al. Image-based dose planning of intracavitary brachytherapy: registration of serial-imaging studies using deformable anatomic templates. Int J Radiat Oncol Biol Phys 2001;51:227–43. 19. Thirion JP. Image matching as a diffusion process: an analogy with Maxwell’s demons. Med Image Anal 1998;2:243–60. 20. Wang H, Dong L, Lii MF, Lee AL, de Crevoisier R, Mohan R, et al. Implementation and validation of a three-dimensional deformable registration algorithm for targeted prostate cancer radiotherapy. Int J Radiat Oncol Biol Phys 2005;61:725–35. 21. Brock KK, Sharpe MB, Dawson LA, Kim SM, Jaffray DA. Accuracy of finite element model (FEM)-based multi-organ deformable image registration. Med Phys 2005;32:1647–59.

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22. Kessler ML, Pitluck S, Petti PL, Castro JR. Integration of multimodality imaging data for radiotherapy treatment planning. Int J Radiat Oncol Biol Phys 1991;21:1653–67. 23. Balter JM, Pelizzari CA, Chen GT. Correlation of projection radiographs in radiation therapy using open curve segments and points. Med Phys 1992;19:329–34. 24. Langmack KA. Portal imaging. Br J Radiol 2001;74:789–804. 25. Pelizzari CA, Chen GT, Spelbring DR, Weichselbaum RR. Accurate three-dimensional registration of CT, PET, and/or MR images of the brain. J Comput Assist Tomogr 1989;13:20–6. 26. van Herk M, Kooy HM. Automatic three-dimensional correlation of CT-CT, CT-MRI, and CT-SPECT using chamfer matching. Med Phys 1994;21:1163–78. 27. Kim J, Fessler JA. Intensity-based image registration using robust correlation coefficients. IEEE Trans Med Imaging 2004;23:1430–44. 28. Viola P, Wells WM. Alignment by maximization of mutual information. Int J Computer Vision 1997;137–54. 29. Maes F, Collignon A, Vandermeulen D, Marchal G, Suetens P. Multimodality image registration by maximization of mutual information. IEEE Trans Med Imaging 1997;16:187–98. 30. Roman S. Introduction to coding and information theory, Undergraduate Texts in Mathematics, ISBN 0-387-94704-3, New York, NY: Springer-Verlag, 1997. 31. Staring M, Klein S, Pluim JP. Nonrigid registration with adaptive, content-based filtering of the deformation field. Proc SPIE Medical Imaging 2005: Image Proc. 2005:212–21. 32. Ruan R, Fessler JA, Roberson M, Balter J, Kessler M. Nonrigid registration using regularization that accommodates for local tissue rigidity. Proc SPIE Medical Imaging Proc. Vol 6144, 2006. (In press). 33. Graves EE, Pirzkall A, Nelson SJ, Larson D, Verhey L. Registration of magnetic resonance spectroscopic imaging to computed tomography for radiotherapy treatment planning. Med Phys 2001;28:2489–96. 34. Munley MT, Marks LB, Scarfone C, Sibley GS, Patz EF Jr, Turkington TG, et al. Multimodality nuclear medicine imaging in three-dimensional radiation treatment planning for lung cancer: challenges and prospects. Lung Cancer 1999;23:105–14. 35. Marks LB, Spencer DP, Bentel GC, et al. The utility of SPECT lung perfusion scans in minimizing and assessing the physiologic consequences of thoracic irradiation. Int J Radiat Oncol Biol Phys 1993;26:659–68. 36. Jaffray DA, Siewerdsen JH, Wong JW, Martinez AA. Flatpanel cone-beam computed tomography for imageguided radiation therapy. Int J Radiat Oncol Biol Phys 2002;53:1337–49. 37. Mackie TR, Kapatoes J, Ruchala K, Lu W, Wu C, Olivera G, et al. Image guidance for precise conformal radiotherapy. Int J Radiat Oncol Biol Phys 2003;56:89–105. 38. Smitsmans MH, Wolthaus JW, Artignan X, de Bois J, Jaffray DA, Lebesque JV, et al. Automatic localization of the prostate for on-line or off-line image-guided radiotherapy. Int J Radiat Oncol Biol Phys 2004;60:623–35. 39. Yan D, Wong J, Vicini F, Michalski J, Pan C, Frazier A, et al. Adaptive modification of treatment planning to minimize the deleterious effects of treatment setup errors. Int J Radiat Oncol Biol Phys 1997;38:197–206. 40. Yan D, Lockman D, Martinez A, Wong J, Brabbins D, Vicini F, et al. Computed tomography guided management of interfractional patient variation. Semin Radiat Oncol 2005;3:168–79. 41. Lam KL, Ten Haken RK, Litzenberg D, Balter JM, Pollock SM. An application of Bayesian statistical methods to adaptive radiotherapy. Phys Med Biol 2005;50:3849–58. 42. Rosu M, Chetty IJ, Balter JM, Kessler ML, McShan DL, Ten Haken RK. Dose reconstruction in deforming lung anatomy: dose grid size effects and clinical implications. Med Phys 2005;32:2487–95.


The British Journal of Radiology, 79 (2006), e1–e4

CASE REPORT

Pyourachus: study of two cases 1

R B THAPAR,

DMRD, DMRE,

2

V U JHA,

MS,

1

R U MEHTA,

MD

and 2G R SHAH,

MS, MCh

Departments of 1Radiology and 2Urology, Sir Hurkisondas Nurrotamdas Hospital & Research Center, Raja RamMohan Roy Road, Mumbai – 400004, India ABSTRACT. The urachus, or median umbilical ligament, is a midline tubular structure that extends upward from the anterior dome of the bladder toward, the umbilicus and represents the vestigial remnant of at least two embryonic structures, the cloaca and the allantois. The tubular urachus normally involutes before birth, remaining as a fibrous band, however its persistence can give rise to various clinical problems, not only in infants and children but also in adults. We report two cases of pyourachus at our institute with a review of the clinical presentation, imaging findings and surgical management. Both our patients were young males, with haematuria being the presenting feature in one case which has not been previously described in literature.

The urachus is a vestigial remnant of the cloaca and the allantois which is present after birth as a midline fibrous cord extending from the anterior dome of the urinary bladder to the umbilicus. The abnormal persistence of the urachal patency may result in various clinical problems at any age. Due to their rarity, these urachal remnant diseases are frequently misdiagnosed and confused with a wide spectrum of midline intraabdominal or pelvic disorders. A detailed understanding of the embryological basis of these urachal disorders and their imaging features is crucial for the correct diagnosis and management. We report two cases of pyourachus at our institute with a brief review of the embryology, anatomy, clinical presentation, imaging findings and surgical management.

Case reports Case 1 A 19-year-old male, with a known case of stricture urethra at the bulbomembranous junction, presented to our institute with pain and fullness in the suprapubic region for the last 3 days. The routine blood investigations revealed a total leukocyte count of 12 800 cm23 and an elevated serum creatinine of 3 mg dl21. The patient then underwent ultrasound of abdomen and pelvis, which revealed bilateral hydronephrosis and hydroureter with a well localized collection above the dome of the bladder, immediately beneath the anterior abdominal wall. This collection showed internal echoes suggestive of turbid contents (Figure 1). The contrast enhanced computed tomography (CECT) scan of abdomen and pelvis revealed bilateral hydronephrosis and hydroureter with a well localized, thick Address correspondence to: Dr Ravikumar B Thapar, 302, Amar Residency, Punjabwadi, S.T.Road, Deonar, Mumbai – 400088, India.


Received 14 January 2005 Revised 1 May 2005 Accepted 27 June 2005 DOI: 10.1259/bjr/96369737 ’ 2006 The British Institute of Radiology

walled, rim enhancing collection in the suprapubic region extending from the anterior dome of the bladder to the umbilicus with adjacent fat stranding (Figure 2). The central portion of the collection had attenuation value of around 45 Hounsfield units. The delayed and prone scans after bladder opacification did not reveal any communication between the bladder and this collection. The anatomical location of this collection correlated with the course of the urachus which was well demonstrated on the sagittal multiplanar reconstruction (MPR) images (Figure 3) and, hence, a diagnosis of infected collection within the urachal remnant (Pyourachus) was suggested. The patient underwent a cystoscopy with visual internal urethrectomy (Cysto VIU) for the urethral stricture. The cystoscopy revealed a grossly trabeculated bladder with an extrinsic indentation on the superior wall showing petechial haemorrhages. A small opening was seen but could not be negotiated. Laparotomy was then performed at the same sitting and the abscess cavity was dissected from umbilicus to the bladder and excised en bloc, along with a cuff of the bladder wall (Figure 4). The histopathology report confirmed the diagnosis of urachal abscess.

Case 2 A 25-year-old male presented with complaints of haematuria and burning micturation over the past 9 days and fever with chills for 3 days. He had a past history of pulmonary Koch’s and had been on antituberculous chemotherapy for 3 months. The blood investigations revealed a total leukocyte count of 6300 cm23, elevated erythrocyte sedimentation rate (ESR) of 64 mm and a normal serum creatinine of 1.1 mg dl21. Routine urine examination showed 80– 100 RBCs / hpf and 10–12 pus cells / hpf. The frontal chest radiograph revealed bilateral hilar adenopathy. e1

R B Thapar, V U Jha, R U Mehta and G R Shah

Figure 3. Sagittal multiplanar reconstruction (MPR) image delineates the classical ‘‘cone-shaped’’ anatomical extent of the collection along the course of the urachus with its tip at the umbilicus and the base over the bladder dome.

Ultrasound of the abdomen and pelvis was then performed, which showed a 2 cm 6 2 cm mixed echogenic polypoidal lesion in the fundus of the bladder. No evidence of calcification was seen. CECT scan of

abdomen and pelvis revealed a focal necrotic nodular lesion involving the anterosuperior wall of the urinary bladder, projecting into the anterior perivesical space and the bladder lumen (Figure 5). There was no evidence of calcification or lymphadenopathy. The kidneys and ureters were normal. Considering the location and morphology of the lesion, a differential diagnosis of infected urachal remnant or an urachal neoplasm was considered. Cystoscopy revealed an indentation on the dome of the bladder with petechial haemorrhages in the overlying mucosa (Figure 6). The patient was taken up for

(a)

(b)

Figure 1. Ultrasound of the pelvis showing a well localized hypoechoic collection indenting the bladder dome.

Figure 2. (a) Contrast enhanced CT (CECT) scan at the level of bladder dome showing a thick walled, rim enhancing collection related to the anterosuperior aspect of the bladder. Note the relatively high density of the collection as compared with the urine in the bladder, suggestive of infected turbid contents. (b) CECT scan at the level of the umbilicus shows the heterogeneously enhancing collection with adjacent fat stranding extending up to the umbilicus. e2


Case report: Pyourachus

Figure 4. Intraoperative photograph showing the infected urachus being excised en bloc.

laparotomy. The urachus was identified and traced to the mass, which was excised en bloc along with the urachus and 1 cm margin of the bladder. The histopathology analysis revealed an urachal cyst abscess with no evidence of malignancy or tuberculosis.

Discussion The urachus is an embryonic remnant resulting from the involution of the allantoic duct and the ventral cloaca [1]. The obliterated urachus, also known as the median umbilical ligament, extends from the anterior dome of

Figure 5. Contrast enhanced CT scan at the level of the bladder shows a focal, heterogeneously enhancing, nodular lesion along the anterosuperior aspect of the bladder.


Figure 6. Cystoscopy showing an indentation on the bladder dome with petechial haemorrhages in the overlying mucosa.

the bladder towards the umbilicus. It measures 3–10 cm in length, 8–10 mm in diameter and lies between the transversalis fascia and the peritoneum, in the space of Retzius. The urachus is encased between the two layers of the umbilicovesical fascia, which tends to contain the spread of urachal disease. The urachus is a muscular tube, and three distinct tissue layers are recognized: (1) an epithelial canal with transitional (70%) or columnar (30%) epithelium; (2) a submucosal connective tissue layer; and (3) an outer layer of smooth muscle that is in continuity with the detrusor muscle and thickest near the bladder [2]. However, once the urachus becomes a fibrous cord, the urachal layers are generally not recognizable. The urachal remnant diseases may be classified as congenital urachal anomalies and acquired urachal remnant diseases which include infection and neoplasm. The congenital urachal anomalies are twice as common in males as in females and 4 types are recognized [3]. 1. patent urachus (about 50% of cases); 2. urachal cyst (about 30% of cases); 3. urachal sinus (about 15% of cases); and 4. urachal diverticulum (about 5% of cases). Patent urachus is usually recognized in the neonate due to discharge of urine through the umbilicus which may occasionally be minimal or intermittent. It is associated with urethral atresia or obstructive posterior urethral valves in 15–30% of patients, in whom it serves as a protective mechanism [4]. However, the other congenital urachal anomalies are generally asymptomatic, the clinical presentation being usually associated with a superadded infection [5]. An urachal cyst develops if the urachus closes both at the umbilicus and at the bladder, but remains patent between these two areas. It develops most commonly in the distal third of the urachus and the majority of them are infected at the time of diagnosis. Therefore, they usually present with lower abdominal pain, fever, voiding symptoms, midline hypogastric tenderness, palpable mass and evidence of urinary infection [2]. Haematuria as a presentation of urachal cyst as seen in case 2 is very rare. Infected cysts occur more commonly e3

R B Thapar, V U Jha, R U Mehta and G R Shah

in adults and only occasionally in children. Non-infected urachal cysts may become symptomatic when they enlarge or may be detected incidentally during routine examination. In an excellent review of urachal remnant diseases, Yu et al have detailed the spectrum of CECT and ultrasound findings [5]. The preferred method of diagnostic imaging is ultrasound, which will show a fixed, midline, cystic, extraperitoneal mass between the umbilicus and the bladder [5]. If there is superadded infection, ultrasound will demonstrate wall thickening, internal debris and complex echogenicity. The CECT characteristics of infected urachal cyst or pyourachus include (a) midline location deep to the rectus abdominis muscle; (b) conical shape extending from a tip at the umbilicus to a base over the bladder dome; (c) peripheral inflammatory changes in subcutaneous tissues, rectus abdominis muscle and mesenteric fat; and (d) intraperitoneal fluid or abscess (if perforation has occurred) [7]. The attenuation of the purulent cyst contents will be higher than that of water. MRI does not usually provide any significant additional information while intravenous urography and voiding cystourethrography may only show an extrinsic impression on the fundus of the bladder. CT or ultrasound guided diagnostic percutaneous aspiration may be performed. However, this does not play a therapeutic role due to the high recurrence rate. The organism most commonly cultured from the cyst fluid is Staphylococcus aureus [8]. The infected urachal cyst may drain along the urachus and result in acquired forms of urachal sinus (drainage through the umbilicus), urachal diverticulum (drainage through the bladder), or alternating sinus where drainage shifts alternately between the umbilicus and the bladder. If the drainage is at the umbilicus, a fistulogram in the lateral view is preferable, while if drainage at the bladder is suspected, a cystogram in the lateral view will better reveal the disease [4]. Other serious complications of an infected urachal cyst include rupture into the peritoneal cavity causing significant peritonitis and, rarely, inflammatory involvement of the adjacent bowel and formation of an enteric fistula [9]. Besides infection, other complications of an urachal cyst include urinary retention, haemorrhage and an increased incidence of adenocarcinoma. If a cyst harbouring a mucinous adenocarcinoma ruptures, pseudomyxoma peritoneii may result [10]. The complex echogenicity at ultrasound and heterogeneous attenuation and contrast enhancement on CECT may cause confusion between an infected urachal remnant and urachal carcinoma, as in case 2. Urachal carcinomas are usually mucinous adenocarcinomas and show punctate, stippled or peripheral curvilinear calcification in 50–70% of cases [11]. The presence of haematuria, mural nodularity and calcification at

e4

CECT, along with lack of adjacent inflammatory change, help to differentiate urachal malignancy from an infected cyst [5]. In such cases, where imaging is unable to differentiate these two entities, CT or ultrasound guided percutaneous needle biopsy or fluid aspiration is mandatory for diagnosis and therapeutic planning. The management of an urachal cyst is usually surgical. Primary excision is the treatment of choice for uninfected cyst. In case of infected cysts, either a one stage excision or two stage drainage with subsequent total excision including a cuff of bladder may be performed. The twostaged approach helps to limit the amount of bladder wall resected and reduces the risk of injury to the adjacent intraperitoneal structures [2]. Total excision of the cyst wall is essential to eliminate the risk of future reinfection and carcinoma. Recently, endoscopic approach for the excision of urachal cyst has been described [12].

References 1. Moore KL. The urogenital system. In: Moore KL, editor. The developing human. 3rd edn. Philadelphia, PA: Saunders, 1982:255–97. 2. Gearhart JP. Exstrophy, epispadias and other bladder anomalies. In: Walsh PC, Retik AB, Vaughan ED, Wein AJ, editors. Campbell’s urology. 8th edn. Philadelphia, PA: Saunders, 2002:2136–96. 3. Mesrobian HG, Zacharias A, Balcom AH, Cohen RD. Ten years experience with isolated urachal anomalies in children. J Urol 1997;158:1316–8. 4. Friedland GW, Devries PA, Matilde NM. Congenital anomalies of the urachus and bladder. In: Pollack HM, McClennan BL, editors. Clinical urography. 2nd edn. Philadelphia, PA: Saunders, 2000:826–51. 5. Yu JS, Kim KW, Lee HJ, Lee YJ, Yoon CS, Kim MJ. Urachal remnant diseases: spectrum of CT and US findings. Radiographics 2001;21:451–61. 6. Khati NJ, Enquist EG, Javitt MC. Imaging of the umbilicus and periumbilical region. Radiographics 1998;18:413–31. 7. Herman TE, Shackelford GD. Pyourachus: CT manifestations. J Comput Assist Tomogr 1995;19:440–3. 8. McMillan RW, Schulinger JN, Santucci VT. Pyourachus: an unusual surgical problem. J Pediatr Surg 1973;8:87–9. 9. Berman SM, Tolia BM, Laor E, Reid RE, Schweizerhof SP, Freed SZ. Urachal remnants in adults. Urology 1988;131:17–21. 10. Sasano H, Shizawas S, Nagura H, Yamaki T. Mucinous adenocarcinomas arising in a giant urachal cyst associated with pseudomyxoma peritoneii and stromal osseous metaplasia. Pathol Int 1997;47:502–5. 11. Brick SH, Friedman AC, Pollack HM, Fishman EK, Radecki PD, Siegelbaum MH, et al. Urachal carcinoma: CT findings. Radiology 1988;169:377–81. 12. Tornero JI, Ponce de Leon J, Huguet J, Rosales A, Caparros J, Villavicencio H. Endoscopic approach of the overinfected urachal cyst. Int Urol Nephrol 2002;34:289–91.



CASE REPORT

Diagnosis of tetralogy of Fallot with anatomically corrected malposition of the great arteries and single coronary artery by multidetector CT 1

A KHOSITSETH,

MD,

2

R PORNKUL,

MD

and 2S SIRIPORNPITAK,

MD

1

Division of Pediatric Cardiology, Department of Pediatrics and 2Department of Radiology, Faculty of Medicine, Ramathibodi Hospital, Mahidol University, Bangkok 10400, Thailand

ABSTRACT. We present a case of an 8-year-old boy diagnosed as a tetralogy of Fallot with anatomically corrected malposition of the great arteries by echocardiography and cardiac catheterization. Multidetector CT nicely elucidated the course of the single coronary artery from the right coronary cusp, which traverses across the right ventricular outflow tract.

Received 14 January 2005 Revised 10 May 2005 Accepted 27 June 2005 DOI: 10.1259/bjr/47408749 ’ 2006 The British Institute of Radiology

An 8-year-old boy presented with dyspnoea on exertion and mild cyanosis. Physical examination revealed mild central cyanosis with oxygen saturation of 80%, clubbing of fingers, single S2 with increased intensity and grade 3/6 systolic ejection murmur at left upper sternal border. Electrocardiography (ECG) demonstrated right axis deviation and right ventricular hypertrophy. ECG showed a large subaortic ventricular septal defect (VSD) with an overriding aorta (,50%), mitral-aortic valve continuity, right ventricular hypertrophy, severe infundibular and valvular pulmonary stenosis (PS), and the aorta is parallel to and left of the pulmonary artery. The coronary arteries were not well visualized. Cardiac catheterization confirmed the anatomical findings detected by echocardiography and, in addition, demonstrated a single coronary artery arising from the right coronary cusp, which supplied both right and left coronary arteries. The diagnosis of tetralogy of Fallot with anatomically corrected malposition of the great arteries and single coronary artery was therefore made. To better delineate the course of the coronary arteries, CT angiography (CTA) of the coronary arteries was performed with a 4-slice multidetector CT (MDCT) (Light Speed Plus; GE, Milwaukee, WI) in the retrospective ECG gated mode, axial plane using 1.25 mm collimation, 1 mm reconstruction interval at 75% of the R-R interval, covering the level from carina to cardiac apex. Respiratory effort was suspended by the anaesthetist for the 17 s data acquisition. Non-ionic contrast media injection via the peripheral vein was introduced at a dose of 2 ml kg21 and a rate of 2 ml s21. The scan started 20 s after injection of the contrast media. Postprocessing three-dimensional volume rendering and multiplanar reformation images were acquired by GE The British Journal of Radiology, July 2006

Medical System software (Volume Analysis Vox two 3.0.26e). This revealed subvalvular and valvular PS, large subaortic VSD (Figure 1), overriding of the aorta and aortic-mitral valve continuity (Figure 2). The great arteries originated above the appropriate ventricles, although there was some degree of overriding of the aorta over the VSD. Moreover, their spatial relationships were abnormal in that the aorta was anterior to and left of the pulmonary artery (Figures 1 and 3). Figure 3 demonstrated a single coronary artery from the right coronary cusp. The left coronary artery (LCA) courses to the left normally, dividing into the left anterior descending and the left circumflex arteries running along the interventricular groove, and then along the left atrium posteriorly, respectively. The right coronary artery (RCA) courses to the right and normally divides into two branches, one along the right atrioventricular groove and the other along the surface of the right ventricle. However, the course of the RCA, which is normal, traverses the right ventricular outflow tract. In contrast, the abnormal LCA from the right coronary cusp traverses to the left, but not across the right ventricular outflow tract.

Discussion Anatomically corrected malposition of the great arteries (ACMGA) is a rare form of conotruncal malformation in which the aorta and pulmonary artery arise from their appropriate ventricles despite the presence of malposition of the great arteries (S,D,L) [1]. The frequent associated anomalies are valvular and subvalvular pulmonary stenosis. This case report described tetralogy of Fallot associated with ACMGA e5

A Khositseth, R Pornkul and S Siripornpitak

Figure 3. Volume-rendered 3D reconstruction of the heart Figure 1. Coronal multiplanar reconstruction from CT angiography with a 1.25 mm slice thickness and cardiac gating technique demonstrates findings compatible with tetralogy of Fallot [right ventricular hypertrophy, right ventricular outflow tract obstruction (#), ventricular septal defect (*)]. The aorta (AO) is to the left of the pulmonary artery (PA). LV, left ventricle; RV, right ventricle.

Figure 2. Oblique coronal multiplanar reconstruction from CT angiography with a 1.25 mm slice thickness and cardiac gating technique demonstrates the fibrous continuity between the mitral and aortic annuli (bold arrows), a large subaortic ventricular septal defect (*) and overriding of the aorta (AO). AV, aortic valve leaflets; MV, mitral valve leaflets as shown by dash arrow; LV, left ventricle; and RV, right ventricle.

e6

and great vessels from multidetector CT demonstrates the aorta (AO) located anterior and leftward to the pulmonary artery (PA). Single coronary artery comes off the right coronary cusp. RCA, right coronary artery; LCA, left coronary artery; LCX, left circumflex artery; LAD, left anterior descending artery; RA, right atrium; RV, right ventricle; and LV, left ventricle.

and single coronary artery. Although the echocardiography and cardiac catheterization can elucidate almost all important findings, the exact course of the coronary arteries is so crucial in determining the best possible technique of surgical repair to avoid injuring the right coronary artery. In the repair of a tetralogy of Fallot (TOF), a transannular pericardial patch along with patch closure of the VSD is needed. The repair is technically difficult in this patient because the path of the RCA anterior to the pulmonary trunk virtually precludes a transannular patch repair. Pulmonary valvulotomy and/ or infundibulectomy have been performed in the majority of reported patients with ACMGA, but incomplete relief of the obstruction occurred in some patients [2–4], which required a conduit repair from the right ventricle to the pulmonary artery [2]. There was a report that a leftward deviation of the RCA away from the right atrioventricular groove associated with a levoposition of the anterior aorta allowed successful transannular subpulmonary incision of the right ventricular outflow tract along the right atrioventricular groove without jeopardizing the RCA [5]. However, the anatomical orientation of the RCA to the right ventricular outflow tract is still the main concern. In this case, the course of the RCA (which normally arises from the right coronary cusp, traverses across the right ventricular outflow tract) potentially hinders the transannular patch repair technique. Finally, this patient underwent patch closure of the VSD and pericardial valve conduit with valve ring of 20 mm connected between the right ventricle and the main pulmonary artery to avoid any injury to the RCA and its branches. He had uneventful surgery with a discharge echocardiography demonstrating no residual VSD and mild residual pulmonary stenosis with a The British Journal of Radiology, July 2006

Case report: Multidetector CT in tetralogy of Fallot and single coronay artery

pressure gradient of 21 mm Hg and mild pulmonary regurgitation. MDCT is new technology, which can be performed safely even in small infants. It provides clear information pertinent to great vessels and coronary arteries and has been proven to be very useful in delineating the complex cardiac anatomy non-invasively in this patient. On the other hand, magnetic resonance angiography (MRA) of the coronary arteries is another useful technique in the evaluation of coronary arteries in congenital heart disease [6]. The major advantage of MRA, particularly in children, is no exposure to ionizing radiation and the fact that it does not involve the use of contrast media which may cause adverse reaction such as renal impairment. However, MRA is technically more difficult and time consuming than MDCT. More importantly, in comparison with MDCT, MRA offers less spatial resolution. Lastly, the availability of the coronary MRA software is not as widely available as MDCT. Therefore, MDCT is a more feasible and applicable tool of investigation in most circumstances. In conclusion, MDCT may be able to replace cardiac catheterization in the assessment of coronary artery anatomy. As illustrated in this patient, MDCT, in combination with echocardiography, provides an excellent, non-invasive method of obtaining detailed images


required for planning in the management of this complex cardiac anatomy.

References 1. Van Praagh R, Durnin RE, Jockin H, Wagner HR, Korns M, Garabedian H, et al. Anatomically corrected malposition of the great arteries (S, D, L). Circulation 1975;51:20–31. 2. Rittenhouse EA, Tenckhoff L, Kawabori I, Mansfield PB, Hall DG, Brown JW, et al. Surgical repair of anatomically corrected malposition of the great arteries. Ann Thorac Surg 1986;42:220–8. 3. Kirklin JW, Pacifico AD, Bargeron LM Jr, Soto B. Cardiac repair in anatomically corrected malposition of the great arteries. Circulation 1973;48:153–9. 4. Tsuchida K, Fujiwara T, Ishihara S, Kurosawa H, Imai Y. [Intracardiac repair in anatomically corrected malposition of the great arteries (SDL): report on 2 successful repair]. Nippon Kyobu Geka Gakkai Zasshi 1989;37:760–5. 5. Morita K, Kurosawa H, Koyanagi K, Nomura K, Uno Y, Naganuma H, et al. Atrioventricular groove patch plasty for anatomically corrected malposition of the great arteries. J Thorac Cardiovasc Surg 2001;122:872–8. 6. Taylor AM, Thorne SA, Rubens MB, Jhooti P, Keegan J, Gatehouse PD, et al. Coronary artery imaging in grown up congenital heart disease: complementary role of magnetic resonance and x-ray coronary angiography. Circulation 2000;101:1670–8.

e7


CASE REPORT

Complete eversion and prolapse of bladder concurrent with primary adenocarcinoma 1

Y H KIM, MD, 1D J SUNG, MD, PhD, 1S B CHO, MD, 1K B CHUNG, MD, 1S H CHA, MD, 2H S PARK, MD and J W UM, MD

3

Departments of 1Radiology, 2Urology and 3Surgery, Korea University College of Medicine, Anam Hospital, Korea University, College of Medicine, #126-1, 5-Ka Anam-dong, Sungbuk ku, Seoul 136705, Korea ABSTRACT. Complete eversion and prolapse of the urinary bladder is extremely rare. To the best of our knowledge, the imaging findings of complete bladder eversion have not been documented in the literature. Here, we report a case of complete eversion and prolapse of the urinary bladder demonstrated on MRI. Concurrent primary adenocarcinoma was found in the thickened wall of the everted urinary bladder.

Received 14 January 2005 Revised 23 April 2005 Accepted 27 June 2005 DOI: 10.1259/bjr/56199075 ’ 2006 The British Institute of Radiology

Case report A 78-year-old female, gravida 10, para 8, presented with a 10-year history of a recurrent protruding mass, which was reduced by herself, in the genital area. On admission to the hospital, she complained of an irreducible mass in the genital area for 1 month. She had a long history of stress urinary incontinence and had suffered from total incontinence during the preceding 2 years. She did not sustain a pelvic injury induced by a trauma or a complicated delivery. The physical examination revealed an erythematous mass, which was covered with whitish plaque at the apex of the mass, in the external genital area and a small orifice below the mass. Usual urethral orifice and vaginal introitus were not found in the external genital area. Before undergoing MRI, sterile saline was infused into the small orifice below the mass through a catheter to demonstrate either the urinary bladder or the vagina and a possible fistula. MRI revealed a protruding mass in the external genital area without any evidence of a detectable bladder in the pelvis. The mass showed heterogeneous signal intensity and stretched both ureters inferiorly on T2 weighted image (Figure 1a,b). The periphery of the mass was enhanced diffusely after administering the gadolinium contrast (Figure 1c). The uterus, distended by saline, descended so that it lay posterior to the pubis, but the cervix did not prolapse through the vagina (Figure 1a). Careful physical examination was performed again and two orifices were found bilaterally at the base of the mass (Figure 2). 5 Fr ureteral catheters were introduced into the orifices and the subsequent tubography showed stretched Address correspondence to: Deuk Jae Sung, Department of Radiology, Anam Hospital, Korea University, College of Medicine, #126-1, 5-Ka Anam-dong, Sungbuk ku, Seoul 136-705, Korea.

e8

ureters below the symphysis pubis with bilateral hydronephroureterosis, which was demonstrated on the MR images (Figure 3). The genital mass was confirmed as an everted and prolapsed urinary bladder with the ureteral orifices being exposed externally. The periphery of the mass, which showed diffuse contrast enhancement, was the thickened wall of the everted bladder. Even though the MRI did not demonstrate a focal mass lesion in the everted bladder, multiple punch biopsies were performed to rule out a potential malignancy because the everted bladder wall, which corresponded to the periphery of the mass, was diffusely thickened with contrast enhancement on the MR images. Histological examination revealed adenocarcinoma with moderate differentiation. Radical cystectomy and urinary diversion were performed. Gross examination showed a diffusely thickened and everted bladder wall peripherally with omental fat centrally. The tumour revealed a diffuse infiltrative pattern in the thickened wall (Figure 4). Further pathological evaluation confirmed the diagnosis of mucinous adenocarcinoma in the everted bladder.

Discussion A MEDLINE search from 1966 to December 1999 revealed only nine prior cases of bladder eversion and all cases were reported in female patients. The diagnosis of prolapse in female patients usually refers to genital prolapse, i.e. uterine or vaginal. The lower urinary tract organs, including the bladder, urethra and ureter can also prolapse. Complete eversion and prolapse of the urinary bladder must be differentiated from urethral prolapse, prolapse of an ureterocele, prolapse of redundant bladder mucosa, or a polypoid tumour of the urethra [1]. The British Journal of Radiology, July 2006

Case report: Complete eversion and prolapse of the bladder

(a)

(b)

(c) Figure 1. (a) Sagittal T2 weighted MR image of the pelvis obtained with the patient at rest shows a protruding mass with heterogeneous signal intensity, which was confirmed as the everted and prolapsed bladder in the external genital area. The omental fat (black arrows) is herniated into the centre of the mass and the everted bladder (white arrows) is seen inferiorly with heterogeneous signal intensity at the periphery of the mass. A small amount of fluid (curved arrow) is seen between the everted bladder and the herniated omental fat. The uterus distended by saline (w) descends and lies posterior to the pubis. The cervix does not prolapse through the vagina (arrowheads). (b) Coronal T2 weighted MR image of the pelvis obtained with the patient at rest shows the everted and prolapsed bladder between the thighs. Both ureters (arrows) are dilated and stretched below the symphysis pubis. The uterus distended by saline (w) has a lower position in the pelvis. (c) Gadolinium-enhanced T1 weighted image shows diffuse thickening of the everted bladder wall (arrows) with contrast enhancement at the periphery of the protruding mass between the thighs. Herniated omental fat of heterogeneously high signal intensity (curved arrow) is seen in the centre of the mass without contrast enhancement.

The pathophysiology of bladder eversion remains speculative because of the extreme paucity of information. Half of the previously reported cases of bladder The British Journal of Radiology, July 2006

eversion had concurrent uterovaginal prolapse, which might predispose the patient to bladder eversion by two mechanisms: (1) widening of the urogenital hiatus can e9

Y H Kim, D J Sung, S B Cho et al

Figure 2. Colour photograph shows an erythematous mass covered with whitish plaques in the external genital area. Two orifices (arrows), which were confirmed as ureteral orifices, are exposed externally at the base of the mass with 5 Fr ureteral catheters.

result in pulling the bladder base and posterior urethra away from the pubic bone, which can open the proximal urethra and allow the bladder to begin inverting;

Figure 3. Tubography through the ureteral catheters shows stretched ureters below the symphysis pubis with bilateral hydronephroureterosis. e10

Figure 4. Colour photograph of the gross surgical specimen shows a large mass partially covered by everted and thickened bladder wall (arrows). The central portion of the mass is herniated omental fat (w). The tumour (curved arrows) reveals a diffuse infiltrative pattern in the thickened wall.

(2) vaginal prolapse can cause obstructed voiding with straining, which can also begin the process of the bladder inverting through the urethra [2]. However, our case did not show the concurrent uterovaginal prolapse. The inordinate laxity of the bladder outlet and urethra is assumed to be the other underlying mechanism [3]. Complete eversion and prolapse of the urinary bladder can cause bilateral hydronephrosis and resultant azotemia. Possible mechanisms for hydroenphrosis in patients with a prolapsed bladder include bladder outlet obstruction, ureteral kinking and stretching by the prolapsed bladder [4]. Our case presented with bilateral hydronephrosis that resulted from the stretched ureters. In the case of a prolapsed bladder, the posterior wall of the bladder descends along an arc, initially moving posteriorly and inferiorly to deform the anterior wall of the vagina and then bulging forward as it exits the introitus, at MRI [5]. However, our case presented with a protruding mass at the vulva without any evidence of a detectable bladder in the pelvis at MRI and excluded the possibility of a cystocele. MRI can show a bulging mass outside the external genitalia in complete eversion of the vaginal vault. The presence or absence of uterovaginal prolapse should be noted in the complete eversion and prolapse of the bladder. Prior reports of transurethral bladder eversion described repair by either an abdominal or vaginal surgical approach. Vaginal obliteration or reconstruction combined with a sling urethropexy can be sufficient to address primary pelvic floor defects in patients with transurethral bladder eversion [2]. Pelvic visceral prolapse frequently involves multiple sites, which is suggestive of global pelvic floor weakness. MRI can clearly show pelvic visceral prolapse and be a useful pre-operative planning The British Journal of Radiology, July 2006

Case report: Complete eversion and prolapse of the bladder

tool in women with multicompartment involvement of pelvic floor weakness for whom a complex repair is planned or who have undergone previous repairs. Adenocarcinoma is the third most frequent malignancy of the bladder and has been reported to represent 0.5–2% of all bladder cancers. Primary non-urachal adenocarcinomas of the bladder usually develop in patients with diffuse intestinalization of the bladder mucosa associated with obstruction, cystocele, nonfunctioning bladder, or chronic irritation. They are believed to develop from metaplasia of the bladder epithelium and are associated with cystitis glandularis [6]. The concurrent mucinous adenocarcinoma in our case was assumed to originate from the chronic irritation on the everted bladder wall. The mucosa can become thick if the vaginal vault prolapse is long-standing [7]. As our case demonstrated, the wall of the chronically everted and prolapsed bladder was thick. A diffusely thickened bladder wall with stranding in the surrounding fat is a frequent imaging finding of primary bladder adenocarcinoma [8]. Our case showed diffuse thickening of the everted bladder wall with heterogeneous contrast enhancement on contrast-enhanced MR images. Even though a variety of inflammatory, infective and fibrotic conditions can also result in diffuse bladder wall thickening, it is important for a radiologist to consider adenocarcinoma in the differential diagnosis [8]. In conclusion, the differential diagnosis must include complete eversion and prolapse of the urinary bladder when the diagnostic imaging reveals a protruding mass


at the vulva without evidence of a detectable bladder in the pelvis. The bladder eversion can only develop in female patients with or without genital prolapse. MRI is useful in the assessment of the urogenital prolapse and its associated abnormalities.

References 1. Underhill TR, Altaffer LF 3rd. Complete eversion and prolapse of bladder. Urology 1983;22:622–3. 2. Mastropietro MA, Clark MH, Hale DS. Transurethral bladder eversion concurrent with uterovaginal prolapse. Obstet Gynecol 2002;99:921–5. 3. Chan YT, Ng WD, Yin TF, Kwok TF. Complete prolapse of the urinary bladder. Br J Urol 1990;66:436–7. 4. McGuire EJ. Urinary incontinece. In: Pollack HM. McClennan BL, editors. Clinical urography. 2nd edn. Philadelphia, PA: WB Saunders, 2000:2409–53. 5. Pannu HK, Kaufman HS, Cundiff GW, Genadry R, Bluemke DA, Fishman EK. Dynamic MR imaging of pelvic organ prolapse: spectrum of abnormalities. Radiographics 2000;20:1567–82. 6. Grignon DJ. Neoplasms of the urinary bladder. In: Bostwick DG, Eble JN, editors. Urologic surgical pathology. St. Louis, MO: Mosby, 1997:257–65. 7. Timmons MC, Addison WA. Vaginal vault prolapse. In: Brubaker LT, Saclarides TJ, editors. The female pelvic floor: disorders of function and support. Philadelphia, PA: Davis, 1996:262–8. 8. Hughes MJ, Fisher C, Sohaib SA. Imaging features of primary nonurachal adenocarcinoma of the bladder. AJR Am J Roentgenol 2004;183:1397–401.

e11


CASE REPORT

Unilateral moyamoya disease with co-existing arteriovenous malformation 1

O NAWAWI,

MBBS, MRad, FRCR,

2

M SINNASAMY,

MBBS, MRad

and 1N RAMLI,

MBBS, FRCR

1

Department of Radiology, University Malaya Medical Centre, Kuala Lumpur 59100 and Department of Medicine, Unit of Radiology, Universiti Putra Malaysia, Kuala Lumpur, Malaysia

2

ABSTRACT. A case of an intracerebral bleed in a young man with a rare combination of arteriovenous malformation (AVM) and unilateral moyamoya disease is presented. The location of the bleed in the left basal ganglia corresponded to the area supplied by the basal moyamoya vessels. The AVM which received supply from collateral moyamoya vessels as well as normal cerebral arteries was located in the ipsilateral parieto-occipital region posterior to the basal ganglia bleed. This is the first reported cerebral AVM coexisting with a unilateral moyamoya disease in the English literature. Unusual features of the case such as the unilaterality of the angiographic abnormalities, their coexistence and hypotheses as to their development are discussed.

The term moyamoya is a Japanese word which means ‘‘a puff of smoke’’. It was coined by Suzuki and Takaku in 1969 to describe the angiographic appearance of the fine, telangiectatic collateral vessels at the base of the brain in a group of 21 patients with internal carotid arterial occlusion [1]. The disease was typically seen involving both internal carotid arteries at their clinoid portion. Moyamoya disease has been associated with other vascular diseases such as aneurysm and AVM. The aneurysms which may either be true saccular aneurysms or microaneurysms arise in the telangiectatic net of collaterals, and are a frequent cause of haemorrhage in moyamoya disease [2]. The coexistence of moyamoya disease and AVM, however, is thought to be very unusual. Currently there are 14 reported cases in the literature, 13 of which have bilateral moyamoya disease involvement. There is one case of unilateral moyamoya disease associated with a dural AVM reported in the Japanese literature [3–14]. The association between these two vascular abnormalities is still unexplained. Previous authors have offered several hypotheses, but the possibility of chance still cannot be excluded.

Case report A previously healthy 35-year-old gentleman presented to the emergency department with a sudden episode of generalized tonic clonic seizure 2 h earlier. The seizure, which lasted for 15 min, occurred while he was having a bath. Another similar episode was observed in the emergency department. There was no history of previous tuberculosis or substance abuse. The patient was not on any medication. On clinical examination, the patient was noted to be drowsy, aphasic and had right hemiplegia. He was e12

Received 23 May 2005 Accepted 27 June 2005 DOI: 10.1259/bjr/40183814 ’ 2006 The British Institute of Radiology

afebrile, and clinical examination of the cardiovascular/ respiratory systems and abdomen were normal. Laboratory studies, which included electrolytes, coagulation profile and urinalysis, were normal. An urgent non-contrasted CT examination of the brain revealed a left basal ganglia bleed measuring 4 cm62.5 cm in diameter. Also seen was a focus of calcification in the left posterior parietal area with no associated mass effect (Figure 1). Cerebral angiogram performed the following day showed stenosis of the supraclinoid left internal carotid artery with occlusion of the proximal left middle cerebral artery. There was also narrowing of the A1 segment of the left anterior cerebral artery. Multiple tiny collaterals were seen arising in the area of the lenticulostriate branches giving a ‘‘puff of smoke’’ appearance (Figure 2). The location of these collateral vessels corresponded to the region of basal ganglia bleed seen on the CT. More peripherally in the left parietal region, an arteriovenous malformation (AVM) was seen with the nidus measuring 3 cm in diameter. The feeding arteries were from the branches of the left anterior cerebral artery, collaterals of the occluded left-middle cerebral artery and branches of the left superior cerebellar artery. This vascular malformation appeared early in the arterial phase of the angiogram. There was retrograde filling of the distal cortical left middle cerebral artery from the AVM nidus. The AVM drained into the left sigmoid sinus with no deep venous drainage seen. The right sided cerebral angiogram was normal. The patient was managed conservatively with antiepileptic medication (Dilantin). In the ward, he gradually regained full consciousness and speech, but with residual right hemiparesis for which physiotherapy was instituted. Unfortunately, the outcome of this patient is unknown as he decided to return to his country of origin for further treatment. The British Journal of Radiology, July 2006

Case report: Unilateral moyamoya disease with co-existing AVM

Figure 1. Non enhanced CT of the brain showing a left basal ganglia haematoma with associated mass effect. A focus of calcification is seen in the left posterior parietal region.

Discussion Moyamoya is an occlusive cerebrovascular disease characterized by cerebral angiographic features consisting of intracranial stenosis or occlusion of the internal carotid artery or its terminal branches associated with telangiectatic collateral vessels at the base of the brain [1]. In addition to the enlarged lenticulostriate and thalamoperforating arteries, multiple leptomeningeal and transdural anastomoses develop between the cortical arteries and those of dura mater, scalp, and orbit [2]. The aetiology of this vascular abnormality remains unclear. Opinion is still divided between a congenital versus acquired aetiology. It is usually a bilateral disease, however, atypical unilateral involvement has been reported [15–17]. Although a number of factors have been suggested as the cause of unilateral diseases, in the majority of cases the cause is still inapparent. The onset of symptoms in moyamoya disease has been known to occur in two peaks, the first decade and the fourth decade. Recurrent episodes of sudden hemiplegia, headache and convulsive manifestations occur more frequently in paediatric patients, while subarachnoid or intracerebral bleeding is usually the presenting finding in adults [18]. After a progressive course for many years, the disease frequently stabilizes with residual disability. Cerebral AVM is a congenital abnormality consisting of a nidus of abnormal dilated tortuous arteries and veins with a tangle of closely packed pathologic vessels, resulting in shunting of blood from arterial to venous side without intermediary capillary bed. It is known to be accompanied by a variety of dysplasias involving the feeding arteries and draining veins. Occlusions, stenoses, aneurysms and ectasias are the dysplastic changes affecting the feeding arteries and draining veins. Flow related arterial aneurysms, as well as venous occlusions The British Journal of Radiology, July 2006

or stenoses, have been shown to have an increased risk of haemorrhage [10, 19, 20]. The combined presence of AVM and moyamoya disease is extremely rare in contrast to the presence of saccular aneurysms in patients with this occlusive disease. To date, there are only 14 cases of moyamoya disease accompanied by AVM reported in the literature. 10 of the patients presented with symptoms of cerebral ischaemia, while the remaining four presented with basal ganglia bleed. In all but one case, cerebral angiography demonstrated bilateral involvement of the internal carotid arteries with moyamoya disease. One patient had two independent AVMs located in the frontal lobes, which were supplied mainly by moyamoya collateral vessels through transdural communicating arteries. 13 out of the 15 AVMs were supplied by moyamoya collateral vessels [3–14]. So far, the pathogenesis and relationship between moyamoya disease and AVM are still obscure. There has been no satisfactory explanation as to why this is such a rare phenomenon. The question to be answered is whether the moyamoya disease caused or was caused by the AVM. It is also possible that two independent vascular abnormalities occur in one patient as two AVMs have been reported not to be supplied by moyamoya collateral vessels but by normal cerebral arteries. In these cases the presence of AVMs and moyamoya disease could purely be coincidental [9, 11]. Some authors believe that progressive vascular occlusion might be induced by AVMs. They observed new or progressive occlusions and stenoses of the arteries that originally supplied the AVMs. They suggested that the stress of high and increased blood flow through an AVM induces intimal hyperplasia at the terminal internal carotid artery, eventually leading to stenosis and occlusion. This process is also thought to represent a built-in protective mechanism against cerebral haemorrhage in AVMs [9, 10]. Besides the haemodynamic stress, the authors also suggested that angiogenesis is responsible for the proliferation of new capillaries observed in the moyamoya phenomenon. The angiogenesis might be induced by the release of an angiogenic factor such as transforming growth factors associated with AVMs [10]. These factors may originate from in situ sequestration of platelets and fibrin within the AVM nidus or its feeding arteries [21]. On the other hand, few authors have implicated angiogenesis process in moyamoya disease as being the cause for AVMs. Lichtor et al [2] postulated that in moyamoya disease the perforating vessels and end capillaries become distended in response to ischaemia. However, the capillary linkage is not rich enough to reach the cortex, so instead the increased blood flow is channelled into the normal draining veins. As a result these veins become dilated, taking on the appearance of an AVM. Unlike a true AVM, these vessels filled late rather than early in the injection phase [2]. Another report supporting AVM as being secondary to moyamoya was by Schmit et al, who demonstrated the development of AVM in a patient with moyamoya by repeated angiography over a 9-year period [12]. These reports suggest that such AVMs were not congenital, but developed as a consequence of angiogenic failure [3]. e13

O Nawawi, M Sinnasamy and N Ramli

Figure 2. Left internal carotid angiogram in Towne’s (upper) and lateral (lower) projection, showing occluded left middle cerebral with associated small collateral lenticulostriate vessels (arrow). Posterolaterally an arteriovenous malformation (AVM) is seen, partly supplied by the moyamoya collateral vessels and draining into the left sigmoid sinus.

Our case is unusual not only due to the coexistence of moyamoya disease and AVM, but also because of the atypical unilateral involvement of the moyamoya disease. There have been few reports on unilateral involvement of moyamoya disease in the literature in both adults and children. The progression to the typical bilateral form is infrequent. However, young children have a higher tendency to develop contralateral lesions over the years [22, 23]. To the best of our knowledge there has been no report of combined presence of AVM and unilateral moyamoya disease in the English literature. Suggested causes for atypical unilateral involvement of moyamoya disease are tuberculous meningitis, irradiation or artherosclerosis [15, 17], none of which were present in our patient. In most cases, the cause remains undetermined. We cannot do more than speculate whether the presence of AVM initiated the development of unilateral moyamoya disease in our case. The intracerebral bleed in our patient was not from the AVM, but from the collateral vessels at the lenticulostriate region. This is not surprising as pathologic e14

study of moyamoya collateral vessels has shown various degrees of luminal stenosis, intimal thickening, discontinuity of the internal elastic lamina and microaneurysm formation, which cause the vessels to be more fragile and susceptible to bleeding [24]. Furthermore, Mawad et al suggested that the arterial stenosis proximal to the AVM serves as a built-in mechanism against cerebral haemorrhage from the AVM by reducing the flow to the nidus [9]. The treatment of moyamoya disease involves establishing as much blood flow as possible to the brain by creating direct or indirect bypass. On the other hand, AVMs are treated by stopping the flow into the nidus. Hence, the management principles in treating the combined presence of AVM and moyamoya disease are in direct conflict. Radiosurgery (gamma knife) has been successfully used to treat cerebral AVM in moyamoya disease [4]. In our case, the treatment was primarily symptomatic and supportive as the patient intended to seek treatment elsewhere. In conclusion, we present a rare case of combined presence of moyamoya disease and cerebral AVM in a The British Journal of Radiology, July 2006

Case report: Unilateral moyamoya disease with co-existing AVM

young and otherwise healthy gentleman. Our case is also unusual as the moyamoya disease was only present on the side where the AVM was found.

12.

References 1. Suzuki J, Takaku A. Cerebrovascular ‘‘moyamoya’’ disease: Disease showing abnormal net-like vessels in the base of brain. Arch Neurol 1969;20:288–99. 2. Lichtor T, Mullan S. Arteriovenous malformation in moyamoya syndrome. J Neurosurg 1987;67:603–8. 3. Nakashima T, Nakayama N, Furuichi M, Kokuzawa J, Murukawa T, Sakai N. Arteriovenous malformation in association with moyamoya disease: report of two cases. Neurosurg Focus 1998;5 (5):Article 6. 4. Seol HJ, Kim DG, Oh CW, Han DH. Radiosurgical treatment of a cerebral arteriovenous malformation in a patient with moyamoya disease: case report. Neurosurgery 2002;51:478–82. 5. Akiyama K, Minakawa T, Tsuji Y, Isayama K. Arterivenous malformation associated with moyamoya disease: case report. Surg Neurol 1994;41:468–71. 6. Fuse T, Takagi T, Fukushima T, Hashimoto N, Yamada K. Arteriovenous malformation associated with moyamoya disease. Childs Nerv Syst 1996;12:404–8. 7. Halatsch ME, Rustenbeck HH, Jansen J. Progression of arteriovenous malformation in moyamoya syndrome. Acta Neurochir (Wien) 1997;139:82–5. 8. Kayama T, Suzuki S, Sakurai Y, Nagayama T, Ogawa A, Yoshimoto T. A case of moyamoya disease accompanied by an arteriovenous malformation. Neurosurgery 1986;18: 465–8. 9. Mawad ME, Hilal SK, Michelsen WJ, Stein B, Ganti SR. Occlusive vascular disease associated with cerebral arteriovenous malformations. Radiology 1984;153:401–8. 10. Montanera W, Marotta TR, ter Brugge K, Lasjaunias P, Willinsky R, Wallace MC. Cerebral arteriovenous malformations associated with moyamoya phenomenon. Am J Neuroradiol 1990;11:1153–6. 11. Okada T, Kida Y, Kinomoto T, Sakurai T, Kobayashi T. Arteriovenous malformation associated with moyamoya


13. 14.

15.

16.

17.

18. 19. 20.

21. 22.

23.

24.

disease: case report. Neurol Med Chir (Tokyo) 1990;30:945–8. Schmit BP, Burrows PE, Kuban K, Goumnerova L, Scott RM. Acquired cerebral arteriovenous malformation in a child with moyamoya disease: case report. J Neurosurg 1995;84:677–80. Scott RM. Arteriovenous malformation and moyamoya disease. Childs Nerv Syst 1997;13:357. Yamasaki F, Hotta T, Taniguchi E, Eguchi K, Hashizume A, Kodama Y, et al. A case of dural arteriovenous malformation in the anterior fossa associated with an occlusion of the unilateral middle cerebral artery with moyamoya phenomenon [in Japanese]. No Shinkei Geka 1995;23:603–7. Debrun G, Sauvegrain J, Aircardi J. Moyamoya, a nonspecific radiological syndrome. Neuroradiology 1975;8:241–4. Hilal SK, Soloman GE, Gold AP, Carter S. Primary cerebral arterial occlusive disease in children. Part II. Neurocutaneous syndromes. Radiology 1971;99:87–93. Mathew NT, Abraham J, Chandy J. Cerebral angiographic features in tuberculosis meningitis. Neurology 1970;10:1015–23. Hardy RC, Williams RG. Moyamoya disease and cerebral hemorrhage. Surg Neurol 1984;21:507–10. Yasargil MG. Microneurosurgery Vol. IIIa. Stuttgart, Germany: Thieme, 1992:138–211. Willinsky R, Lasjaunias P, terBrugge K, Pruvost P. Brain arteriovenous malformations (BAVMS): analysis of the angioarchitecture in relationship to hemorrhage. J Neuroradiol 1988;15:225–37. Folkman J, Klagsbrun M. Angiogenic factors. Science 1987;235:442–7. Hirotsune N, Meguro T, Kawada S. Long-term follow-up study of patients with unilateral moyamoya disease. Clin Neurol Neurosurg 1997;99:S178–81. Houkin K, Abe H, Yoshimoto T, Takahashi A. Is ‘‘unilateral’’ moyamoya disease different from moyamoya disease? J Neurosurg 1996;85:772–6. Yamashita M, Oka K, Tanaka K. Histopathology of the brain vascular network in moyamoya disease. Stroke 1983;14:50–8.

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CASE REPORT

Schwannoma of the nasal septum 1

S RAJAGOPAL, MBBS, 1V KAUSHIK, 1 R K BHATNAGAR, FRCS

MRCS, DLO,

2

K IRION,

MD, PhD,

3

M E HERD,

FRCPath

and

1

Department of Otolaryngology, Head & Neck Surgery, 2Department of Radiology and 3 Department of Pathology, Fairfield General Hospital, Bury, Lancashire BL9 7TA, UK ABSTRACT. Schwannomas of the nasal cavity and paranasal sinuses are very rare. We report the case of a 54-year-old woman with a schwannoma arising from the nasal septum. We discuss the clinical presentation, differential diagnosis, imaging characteristics and treatment of this rarely encountered lesion.


Schwannoma (neurilemmoma) is a benign neoplasm arising from Schwann cells in the peripheral nerve sheath. It can arise throughout the body, but is most commonly observed in the head and neck. This region accounts for 25–45% of all schwannoma [1]. Tumours arising from nasal septum are extremely rare with only 16 cases having been reported in the English literature [1–8].

Case report A 54-year-old woman was referred to our department with a 6-month history of progressive left-sided nasal obstruction. She denied rhinorrhoea, epistaxis, anosmia, facial pain, headache and recent nasal trauma. Her medical and family histories were unremarkable. Anterior rhinoscopy revealed a large polypoid mass completely filling the left nasal cavity. The polyp was firm in consistency and appeared to be covered by normal nasal mucosa. It did not bleed on touch. The attachment of the polyp was difficult to determine. The nasal septum was central and right nasal cavity was clear. A CT scan of the paranasal sinuses was requested to determine the nature and extent of the polyp. This showed a homogeneous mass, lying within the midportion of the left nasal cavity, probably arising from the nasal septum (Figure 1). The lesion was well defined with smooth margins and without calcification (Figure 2). The paranasal sinuses were clear. A contrast study was not performed due to the absence of bony destruction. The patient agreed and consented to undergo removal of the polyp under general anaesthesia. The polypoid mass was attached to middle part of the left side of the nasal septum, opposite the anterior end of the middle turbinate. The mass was excised completely along with a cuff of septal mucoperichondrium. The histopathological e16

diagnosis was benign schwannoma (Figures 3 and 4). The patient continues to be followed-up and remains free of recurrence at 6 months.

Discussion Schwannoma is a benign tumour of nerve sheath origin that can arise on any myelinated nerve. The most frequent site affected in the head and neck is the eighth cranial nerve (vestibular nerve). Other observed locations include the scalp, face, oral cavity, pharynx, larynx, trachea, parotid gland, middle ear and external auditory meatus. Whilst schwannomas almost always occur as solitary lesions with no associated genetic syndromes, in some instances they are multiple and occur in association with neurofibromatosis type 2 [8]. Malignant transformation in schwannoma is very rare. Schwannomas of the sinonasal tract are very infrequent, representing less than 4% of the schwannomas of the head and neck [6]. In this location they have been reported in patients between the ages of 6 years and 78 years. There is no sex or racial predilection [6]. The ethmoidal sinus is most commonly involved, followed by the maxillary sinus, nasal fossa and sphenoid sinus [6, 9, 10]. Localization to the nasal septum is exceedingly rare. Septal schwannomas arise from the autonomic or sensory nerves within the nasal septum. There is no apparent site predilection on the septum. The clinical presentation of sinonasal schwannoma is often varied and non-specific. Patients may complain of nasal obstruction, epistaxis, rhinorrhoea, anosmia, or facial swelling and pain [3]. There are no distinctive features to be noted on examination. Consequently, the diagnosis is only likely to be made once histology results are available. The differential diagnosis of a nasal tumour includes a wide variety of pathology including inflammatory polyps, juvenile angiofibroma, inverted papilloma, The British Journal of Radiology, July 2006

Case report: Septal schwannoma

Figure 3. Microscopic appearance of tumour showing Antoni A pattern with nuclear pallisading (H & E staining).

Figure 1. Coronal CT scan of paranasal sinus showing mass in the left nasal cavity.

meningioma, neurofibroma, melanoma and olfactory neuroblastoma (esthesioneuroblastoma) [1, 11]. Macroscopically, schwannomas appear as gelatinous or cystic, well encapsulated masses. Cystic degeneration, necrosis, lipidization and formation of angiomatous clusters of blood vessels with focal thrombi are degenerative processes that can occur. Microscopically, schwannomas are traditionally classified into two major histological types. Antoni A is characterized by a compact arrangement of spindle cells. Antoni B is typified by loose myxoid stroma with spindle cells running in a haphazard manner. The distinction is considered to have only academic interest [6].

Figure 2. Axial cuts showing homogeneous mass in midportion of left nasal cavity. The British Journal of Radiology, July 2006

Securing the diagnosis on the basis of high-resolution imaging is difficult. In general, the appearances on CT are not specific enough to enable it to be distinguished confidently from other tumours in this region. On CT, paranasal schwannoma usually have mottled central hypodense foci with peripheral enhancement after injection of contrast medium [12]. The heterogeneous appearance is related to areas of increased vascularity with adjacent non-enhancing cystic or necrotic regions. This is important in distinguishing it from inflammatory polyps [13]. Unlike schwannomas originating from the paranasal sinuses, the small confines of the nasal cavity mean that septal schwannomas tend to become symptomatic at an earlier stage and are comparatively smaller in size at presentation. As a result, they are usually excised without the need for radiographic imaging. In cases where CT was performed [4–8] the septal schwannoma appeared as a homogeneous soft tissue mass, with mottled enhancement, occupying the nasal cavity and

Figure 4. Immunohistology slide confirming the diagnosis of schwannoma.

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S Rajagopal, V Kaushik, K Irion et al

sometimes extending into the nasopharynx. The origin from the nasal septum was not always apparent radiologically, and in some cases was only determined at the time of surgery. Trans-septal extension into the opposite nasal cavity was noted on one scan [5]. Large septal schwannomas have caused septal deviation and bowing of both the medial walls of the maxillary sinus and orbit. In many cases CT showed retention of secretions within adjacent sinuses [4–8]. Erosion of the ipsilateral middle and inferior turbinate was seen in one case [6]. To date, there are no reported cases of septal schwannomas extending to the skull base. The imaging characteristics of sinonasal schwannoma on MRI are similar to those of schwannoma observed elsewhere in the body. An intermediate signal is observed on T1 weighted images, whereas on T2 weighted images the signal varies from intermediate to high. A more uniform enhancement pattern after gadolinium administration has been observed [12–14]. MRI has been performed in two of the reported cases of septal schwannoma [5, 7], but the imaging characteristics were not described in detail. Treatment is complete surgical excision, which can range from simple excision under local anaesthesia to a more extensive facial degloving approach [6]. To date, there are no reported cases of recurrence.

References 1. Pasic TR, Makielski K. Nasal schwannoma. Otolaryngol Head Neck Surg 1990;103:943–6. 2. Bogdasarian RM, Stout AP. Neurilemmoma of the nasal septum. Arch Otolaryngol 1943;38:62–4.

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3. Higo R, Yamasoba T, Kikuchi S. Nasal neurinoma: case report and review of literature. Auris Nasus Larynx 1993;20:297–301. 4. Thomas JN. Massive schwannoma arising from the nasal septum. J Laryngol Otol 1977;91:63–8. 5. Oi H, Watanabe Y, Shojaku H, Mizukoshi K. Nasal septal neurinoma. Acta Otolaryngol (Stockh) 1993;Suppl. 504:151–4. 6. Berlucchi M, Piazza C, Blanzuoli L, Battaglia G, Nicolai P. Schwannoma of the nasal septum: a case report with review of the literature. Eur Arch Otorhinolaryngol 2000;257:402–5. 7. Wada A, Matsuda H, Matsuoka K, Kawano T, Furukawa S, Tsukuda M. A case of schwannoma of nasal septum. Auris Nasus Larynx 2001;28:173–5. 8. Butugan O, Schuster GS, de Almeida ER, Miniti A. Schwannoma of the nasal septum: report of two cases. Rev Laryngol Otol Rhino 1993;114:33–6. 9. Batsakis J. Tumours of the head and neck. Clinical and pathological considerations, 2nd edn. Baltimore, MD: Williams and Wilkins, 1979:313–33. 10. Calcaterra TC, Rich R, Ward PW. Neurilemmoma of the sphenoid sinus. Arch Otolaryngol 1974;100:383–5. 11. Kaufman SM, Conrad LP. Schwannoma presenting as nasal polyp. Laryngoscope 1976;69:463–6. 12. Younis RT, Gross CW, Lazar RH. Schwannoma of the paranasal sinus. Arch Otolaryngol Head Neck Surg 1991;117:677–80. 13. Fujiyoshi F, Kajiya Y, Nakajo M. CT and MR imaging of nasoethmoid schwannoma with intracranial extension. AJR Am J Roentgenol 1997;169:1754–5. 14. Dublin AB, Dedo HH, Bridger WH. Intranasal schwannoma: magnetic resonance and computed tomography appearance. Am J Otolaryngol 1995;16:251–4. 15. Cakmak O, Yavuz H, Yucel T. Nasal and paranasal schwannomas. Eur Arch Otorhinolaryngol 2003;260:195–7.



CASE REPORT

Isolated renal mucormycosis: an unusual cause of acute renal infarction in a boy with aplastic anaemia 1

R SHARMA, MD, 1G SHIVANAND, and 3M C SHARMA, MD

MD,

2

R KUMAR,

MD,

2

S PREM,

MD,

1

H KANDPAL,

MD,

1

C J DAS,

MD

1

Department of Radiodiagnosis, 2Department of Hematology and 3Pathology, All India Institute of Medical Sciences, New Delhi –110029, India ABSTRACT. Mucormycosis is an opportunistic fungal infection occurring most often in immunocompromised patients. The most common clinical presentation of infection with the fungus is rhinocerebral, followed by primary pulmonary, disseminated, gastrointestinal and cutaneous manifestations. Renal involvement in disseminated mucormycosis is well described, however, there are only few case reports describing the isolated renal mucormycosis in literature. We present an unusual case of renal mucormycosis in a patient of aplastic anaemia who presented with right flank pain and, on imaging, was found to have renal infarction and thrombosis of renal vessels. Histopathology revealed fungal hyphae infiltrating the entire kidney, renal vessels and perinephric space.

Mucormycosis is a rare opportunistic infection caused by fungi of the order Mucorales and the genera Rhizopus, Absidia and Mucor. Renal involvement in disseminated mucormycosis ranges from 14% to 19% of patients [1]. Isolated renal mucormycosis is extremely rare. Only few case reports have been reported in English literature [2– 9]. To the best of our knowledge, this is the first report of renal mucormycosis in a patient with aplastic anaemia presenting with renal infarction.

Case report A 14-year-old boy with aplastic anaemia presented to the emergency department with a 2-day history of right flank pain. He had no fever, chills, haematuria or dysuria. The patient was on treatment with antithymocyte globulin (ATG) for the aplastic anaemia. Physical examination showed a normal body temperature and right flank tenderness. A mass was palpable in the lumbar region. Laboratory examination revealed low white blood cell, neutrophil and platelet counts. Blood urea and serum creatinine were raised. Chest radiograph revealed no abnormalities. Blood and urine cultures showed no growth of either bacteria or fungus. Ultrasound combined with colour and power Doppler study (Figure 1) revealed enlarged hypoechoic right kidney with no intrarenal vascularity. The right renal vein was enlarged and filled with thrombus, but the inferior vena cava was patent. The right main renal artery could not be visualized, suggestive of either absent flow due to long standing renal vein thrombosis or coexisting thrombosis of renal artery as well. Contrast enhanced CT scan (Figure 2) revealed enlarged non-enhancing right kidney with perinephric The British Journal of Radiology, July 2006

Received 3 May 2005 Revised 13 June 2005 Accepted 30 June 2005 DOI: 10.1259/bjr/17821080 ’ 2006 The British Institute of Radiology

fluid collection. Renal radionuclide scan revealed absent perfusion of right kidney. Based on clinical presentation and fluid collection in perinephric area, presumptive diagnosis of infective aetiology for the infarction was made. Patient was started on a combination of antifungal (Amphotericin B) and antibacterial agents. He was also given blood and platelet transfusions. Subsequently, the patient underwent nephrectomy. Histopathology of nephrectomy specimen revealed fungal hyphae suggestive of mucormycosis infiltrating the entire right kidney, extending into renal vessels and perinephric space (Figure 3). During post-operative period, the patient developed acute respiratory distress syndrome and progressive hyponatraemia and he expired on the fifth post-operative day.

Discussion Mucormycosis (phycomycosis, zygomycosis) is a ubiquitous, opportunistic infection caused by species in the order Mucorales of the class Zygomycetes. Patients who develop this type of infection are usually immunocompromised. The conditions that predispose to this infection are burns, diabetes mellitus, diabetic ketoacidosis, cancer, steroid therapy, alcoholism, illicit drug abuse, solid organ transplantation, malnutrition, renal failure, surgical procedures, trauma, penetrating wounds to the skin, insect bites, use of needles, broad spectrum antibiotic therapy and seropositive HIV status [10, 11]. In our case, the patient had aplastic anaemia and a low neutrophil count. The four main presentations of mucormycosis described are rhinocerebral, pulmonary, gastrointestinal and disseminated forms [11]. Infection of single organs e19

R Sharma, G Shivanand, R Kumar et al

(a)

(b)

Figure 1. Ultrasound with colour Doppler image shows (a) enlarged hypoechoic right kidney with no intrarenal vascularity (b) patent inferior venacava and no flow in renal vein.

such as bone, heart and kidney also rarely occurs. The organism characteristically invades the walls of adjacent blood vessels, producing thrombosis and infarction, but rarely disseminates through the vessels. Renal involvement occurs as part of disseminated zygomycosis in 22% of cases, but isolated renal disease has also been documented in case reports [12]. The patients may present with fever, flank pain and tenderness, gross haematuria and pyuria. These symptoms would be consistent with a diagnosis of acute pyelonephritis. Occasionally, patients present with renal failure in bilateral disease. In a large series of renal mucormycosis reported by Gupta et al, renal infarction was observed in 12 out of 18 patients [11]. Thrombi occluding lumina of both the main renal artery and vein were observed in seven patients, as in our case. The thrombi contained the fungal hyphae in all these seven cases. Isolated renal mucormyscosis may be unilateral or bilateral. In the series reported by Gupta et al, most (13 of

Figure 2. Axial contrast enhanced CT scan shows absent perfusion of right kidney with perinephric collection.

e20

18 patients) of the cases were bilateral [11], unlike our patient who had unilateral involvement.

Figure 3. Photomicrographs showing coagulative necrosis (infarct) of kidney parenchyma (A 6100) with invasion of blood vessels by non-septate fungal hyphae of mucormycosis (B, PAS 6200). The British Journal of Radiology, July 2006

Case report: Isolated renal mucormycosis

The imaging findings include enlarged kidney with or without perinephric collection on ultrasound. Contrast enhanced CT scan may reveal enlarged kidney, with reduced or absent enhancement. There may be multiple low attenuation areas within the parenchyma suggestive of abscesses [9]. In our case, ultrasound combined with colour Doppler revealed enlarged kidney with absent perfusion. CT confirmed these findings and, in addition, there was perinephric collection. To the best of our knowledge, these imaging findings have hitherto not been described in the radiology literature. Patients usually present with flank pain, fever and pyuria. There are no serological tests to identify mucormycosis infections. Blood and urine cultures are usually sterile. Most of the patients are treated for bacterial pyelonephritis and, not surprisingly, do not respond to antibiotics. Diagnosis requires demonstration of fungi in tissues. Most of the Mucorales will grow in routine laboratory culture media and identification is by its morphologic features, which show broad aseptate hyphae with right-angle branching. A high index of clinical suspicion and knowledge of the varied manifestations is very important in diagnosing this condition. If a seriously ill, immunocompromised patient develops clinical features of acute pyelonephritis and ultrasound shows enlarged kidney with or without perinephric collection, contrast enhanced CT should be performed. In the presence of either renal infarction or multiple abscesses in renal parenchymal, biopsy is indicated to confirm the diagnosis. Treatment of renal mucormycosis involves combined surgical and medical treatment. Treatment consists of extensive debridement of necrotic tissue and administration of Amphotericin B (0.6–1 mg kg21 day21). This treatment is usually not successful [11]. Most of the patients with renal mucormycosis will need nephroureterectomy as the fungal bezoar extends down the ureter and because the hyphal elements have a propensity for vascular invasion, causing infarction of tissue which


does not respond well to simply removing the fungal cast. Overall survival for isolated renal mucormycosis is 65%. Mortality is nearly 100% for those who did not undergo nephrectomy. In conclusion, isolated renal mucormycosis is a fatal and unusual cause of renal infarction which needs a high index of clinical suspicion. A combination of aggressive surgical and medical treatment, if instituted, may improve the outcome.

References 1. Raghavan R, Date A, Bhaktaviziam A. Fungal and nocardial infections of the kidney. Histopathology 1987;11:9–20. 2. Davila RM, Moser SA, Grosso LE. Renal mucormycosis: a case report and review of the literature. J Urol 1991;145:1242–4. 3. Pickles R, Long G, Murugasu R. Isolated renal mucormycosis. Med J Aust 1994;160:514–6. 4. Langston C, Roberts DA, Porter GA, et al. Renal phycomycosis. J Urol 1973;109:941–4. 5. Prout GR Jr, Goddard AR. Renal mucormycosis. N Engl J Med 1960;263:1246. 6. Low AI, Tulloch AGS, England EJ. Phycomycosis of the kidney associated with a transient immune defect and treated with clotrimazole. J Urol 1974;111:732–4. 7. Flood HD, O’Brien AM, Kelly DG. Isolated renal mucormycosis. Postgrad Med J 1985;61:175–6. 8. Santos J, Espigado P, Romero C, et al. Isolated renal mucormycosis in two AIDS patients. Eur J Clin Microbiol Infect Dis 1994;13:430–2. 9. Chugh KS, Sakhuja V, Gupta KL, et al. Renal mucormycosis computerized tomographic findings and their diagnostic significance. Am J Kidney Dis 1993;22:393. 10. Vesa J, Bielsa O, Llado C, et al. Massive renal infarction due to mucormycosis in an AIDS patient. Infection 1992;20:234–6. 11. Gupta KL, Joshi K, Sud K, Kohli S, Jha V, Bishan D, et al. Renal zygomycosis: an under-diagnosed cause of acute renal failure. Nephrol Dial Transplant 1999;14:2720–5. 12. Ingram CW, Sennesh J, Cooper JN, Perfect JR. Disseminated zygomycosis: report of four cases and review. Rev Infect Dis 1989;11:741–54.

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CASE REPORT

Iatrogenic bilateral pneumothorax following unilateral transbronchial lung biopsy 1

S FINDIK,

MD, FCCP,

1

L ERKAN,

MD

and 2R W LIGHT,

MD, FCCP

1

Department of Pulmonary Medicine, Faculty of Medicine, Ondokuz Mayis University, Samsun, Turkey and 2Vanderbilt University, Pulmonary Disease Program, Saint Thomas Hospital, Nashville, TN, USA ABSTRACT. Bilateral pneumothoraces are recognized complications of thoracic procedures in patients who have undergone heart or heart-lung transplantation. Bilateral simultaneous pneumothoraces developing following a unilateral transbronchial lung biopsy in the absence of previous thoracic surgery has not been reported previously.


Bilateral pneumothoraces [1, 2] are well known complications of thoracic procedures in patients who have undergone heart or heart-lung transplantation. To date, the only paper describing the development of bilateral pneumothoraces after unilateral transbronchial biopsies occurred in patients who had undergone heartlung transplantation 2 years previously [2]. We present a patient without a history of any thoracic surgical procedure who developed bilateral pneumothoraces following unilateral transbronchial biopsies.

Case report A 35-year-old woman presented to our clinic with a complaint of progressive dyspnoea for 3 years. She denied cough, sputum, haemoptysis, chest pain, wheezing or other respiratory symptoms. She is a housewife who denies any history of illicit drug use, smoking or alcohol ingestion. She had pneumonia while she was in primary school. Her mother has diabetes mellitus type II. On physical examination, she was a well-nourished, well-developed woman in no distress. Her vital signs were within normal limits. The only abnormality was decreased breath sounds over both lung bases, especially on the right side. The only laboratory abnormalities were a C-reactive protein (CRP) elevated to 8.33 mg l21 (normal 0–5) and an IgA elevated to 3.49 g l21 (normal 0.6–3.3). The pulmonary function tests revealed an obstructive pattern with an increase in lung volumes (FVC 68%, FEV1 58%, TLC 84% and RV 115% predicted). The single breath The work was performed at Ondokuz Mayis University, Samsun, Turkey. Address correspondence to: Richard W Light, MD, Saint Thomas Hospital, 4220 Harding Road, Nashville, TN 37205, USA.

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diffusing capacity (DLco) was 5.6 mmol kPa21 min21 (70% predicted), while the DLco/alveolar volume was 95% of predicted. The chest X-ray revealed bilateral hyperlucent lungs. The thoracic CT scan and high resolution CT (HRCT) scan revealed multiple parenchymal cysts, especially in the middle and lower lung zones. Among the cysts, there were both normal parenchyma and atelectatic areas. Interlobular septal thickening was also seen (Figure 1). The initial impression was probable pulmonary lymphangioleiomyomatosis (LAM). Fibre optic bronchoscopy was normal except for the right middle lobe bronchus, which was compressed inferiorly. Multiple transbronchial lung biopsies were taken from the posterior segment of the right upper lobe and the right middle lobe.

Figure 1. CT scan of chest demonstrating multiple intraparenchymal cysts, especially in the middle and lower lung zones. The British Journal of Radiology, July 2006

Case report: Iatrogenic bilateral pneumothorax

Figure 2. CT scan of chest demonstrating small bilateral pneumothoraces.

The chest radiograph taken after bronchoscopy was suspicious for bilateral pneumothoraces, which were confirmed immediately thereafter with a chest CT scan revealing bilateral pneumothoraces, each occupying less than 10% of the hemithorax (Figure 2). Since the presence of a left pneumothorax in addition to the right pneumothorax on the side of the biopsy was distinctly unusual, we hypothesized that there might be a defect in the mediastinum allowing the passage of air from the right pleural space to the left. Thus, we carefully reexamined the higher cuts of the CT scans and noted that the anterior mediastinal line was much less apparent (Figure 3) than it was on the original CT scan (Figure 1). We concluded that air in the right pleural space, which developed following the transbronchial biopsy, probably passed to the left pleural space through the defect in the anterior junction line. The patient was hospitalized and treated with 7–8 l of oxygen per minute via nasal prongs to facilitate resolution of the bilateral pneumothoraces. She did not complain of dyspnoea or chest pain and did not develop

Figure 3. CT scan of chest demonstrating that the anterior junctional line (white arrow) is much less clear than on the comparable CT scan before the pneumothorax (Figure 1). The British Journal of Radiology, July 2006

cyanosis or hypotension. Serial chest radiographs demonstrated progressive resolution of the pneumothoraces and 3 days post-bronchoscopy the chest radiograph and chest CT scans revealed complete resolution of both pneumothoraces. A sagittal CT reconstruction and transverse images confirmed the defect of the anterior junction line. The pathologist at our hospital interpreted the biopsy specimen as emphysematous lung. The biopsy specimens, clinical information and radiological images were sent to Prof. Cagle and Dr Akpolat at Baylor College of Medicine, Houston, USA. They reported that although the case was consistent with pulmonary LAM clinically and radiologically, the biopsies were not sufficient for the diagnosis of LAM and suggested; (1) actin staining and (2) open lung biopsy. Repeat actin staining of the biopsies was negative. The patient refused open lung biopsy and was lost to follow-up.

Discussion The pleural spaces are completely separated from each other in human beings [3]. Radiographically, anteriorly and superiorly to heart, this separation is manifested by the anterior junction line [4], which represents the visceral and parietal pleura of both lungs with varying amounts of adipose tissue. However, the two pleural spaces are separated by only the two parietal pleura and the adipose tissue. During heart-lung transplantation, and occasionally other cardiothoracic surgeries performed via median sternotomy, the two parietal pleurae may become severed, resulting in a communication between the two pleural spaces. This is sometimes referred to as the ‘‘iatrogenic buffalo chest’’ [5, 6]. Persistent pleural connections after heart or heart-lung transplantation have been reported in 33–40% of patients [3]. There are only three reports of iatrogenic buffalo chest in patients who had undergone non-transplant thoracic surgery [6]. One patient had undergone multiple cardiac surgeries for congenital heart disease and developed bilateral pneumothoraces after a thoracentesis. A second patient had undergone coronary artery bypass graft surgery and developed bilateral pneumothoraces after attempts to catheterize a subclavian vein. A third patient had undergone transthoracic oesophageal resection and had developed bilateral pneumothoraces following the percutaneous needle biopsy of a lung nodule [6]. The occurrence of iatrogenic bilateral pneumothoraces as a consequence of a unilateral transbronchial biopsy in a patient who had not undergone any mediastinal or thoracic surgery has never before been reported. One might hypothesize that the pneumothorax on the left was secondary to the pulmonary LAM rather than the communication between the pleural spaces, since pneumothorax is common with LAM [7]. Simultaneous bilateral pneumothoraces have also been reported in patients with other cystic or bullous diseases [8–11]. We think that the bilateral simultaneous pneumothoraces are secondary to the procedure rather than the underlying disease process, for the following reasons: (a) the pneumothoraces appeared synchronously and disappeared synchronously, (b) the CT scans demonstrated e23

S Findik, L Erkan and R W Light

the communication between the pleural spaces and (c) the pneumothoraces were the same size. Pulmonary LAM is a rare lung disease that affects women of childbearing age [7]. It is characterized pathologically by the proliferation of atypical pulmonary interstitial smooth muscle and by cyst formation. The mechanism of cyst formation and emphysema-like disease is unknown. Smooth muscle proliferation within the airways and/or elastic fibre degradation may be contributing factors. Pulmonary LAM has more in common clinically, radiographically and physiologically with pulmonary emphysema than with other interstitial lung diseases. In general, the diagnosis should be strongly suspected in any young woman who presents with emphysema, recurrent pneumothorax or a chylothorax. HRCT often strongly suggests the diagnosis, and tissue confirmation is not always necessary [12]. The findings of diffuse, homogeneous, small, thin-walled cysts are highly suggestive of the diagnosis in the appropriate clinical context [13]. It is possible that the defect in the septum formed by the two parietal pleura was due to pulmonary LAM. However, we feel that this is unlikely since there are no previous reports of such defects in patients with LAM. Moreover, since pneumothorax is relatively common in LAM, one would expect other cases of simultaneous bilateral pneumothoraces if patients with LAM had defects in mediastinal septum. A recent review of bilateral spontaneous pneumothoraces revealed no cases due to LAM [14].

Conclusion In conclusion, physicians should be aware that a communication between the pleural spaces may be present in patients without a history of mediastinal surgery or median sternotomy.

Acknowledgments The authors thank Drs Cagle and Akpolat from Baylor College of Medicine, Houston, USA, for pathologic

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examination of the biopsies, Drs Oner Dikensoy and Heather Misra and Ms Bonnie Cathey for their careful review of the manuscript.

References 1. Paranjpe DV, Wittich GR, Hamid LW, et al. Frequency and management of pneumothoraces in heart-lung transplant recipients. Radiology 1994;190:255–6. 2. Lee YC, McGrath GB, Chin WS, et al. Contralateral tension pneumothorax following unilateral chest tube drainage of bilateral pneumothoraces in a heart-lung transplant patient. Chest 1999;116:1131–3. 3. Wittich GR, Kusnick CA, Starnes VA, et al. Communication between the two pleural cavities after major cardiothoracic surgery: relevance to percutaneous intervention. Radiology 1992;184:461–2. 4. Fraser RS, Muller NL, Colman N, Pare PD. The mediastinum. In: Diagnosis of diseases of the chest (4th edn). Philadelphia, PA: WB Saunders Company, 1999;196–234. 5. Schorlemmer GR, Khouri RK, Murray GF, et al. Bilateral pneumothoraces secondary to iatrogenic buffalo chest: an unusual complication of median sternotomy and subclavian catheterization. Ann Surg 1984;199:372–4. 6. Johri S, Berlin D, Sanders A. Bilateral pneumothoraces after unilateral transthoracic needle biopsy of a lung nodule. Chest 2003;123:1297–9. 7. Kalassian KG, Doyle R, Kao P, et al. Lymphangioleiomyomatosis: new insights. Am J Respir Crit Care Med 1997;155:1183–6. 8. Wait MA, Estrera A. Changing clinical spectrum of spontaneous pneumothorax. Am J Surg 1992;164:528–31. 9. Graf-Deuel E, Knoblauch A. Simultaneous bilateral spontaneous pneumothorax. Chest 1994;105:1142–6. 10. Pettersson GB, Gatzinsky P, Selin K. A case of total bilateral spontaneous pneumothorax. Scand J Thorac Cardiovasc Surg 1983;17:175. 11. Donovan PJ. Bilateral spontaneous pneumothorax: a rare entity. Ann Emerg Med 1987;16:1277–80. 12. Chu SC, Horiba K, Usuki S, et al. Comprehensive evaluation of 35 patients with lymphangioleiomyomatosis. Chest 1999;115:1041–52. 13. Muller NL, Chiles C, Kullnig P. Pulmonary lymphangioleiomyomatosis: correlation of CT with radiographic and functional findings. Radiology 1990;175:335–9. 14. Sayar A, Turna A, Metin M, et al. Simultaneous bilateral spontaneous pneumothorax report of 12 cases and review of the literature. Acta Chir Belg 2004;104:572–6.



CASE REPORT

Pulmonary and nodal multiple myeloma mimicking lymphoma P O’SULLIVAN,

FFRRCS

¨ LLER, and N L MU

MD, PhD

Department of Radiology, University of British Columbia, Vancouver, BC, Canada ABSTRACT. Multiple myeloma is a haematological malignancy characterized by the occurrence of plasma cell tumours within the bone marrow. In advanced multiple myeloma, metastatic deposits outside the bone marrow (extramedullary) are rare. Reported extramedullary sites include the upper respiratory tract and lymph nodes. Parenchymal pulmonary plasmacytoma is exceptionally rare. We describe such a case in a 51-year-old male, who underwent allogenic marrow transplantation. 6 months posttransplantation developing nodal and hilar pulmonary multiple myeloma, the radiological appearances mimicking lymphoma.

Multiple myeloma (MM) is a systemic disease process primarily involving the bone marrow. In MM the bone marrow is replaced by malignant plasmacytes, which produce monoclonal proteins. This disease process primarily involves the axial skeleton. Extramedullary plasmacytoma (EMP) are tumours of malignant plasmacytes occurring outside the bone marrow. A solitary plasmacytoma is a single focal tumour of malignant plasmacytes that may occur in or outside the bone marrow. EMP are uncommon and typically manifest as a solitary plasmacytoma. Solitary plasmacytomas occur most commonly in the upper respiratory tract [1] such as nasal cavities, paranasal sinuses and nasopharynx. Approximately 15% of patients with extramedullary plasmacytoma progress to full multiple myeloma [2]. Localized intrapulmonary plasmacytoma is rare, representing fewer than 5% of all extramedullary plasmacytomas [3]. We describe a case of MM with pulmonary and lymph node involvement. To our knowledge the CT findings of extramedullary pulmonary multiple myeloma have not been previously described.

Case report A 51-year-old man with a background history of light chain disease MM stage 4b had previously undergone an allogenic sibling bone marrow transplant. Initially, he made a satisfactory clinical recovery. However, a followup bone marrow aspirate performed at 2 months posttransplant revealed some atypical and immature plasma cells raising concern for residual disease. 6 months after transplantation, the patient re-presented to the emergency department with increasing cough and shortness of breath. A chest radiograph was performed which demonstrated bilateral hilar lymphadenopathy and Address correspondence to: Dr Paul O’Sullivan, Department of Radiology, Vancouver General Hospital, 899 West 12th Avenue, Vancouver, BC, Canada, V5Z 1M9.


Received 13 May 2005 Revised 11 July 2005 Accepted 14 July 2005 DOI: 10.1259/bjr/77207966 ’ 2006 The British Institute of Radiology

lung nodules. The findings raised concern for a recurrence and a bone marrow aspirate was performed, which revealed a hypocellular marrow with trilineage engraftment and monoclonal plasma cells inkeeping with recurrence. In view of the suspicion of malignancy on the chest radiograph, an intravenous contrast enhanced CT was ordered. This confirmed the presence of extensive bilateral hilar lymphadenopathy (Figure 1a) and multiple lung nodules. The lung nodules were round, smooth or lobulated, and had soft tissue attenuation. They measured from 1.1 cm 6 1.0 cm in size (anteroposterior 6 transverse diameter) in the left upper lobe to 4.9 cm 6 4.8 cm diameter in the left lower lobe (Figure 1b). The initial differential diagnosis, based on the CT findings included lymphoma and post-transplant lymphoproliferative disease. The hospital respirology and haematology services were consulted, and the patient subsequently underwent a bronchoscopy and transbronchial biopsy. The biopsy samples contained a homogeneous cellular infiltrate with immunophenotypic features of plasma cells, consistent with recurrent multiple myeloma (Figure 2). Despite haematological and intensive care supportive measures, including treatment for opportunistic pathogens, the patient’s respiratory function progressively deteriorated over the next 2 weeks. The patient died 16 days after admission of progressive respiratory failure. No post mortem was performed.

Discussion The occurrence of extramedullary disease is uncommon in MM. In one study of a group of 432 patients with MM, only 19 (4.4%) patients were identified as having extramedullary plasmacytoma (EMP) [4]. In the same study, a total of 48 solitary plasmacytomas were identified in 19 patients with extramedullary disease. The most common sites were lymph nodes in 10 (20.8%), the pleura in 8 (16.6%) and soft tissues in 6 (12.5%). Only e25

P O’Sullivan and N L Mu¨ller

(a)

(b)

Figure 1. A 51 year-old man with extramedullary myeloma. (a) Mediastinal windows: CT image shows large bilateral hilar lymph nodes (right arrow) compressing the pulmonary artery and bronchus intermedius. (b) Lung windows: CT image shows bilateral lung nodules and masses (Hounsfield range 65¡12) (arrows).

3 (6.2%) lesions occurred within the lung parenchyma. The prognosis of patients with pulmonary multiple myeloma is poor [5]. This contrasts with reports of long survival rates with primary pulmonary plasmacytomas of the lung [6]. The most typical thoracic manifestations of multiple myeloma are bony involvement of the thoracic cage or a pulmonary infiltrate secondary to infection [7]. One author has recently described 13 patients with multiple myeloma and pulmonary complications. Six (46%) patients had pneumonia, 2 (15%) had mass lesions, 2 (15%) had multiple nodular lesions and 3 (23%) had interstitial infiltrates [7]. Other described manifestations of myeloma in the lung include; a diffuse reticulonodular pattern [8] or pulmonary calcification [9]. Amyloid, often seen with multiple myeloma has several manfestations in the lung. Nodular parenchymal disease (multiple welldefined nodules), tracheobronchial (focal or diffuse

thickening of the airway) and diffuse parenchymal (also termed alveolar septal) amyloidosis [10]. The CT findings in our patient of bilateral hilar adenopathy with round pulmonary nodules are unusual for MM, and resemble more those of lymphoma and post-transplant lymphoproliferative disorder (PTLD). Immunophenotyping is a technique used to differentiate individual cell lineages. This is done by attaching surface antibodies to cell specific surface antigens. Myeloma has been shown to stain positively with CD138 [11]. Both bronchoscopic biopsy samples stained strongly positive for CD 138 plasma cell markers. Also, in PTLD we would expect immunophenotypic features of B cell lymphoma and Epstein Barr virus (EBV) positivity, both of which were absent. In conclusion, although the CT features and clinical history in this case would strongly support PTLD or lymphoma, the biopsy and immunophenotyping revealed a very unusual manifestation of multiple myeloma.

References

Figure 2. Medium power view (206) of bronchoscopy biopsy infiltrate, with CD 138 plasma cell marker, exhibiting a strongly positive brown stain. These features are characteristic of multiple myeloma. e26

1. Wise JN, Schaefer RF, Read RC. Primary pulmonary plasmacytoma: a case report. Chest 2001;120:1405–7. 2. Galieni P, Cavo M, Pulsoni A, Avvisati G, et al. Clinical outcome of extramedullary plasmacytoma. Haematologica 2000;85:47–51. 3. Lacaze O, Khaddage A, Court-Fortune I, et al. [Isolated intrapulmonary plasmacytoma; diagnostic and therapeutic difficulties]. Rev Mal Respir 2002;19:648–50. 4. Damaj G, Mohty M, Vey N, et al. Features of extramedullary and extraosseous multiple myeloma: a report of 19 patients from a single center. Eur J Haematol 2004;73:402–6. 5. Oymak FS, Karaman A, Soyuer I, et al. [Pulmonary and chest wall involvement in multiple myeloma]. Tuberk Toraks 2003;51:27–32. 6. Koss MN, Hochholzer L, Moran CA, et al. Pulmonary plasmacytomas: a clinicopathologic and immunohistochemical study of 5 cases. Ann Diagn Pathol 1998;2:1–11. 7. Oymak FS, Karaman A, Soyeur I, et al. [Pulmonary and chest wall involvement in multiple myeloma]. Tuberk Toraks 2003;51:27–32 [In Turkish].


Case report: Multiple myeloma mimicking lymphoma 8. Shin MS, Carcelen MF, Ho KJ. Diverse roentgenographic manifestations of the rare pulmonary involvement in myeloma. Chest 1992;102:946–8. 9. Weber CK, Friedrich JM, Merkle E, Prummer O, et al. Reversible metastatic pulmonary calcification in a patient with multiple myeloma. Ann Hematol 1996;72: 329–32.


10. Kim HY, Im JG, Song KS, et al. Localised amyloidosis of the respiratory system: CT features. J Comput Assist Tomog 1999;23:627–31. 11. Costes V, Magen V, Legouffe E, Durand L, et al. The Mi15 monoclonal antibody (anti-syndecan-1) is a reliable marker for quantifying plasma cells in paraffin-embedded bone marrow biopsy specimens. Hum Pathol 1999;30:1405–11.

e27


CASE REPORT

Diffusion-weighted imaging findings in Balo concentric sclerosis E C KAVANAGH, MRCPI, FFR, RCSI, M K S HERAN, MD, D M FENTON, MD, J S LAPOINTE, MD, R A NUGENT, MD and D A GRAEB, MD Division of Neuroradiology, Department of Radiology, Vancouver General Hospital, 899 West 12th Avenue, Vancouver, British Columbia, V5Z 1M9, Canada ABSTRACT. Balo concentric sclerosis is a rare demyelinating disease. Pathognomonic features have been previously described. Diffusion-wighted imaging findings have not been previously described in Balo concentric sclerosis. We describe the diffusionweighted imaging findings in a 45-year-old lady with Balo concentric sclerosis. Diffusion-weighted imaging offers insight into the possible pathophysiology of this rare disease.

Received 1 June 2005 Revised 12 July 2005 Accepted 19 July 2005 DOI: 10.1259/bjr/36636301 ’ 2006 The British Institute of Radiology

Balo concentric sclerosis is an uncommon demyelinating disorder, believed to be a variant of multiple sclerosis [1, 2]. Imaging features pathognomonic for Balo concentric sclerosis include alternating rings of demyelinated and myelinated white matter seen on T2 weighted images with concentric ring enhancement on T1 weighted images post-administration of gadolinium [2]. We describe the diffusion-weighted MRI findings in a 45year-old woman with Balo concentric sclerosis.

Case report A 45-year-old, right handed lady presented with a 1 month history of intermittent right arm weakness and slurred speech. Motor and reflex examinations were normal. A non-contrast CT scan performed at another institution had suggested the presence of an infarct in the left corona radiata. An MR of the brain was then performed which revealed a lesion in the left posterior corona radiata with multiple concentric rings of T2 hyperintensity (Figure 1a). The central core of the lesion showed high signal on diffusion-weighted images, consistent with T2 shine-through. The intervening ring showed evidence of unrestricted diffusion (Figure 1b,c). The outer ring showed evidence of mildly restricted diffusion in its most anterior aspect (Figure 1c, arrow), with an apparent diffusion coefficient (ADC) value of 0.8461023 mm2 s21, while the remainder of this ring demonstrated T2 shine-through. The calculated ADC values were 1.3861023 mm2 s21 for the central core, 1.3361023 mm2 s21 for the remainder of the outer ring and 2.4061023 mm2 s21 for the intervening ring of unrestricted diffusion. The ADC value in the posterior corona radiata of the normal contralateral hemisphere measured 1.0161023 mm2 s21. Mild concentric enhancement was seen post-administration of gadolinium. The patient was commenced on a course of oral steroids and her neurological symptoms improved after 2 weeks of e28

therapy. A repeat MR performed 7 months later showed marked reduction in size of the lesion (Figure 2a). The previously noted areas of T2 shine-through had almost completely resolved (Figure 2b,c). The ADC value for the entire lesion was 0.9461023 mm2 s21.

Discussion Many case reports of Balo concentric sclerosis exist in the medical literature [1–6]. Historically the diagnosis was made post mortem and, until recently, Balo concentric sclerosis was believed to be a rapidly progressive and invariably fatal disease [3]. MRI can allow for this diagnosis to be made early, thus allowing appropriate therapy to be commenced. Corticosteroids have been shown to be effective in treating the neurological deficits associated with Balo concentric sclerosis [4]. The pathogenesis of this disease has been a topic of much debate within the pathology literature. Various theories exist as to the development of the typical findings of alternating zones of myelinated and demyelinated white matter. It has been proposed that these bands may represent areas of remyelination at the borders of areas of successive acute demyelination [5]. Others have postulated that the partially myelinated bands may represent the early stages of acute demyelination, rather that remyelination. Pathognomonic MRI features in Balo concentric sclerosis have been proposed [2]. These features include alternating bands of demyelinated and myelinated white matter seen on T2 weighted images and the presence of concentric ring enhancement on T1 weighted images post-administration of gadolinium. MR spectroscopy demonstrated an increased choline peak and a decreased N-acetyl aspartate peak in one reported case of Balo concentric sclerosis [1]. Diffusion-weighted imaging findings in Balo concentric sclerosis have not been The British Journal of Radiology, July 2006

Case report: DWI for Balo concentric sclerosis

(a)

(b)

(c) Figure 1. A 45 year-old lady with Balo concentric sclerosis. (a) Axial T2 weighted image shows a lesion in the posterior left corona radiata with concentric rings of T2 hyperintensity. (b) Axial diffusion-weighted (b1000) image shows high signal within the centre and periphery of the lesion. An intervening ring of relatively unrestricted diffusion is seen. (c) Axial apparent diffusion coefficient map confirms the presence of restricted diffusion within the most anterior aspect of the outer ring (arrow). The intervening ring of relatively unrestricted diffusion is also confirmed.


e29

E C Kavanagh, M K S Heran, D M Fenton et al

(a)

(b)

(c) Figure 2. A 45-year-old lady with Balo concentric sclerosis (MRI performed 7 months after Figure 1). (a) Axial T2 weighted image shows reduction in size of the previously identified lesion in the posterior left corona radiata. (b) Axial diffusion-weighted image shows some residual ill-defined high signal in the lesion, consistent with T2 shine-through. (c) Axial apparent diffusion coefficient map shows no evidence of restricted diffusion.

e30


Case report: DWI for Balo concentric sclerosis

previously described. MRI findings are variable in patients with multiple sclerosis and lesions may appear hyperintense, isointense or hypointense on diffusion-weighted sequences [7, 8]. High signal on diffusion-weighted images in multiple sclerosis has been ascribed to T2 shine-through, making interpretation of apparent diffusion coefficient maps important. In our case, the high signal on diffusion-weighted images in the anterior aspect of the outermost ring truly represented restricted diffusion as shown on the apparent diffusion coefficient map. Restricted diffusion of free water has been described in areas of vacuolization and rapid demyelination [8]. The area of restricted diffusion seen in the anterior aspect of the lesion in this case may represent an area of active demyelination. This area did not show evidence of significantly restricted diffusion 7 months later. The concentric ring of unrestricted diffusion likely represents an area of encephalomalacia secondary to demyelination. The central core of the lesion showed high signal on the T2 and diffusionweighted images, without evidence of ADC abnormality, consistent with T2 shine-through. The central core showed less T2 hyperintensity on follow-up and likely represents an area demyelinated white matter, without established encephalomalacic change. In this case more concentric rings were seen on T2 weighted images than on the diffusion-weighted images. T2 weighted sequences have higher spatial resolution than diffusionweighted images and the smaller rings seen on T2 weighted images are likely to be beyond the resolution of the diffusion-weighted sequences. The underlying pathophysiological explanation for the concentric rings seen in cases of Balo concentric sclerosis remains unknown, but it is likely that the various rings represent different stages of demyelination, a theory that is


supported by the diffusion-weighted findings shown here. In this case report we have demonstrated the diffusion-weighted imaging findings in Balo concentric sclerosis. Diffusion-weighted imaging offers insight into the possible pathophysiology of this unusual manifestation of multiple sclerosis.

References 1. Karaarslan E, Altintas, A, Senol U, Yeni N, Dincer A, Bayindir C, et al. Balo’s concentric sclerosis: Clinical and radiologic features of five cases. AJNR Am J Neuroradiol 2001;22:1362–7. 2. Caracciolo JT, Murtagh RD, Rojiani AM, Murtagh FR. Pathognomonic MR imaging findings in Balo concentric sclerosis. AJNR Am J Neuroradiol 2001;22:292–3. 3. Balo J. Encephalomyelititis periaxialis concentrica. Arch Neurol Psychiatry 1928;19:242–64. 4. Ng SH, Ko SF, Cheung YC, Wong HF, Wan YL. MRI features of Balo’s concentric sclerosis. Br J Radiol 1999;72:400–3. 5. Moore GR, Neumann PE, Suzuki K, Lijtmaer HN, Traugott U, Raine CS. Balo’s concentric sclerosis: new observations on lesion development. Ann Neurol 1985;17:604–11. 6. Chen CJ, Chu NS, Lu CS, Sung CY. Serial magnetic resonance imaging in patients with Balo’s concentric sclerosis: natural history of lesion development. Ann Neurol 1999;46:651–6. 7. Castriota-Scanderbeg A, Tomaiuolo F, Sabatini U, Nocentini U, Grasso MG, Caltagirone C. Demyelinating plaques in relapsing-remitting and secondary progressive multiple sclerosis: assessment with diffusion MR imaging. AJNR Am J Neuroradiol 2000;21:862–8. 8. Stadnik TW, Demaerel P, Luypaert RR, Chaskis C, Van Rompaey KL, Michotte A, et al. Imaging tutorial: differential diagnosis of bright lesions on diffusion-weighted MR images. Radiographics 2003;23:e7.

e31


CASE REPORT

Intravasation of barium sulphate at barium enema examination 1

J S WHITE, BMedSci(Hons), PhD, FRCS, 1R T SKELLY, MB, FRCS, 1K R GARDINER, MD, MCh, FRCS, 2J LAIRD, FRCR and 1M C REGAN, MD, FRCS Departments of 1Surgery and 2Radiology, Royal Victoria Hospital, Grosvenor Road, Belfast BT12 6BJ, UK ABSTRACT. We report a case of venous intravasation of barium sulphate occurring during a routine barium enema examination for investigation of rectal bleeding. The patient suffered a cardiopulmonary arrest, but made a full recovery after organ support in intensive care. Review of radiographs from the examination showed intravasated barium in pelvic vessels. We review the literature on this rare, but serious, complication of barium enema examination and suggest measures by which intravasation can be prevented.

Received 18 April 2005 Revised 12 August 2005 Accepted 2 September 2005 DOI: 10.1259/bjr/57839881 ’ 2006 The British Institute of Radiology

Case report Over 15 000 barium enema examinations are performed annually in Northern Ireland. Circulatory intravasation of barium sulphate is a rare, but potentially fatal, complication of this procedure. We report a case of barium intravasation occurring during a routine barium enema examination. A 77-year-old woman was admitted with a short history of passing dark blood mixed with the stool. She had a history of colonic diverticular disease, hypertension, previous stroke and exertional angina. Digital rectal examination, gastroscopy and rigid sigmoidoscopy were normal. Barium enema examination was carried out using a balloon catheter, which was placed and inflated by an experienced radiographer. Barium sulphate was instilled and screening commenced. Contrast was noted in a linear structure outside the rectum and instillation of barium was halted, the balloon deflated and the catheter removed. The chest and abdomen were screened for signs of barium, but none were apparent. Within a few minutes, the patient complained of feeling faint and suffered a cardiopulmonary arrest. Cardiopulmonary resuscitation was successful and the patient was transferred to intensive care, where she required ventilatory and supportive therapy. Electrocardiography and cardiac enzyme profiles were normal and there were no neurological findings suggesting a cerebrovascular event. The patient developed disseminated intravascular coagulation, which responded to treatment with fresh frozen plasma and platelets. After 3 days in intensive care, she returned to the ward. Review of the barium enema radiographs confirmed the presence of barium sulphate in pelvic veins, consistent with intravasation from the vagina (Figure 1), although examination per Address correspondence to: Mr Jonathan S White, Royal Alexandra Hospital, 10240 Kingsway, Edmontion, AB, T5H 3V, Canada.

e32

rectum and per vagina showed no evidence of injury. The patient subsequently made a full recovery and returned home again.

Discussion Barium enema, one of the most common radiological investigations, is considered safer than colonoscopy as an investigation for colorectal symptoms [1, 2]. The most common serious complication of barium enema is colonic perforation, which occurs in 0.004–0.04% of procedures [3–5]. Barium may also intravasate into the venous drainage of the large bowel and enters the circulation as a ‘‘barium embolus’’. This complication occurs rarely: in addition to the case described here, 36 cases of barium intravasation have been reported in the last 50 years [6–25] (Table 1). Barium intravasation due to rectal injury at the time of barium enema is thought to account for 8 of the 36 cases reported. Thinning of the rectal wall with age and proximity of the haemorrhoidal venous plexus may contribute to intravasation [6]. Intravasation at a site in the colon affected by disease accounted for a further 12 cases. It is thought to occur when intraluminal pressure overcomes the resistance of the colonic wall affected by colitis, diverticulitis or intestinal obstruction [7, 26]. 14 cases of barium intravasation occurring after vaginal insertion of the rectal catheter have also been reported [8–10, 13, 16]. These cases occur most often in elderly patients [8] and contributing factors include deficient perineal musculature, patient confusion, age-related vaginal thinning and use of a balloon catheter, which may prevent reflux of barium from the vagina [19]. In the two remaining cases, the site of barium entry into the venous system was unclear. Barium intravasating from the lower rectum or vagina enters the internal iliac venous system and is carried in The British Journal of Radiology, July 2006

Case report: Intravasation of barium sulphate

Figure 1. Screening radiograph from barium enema examination showing intravasation of barium sulphate into pelvic veins from a presumed vaginal injury.

the systemic circulation to the lungs, where it causes occlusion of the pulmonary circulation. Barium may also circulate within the systemic circulation to the heart, kidney and brain [8]. Barium intravasating from the upper rectum and colon enters the portal circulation and is carried to the liver before entering the systemic circulation. The mortality associated with portal embolisation is much lower than that for systemic embolisation (25% vs 60%, Table 1); this is probably because hepatic reticuloendothelial cells take up barium and reduce the amount entering the systemic circulation [6]. Portal intravasation tends to affect a younger age group than systemic intravasation (average age 44 years vs 77 years, Table 1); this may reflect a decreased likelihood of vaginal injury in younger patients due to better vaginal sensation. Systemic emboli also tend to involve a larger amount of barium with more significant cardiovascular effects [6]. As intravasation of barium sulphate is encountered very rarely, it may be difficult to recognize. It may be detected during screening at the start of the examination: if this is the case, it is recommended that the enema table is tilted to a ‘‘head up’’ position and the patient turned to

Table 1. Summary of fatal and non-fatal cases of barium intravasation published to date (chronological order)

Non-fatal cases Isaacs et al, 1950 Zatkin et al, 1964 Nordahl et al, 1973 Ansell et al, 1976 Juler et al, 1976 ten Kate et al, 1978 Archer et al, 1981 Blom et al, 1983 Chan et al, 1987 Fowlie et al, 1987 Kanehan et al, 1988 Baer et al, 1989 Taylor et al, 1990 Wheatley et al, 1991 Steinman et al, 1994 Zalev et al, 1997 Zalev et al, 1997 Chapman et al, 1998 Fatal cases Roman et al, 1952 Rosenberg et al, 1959 Truemner et al, 1960 Noveroske et al, 1966 Geipel et al, 1967 Frecker et al, 1968 Cove et al, 1974 Bayer et al, 1974 Partanen-Talsta et al, 1975 Ansell et al, 1976 Ansell et al, 1976 Salvo et al, 1976 Croiset van Uchelen et al, 1978 Bonte et al, 1980 Tsoutsanis et al, 1981 David et al, 1983 Haffner et al, 1993 Rossmann et al, 1997

Age (years) Sex

Balloon catheter used? Diagnosis

Site of intravasation

Type of embolus

25 65 39 62 35 68 73 73 36 38 59 58 79 27 57 29 75 85

F F M F M F M F F F M M F M F M F F

– yes yes – yes yes yes yes – yes – – no – – yes yes yes

colitis – diverticular – diverticular – carcinoma diverticular – Crohn’s diverticular diverticular – abscess – colitis diverticular –

proximal colon vagina sigmoid colon vagina sigmoid colon vagina proximal colon rectum vagina proximal colon sigmoid colon sigmoid colon rectum sigmoid colon vagina sigmoid colon vagina vagina

systemic systemic portal systemic portal systemic portal systemic portal portal portal portal systemic portal systemic portal systemic systemic

78 73 76 72 77 81 64 69 72 78 72 36 67 63 67 74 68 32

F F F F F F F F F F F F M F M F F M

yes yes yes yes – yes yes – – – – yes – – – yes yes –

– – – – – – diverticular – – – – colitis – – colitis carcinoma – diverticular

rectum rectum rectum rectum vagina rectum sigmoid colon unknown vagina vagina vagina proximal colon unknown vagina rectum vagina vagina sigmoid colon

systemic systemic systemic systemic systemic systemic systemic systemic systemic systemic systemic portal portal systemic systemic systemic systemic portal


e33

J S White, R T Skelly, K R Gardiner et al

the right to retard the passage of barium into the pulmonary circulation [27]. A plain pelvic radiograph may be helpful in establishing the diagnosis: Zatkin and Irwin described barium within the pelvic veins as a tangled mass of tubular structures resembling a ‘‘Medusa’s head’’ [28]. In portal intravasation, specific vessels such as the inferior mesenteric vein may be visible on plain radiography [7], and may be mistaken for contrast within the ureter [18]. Barium may also be detected within the liver, spleen, heart or lung on plain radiographs of chest and abdomen. CT is effective in demonstrating barium within the lungs and solid organs [15], but MRI is of little benefit [22]. Of the 37 cases of venous intravasation of barium sulphate reported to date, 18 (48.6%) proved fatal. Mortality in patients under 60 years of age was much lower than for those over 60 years (17% vs 64%, Table 1). Massive systemic embolisation usually causes immediate circulatory arrest; if a lesser amount of barium is involved, death may occur a few hours or days after the incident [6, 7, 16]. In younger patients in whom a small amount of barium enters the portal venous system, recovery is likely with supportive management. Patients with respiratory distress require endotracheal intubation and ventilation in an intensive care unit, with monitoring and support of cardiac, renal and hepatic function. Coagulation may be impaired after intravasation of barium and transfusion of clotting factors may be required [29]. Some authors have reported survival following treatment with corticosteroids and antibiotics [6]. Indications for surgery in the management of barium intravasation are unclear. Of the eight cases in which surgical intervention was attempted, four survived following partial colectomy [12, 18, 20, 30] while four others perished following laparotomy, attempted vaginal drainage or colectomy [8, 21, 27]. Two main strategies have been suggested to prevent barium intravasation at barium enema: confirmation of correct placement of the rectal catheter and reduction of the use of balloon catheters. It has been proposed that an appropriately-trained member of staff should inspect the perineum and perform a digital rectal examination before guiding the catheter into the rectum. A small amount of barium or air may also be insufflated at the start of screening, with the patient in a lateral position, to confirm that the catheter is correctly placed and to inspect the point of contact of the rectal catheter with the rectal wall [13, 19, 31]. There is considerable controversy over use of balloon catheters at barium enema. 20 of the 37 reports of barium intravasation were associated with use of a balloon catheter (Table 1). Over-inflation of the balloon may cause direct injury to the rectum or vagina; the balloon also prevents barium reflux, producing higher inflation pressures and increasing the risk of perforation or intravasation [17, 26]. Other authors believe that balloon catheters are safe if used carefully [30]. According to a recent UK survey, 22% of radiologists employ a balloon catheter routinely for barium enema and a further 43% use such a catheter occasionally, although only 5% follow all of the published safety recommendations, such as examining the rectal wall digitally before the balloon is inflated [31, 32]. The same study suggested that extraperitoneal perforation of the colon at barium enema was 2.5 times more likely when a e34

Table 2. Suggested guidelines for prevention of barium sulphate intravasation at barium enema 1. Use of balloon catheters Balloon catheters are only used in exceptional circumstances, and are only inserted by an appropriately-trained member of staff familiar with their use. 2. Screening with air Position of the rectal catheter is confirmed by screening while a small quantity of air from the tubing of the barium bag is introduced at the start of the procedure. This practice is based on the observation that intravasation of air is much less dangerous than intravasation of barium. Barium is only introduced once placement within the rectum is confirmed. 3. Screening with barium The rectum is observed fluoroscopically during the early stages of the examination to detect catheter malposition or barium intravasation. 4. Control of insufflation pressure Insufflation pressure is controlled by raising the barium bag to a fixed height above the table top: the bottom of the bag is placed 20 cm above the table top for introduction of air, and 45 cm above the table for introduction of barium.

balloon catheter was used: some centres have subsequently chosen to abandon the balloon catheter for routine use in barium enema. Close control of inflation pressures during the examination can be achieved by limiting the height of the barium bag, or by the use of pressure-controlled insufflation systems, and may reduce the risk of intravasation of barium. Subsequent to the case reported here, several changes were made in the practice of barium enema in this unit: reduction of the use of balloon catheters, pre-screening with air, careful initial screening with barium and control of insufflation pressure (Table 2). It is hoped that these measures will reduce the chances of encountering a further example of this serious complication of barium enema. In conclusion, intravasation of barium sulphate at barium enema is a rare complication of a common procedure and is associated with a high mortality rate. It is more likely to occur in elderly patients and in those with colorectal disease and has been associated with inadvertent vaginal placement of the rectal catheter. The diagnosis should be considered in any patient who collapses during or shortly after barium enema, and in those who become suddenly unwell in the hours following the procedure. The diagnosis can be confirmed by simple plain radiography; CT scanning may also be useful to detect dissemination of barium sulphate. Survival is more likely in younger patients, in cases of embolisation to the portal circulation and in those cases in which only a small amount of barium is involved. In those patients who survive the initial insult, conservative management with support of organ function is often successful. The complication may be prevented by ensuring correct placement of the rectal catheter and by reducing the use of balloon catheters.

References 1. de Zwart IM, Griffioen G, Shaw MP, Lamers CB, de Roos A. Barium enema and endoscopy for the detection of colorectal


Case report: Intravasation of barium sulphate

2.

3.

4.

5.

6.

7. 8.

9.

10.

11. 12.

13.

14.

15.

16.

neoplasia: sensitivity, specificity, complications and its determinants. Clin Radiol 2001;56:401–9. Kewenter J, Brevinge H. Endoscopic and surgical complications of work-up in screening for colorectal cancer. Dis Colon Rectum 1996;39:676–80. Gedebou TM, Wong RA, Rappaport WD, Jaffe P, Kahsai D, Hunter GC. Clinical presentation and management of iatrogenic colon perforations. Am J Surg 1996;172:454–7. Hakim NS, Sarr MG, Bender CE, Nivatvongs S. Management of barium enema-induced colorectal perforation. Am Surg 1992;58:673–6. Blakeborough A, Sheridan MB, Chapman AH. Complications of barium enema examinations: a survey of UK Consultant Radiologists 1992 to 1994. Clin Radiol 1997;52: 142–8. Fowlie S, Barton JR, Fraser GM. Barium embolisation during barium enema examination: a report of a case and a review of the literature. Br J Radiol. 1987;60:404–6. Archer FH, Freeman AH. A case of non-fatal intravasation of barium during barium enema. Br J Radiol 1981;54:69–72. David R, Berezesky IK, Bohlman M, et al. Fatal barium embolization due to incorrect vaginal rather than colonic insertion. An ultrastructural and x-ray microanalysis study. Arch Pathol Lab Med 1983;107:548–51. Haffner HT, Graw M. Fatal complications of a colon contrast enema due to erroneous insertion of the infusion catheter in the vagina. Dtsch Med Wochenschr 1993;118: 181–4. Partanen-Talsta A, Hyyppa S, Hirvonen J. Generalized fatal BaSO4 embolism from vaginal application of colonographic enema. A case report. Forensic Sci 1975;6:9–15. Tsoutsanis J. Intravasation of barium sulphate during roentgenography of the colon. Diagn Imaging 1981;50:1–3. Wheatley MJ, Eckhauser FE. Portal venous barium intravasation complicating barium enema examination. Surgery 1991;109:788–91. Zalev AH. Venous barium embolization, a rare, potentially fatal complication of barium enema: 2 case reports. Can Assoc Radiol J 1997;48:323–6. Zalev AH, Warren RE, Burnstein MJ. Venous intravasation of barium: CT findings. J Comput Assist Tomogr 1993;17:813–5. Chan FL, Tso WK, Wong LC, Ngan H. Barium intravasation: radiographic and CT findings in a nonfatal case. Radiology 1987;163:311–2. Bonte W, Sprung R. Lethal barium sulfate embolism after accidental vaginal application. Z Rechtsmed 1980;86:71–8.


17. Juler GL, Dietrick WR, Eisenman JL. Intramesenteric perforation of sigmoid diverticulitis with nonfatal venous intravasation. Am J Surg 1976;132:653–6. 18. Baer HM, Siegelbaum MH, Seidmon EJ, Kendall AR. Venous intravasation of barium into the inferior mesenteric vein mimicking a coloureteral fistula: a case report and literature review. J Urol 1989;141:117–9. 19. Chapman AH, Blakeborough A. Complications from inflation of a retention rectal balloon catheter in the vagina at barium enema. Clin Radiol 1998;53:768–70. 20. Kanehann LB, Caroline DF, Friedman AC, Lev-Toaff AS, Radecki PD. CT findings in venous intravasation complicating diverticulitis. J Comput Assist Tomogr 1988;12:1047–9. 21. Rossmann MD, Burr LJ, Thorpe PE. Colovenous fistula complicating diverticulitis: CT and radiographic findings. Abdom Imaging 1997;22:513–5. 22. Steinman RM, Torres G, Stoupis C. CT of abdominal barium intravasation. J Comput Assist Tomogr 1994;18:149–51. 23. Taylor DB, Yoong P. Non fatal barium intravasation during barium enema. Australas Radiol 1990;34:165–7. 24. ten Kate RW, Kalsbeek F. Non fatal intravasation of the barium enema accidentally introduced into the vagina. Tijdschr Gastroenterol 1978;21:277–80. 25. Croiset van Uchelen C, Bendel CJ, Nadorp JM, Bast TJ. Colonic perforation and barium sulfate intravasation in the course of roentgen studies. Ned Tijdschr Geneeskd 1978;122:1905–9. 26. Lewars MD. Barium embolisation during barium enema examination. Br J Radiol 1987;60:1236–7. 27. Ansell G. Diagnostic radiology. In: Diagnostic radiology. Oxford: Blackwell Scientific Publications, 1976:342–6. 28. Zatkin HR, Irwin GAL. Non-fatal intravasation of barium. AJR Am J Roentgenol 1964;92:1169–72. 29. Blom H, Nauta EH, van Rosevelt RF, ten Cate JW. Disseminated intravascular coagulation and hypotension after intravasation of barium. Arch Intern Med 1983;143:1253–5. 30. Fowlie S, Barton JR, Fraser GM. Barium embolisation during barium enema examination. Br J Radiol 1987;60:1236–7. 31. Dodds WJ, Stewart ET, Nelson JA. Rectal balloon catheters and the barium enema examination. Gastrointest Radiol 1980;5:277–84. 32. Blakeborough A, Sheridan MB, Chapman AH. Retention balloon catheters and barium enemas: attitudes, current practice and relative safety in the UK. Clin Radiol 1997;52:62–4.

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CASE REPORT

Epithelioid angiosarcoma of the lung: a rare late complication of Lucite plombage 1

D PATSIOS, BM BCh, MRCP, FRCR (UK), 2M DE PERROT, MD, FRCPC, 3M-S TSAO, MD, FRCPC and 1G WEISBROD, MD, FRCPC Departments of 1Medical Imaging and 2Thoracic Surgery, University Health Network, Toronto General Hospital, 200 Elizabeth Street, Toronto, Ontario, M5G 2C4 and 3Department of Pathology, University Health Network, University of Toronto, Toronto, Ontario M5G 2M9, Canada ABSTRACT. Epithelioid angiosarcoma of the lung is a rare late complication of Lucite plombage treatment of pulmonary tuberculosis. We describe the clinical, radiological and pathologic findings of a case of epithelioid angiosarcoma of the lung presenting with persistent haemoptysis who had undergone remote lung collapse therapy with Lucite plombage.


The discovery of effective antimicrobial therapy for the treatment of tuberculosis has effaced collapse therapy as the mainstream of therapy. Such therapy, until the 1950s, included artificial pneumothorax, phrenic nerve crush, thoracoplasty and extrapleural plombage. The latter therapeutic option has been associated with a number of late complications such as chronic pyothorax, migration of plombage, erosion of blood vessels and adjacent organs, and lung cancer [1, 2]. We describe an unusual case of an epithelioid angiosarcoma that developed in a collapsed lung, treated with Lucite plombage, presenting with persistent haemoptysis.

Case report A 76-year-old woman presented with haemoptysis of half a cup a day. She was well until 3 weeks prior to presentation. Her medical history included right upper hemithorax plombage collapse for right upper lobe tuberculosis in 1952. There was no history of tobacco smoking. A chest radiograph showed deviation of the trachea to the left and narrowing of the lumen of the distal thoracic trachea secondary to mass effect from the right upper lung, which contained 16 air-filled welldefined plombage balls, some of which demonstrated an air fluid level. The collapsed right upper lobe demonstrated soft tissue opacity in between the Lucite plombage and caused bulging of the elevated right horizontal fissure and the right oblique fissure on the lateral view of the chest radiograph (Figure 1). There was right upper chest wall rib deformity consistent with previous remote thoracic surgery and extrapleural plombage. A CT showed similar findings and confirmed the presence of soft tissue opacity between the Lucite e36

plombage, causing mass effect with deviation of the trachea to the left, an effaced right upper lobe bronchus and bulging of the oblique fissure and horizontal fissure (Figure 2). The differential diagnosis included an empyema, erosion of one of the Lucite balls into the remaining right upper lobe and neoplasia of the right upper lobe. Bronchoscopy demonstrated that bleeding was originating from the right upper lobe. The upper central airways were tortuous secondary to the pressure caused by the plombage cavity. The right upper lobe bronchus was extremely compressed and the orifice was narrowed to a few millimetres. There were no endobronchial lesions seen. Bronchial artery embolisation was attempted but failed. Thoracotomy revealed a foul smelling large plombage cavity filled with 16 Lucite plombage balls (Figure 3) and fibrin. The right pleural cavity was therefore decorticated and the atelectatic right upper lobe was resected. No tumoural lesion was seen macroscopically at the time of surgery. The patient had an uneventful post-operative recovery, her haemoptysis stopped, and she left the hospital after 1 week with no symptoms. Unfortunately, the patient re-presented 1 month later with septic shock secondary to a right haemorrhagic empyema. A right thoracotomy and decortication were performed in emergency, but the patient died 48 h later from septic shock. No organism was grown from the right pleural cavity during the first or second operation.

Pathology The resected right upper lobe showed thickened pleura that surrounded the lung parenchyma, which The British Journal of Radiology, July 2006

Case report: Epithelioid angiosarcoma of the lung

(a)

(b)

Figure 1. (a) The frontal radiograph demonstrated deviation of the trachea to the left, narrowing of the lumen of the distal thoracic trachea secondary to mass effect from the right upper lung, which contained multiple air filled well defined plombage balls, some of which demonstrated an air fluid level. The collapsed right upper lobe demonstrated soft tissue opacity in between the Lucite plombage. (b) The lateral view shows bulging of the elevated right horizontal fissure and the right oblique fissure.

contained a haemorrhagic and necrotic tumour of 7.6 cm. Histological examination of the pleura and tumour edges demonstrated infiltrating clusters of epithelioid tumour cells, which also showed prominent permeation of the lumens of capillaries and venules (Figure 4a,b). Within or bordering the necrotic areas, tumour cells assumed a

spindle cell appearance. Tumour cells demonstrated markedly pleomorphic nuclei with prominent nucleoli. Many tumour cells showed intracellular or intercellular vacuoles containing red blood cells. Their mitotic activity was brisk. Immunohistochemistry performed using routinely used diagnostic antibodies demonstrated

(a)

(b)

Figure 2. (a) The CT of the thorax confirmed the presence of a soft tissue opacity (big white block arrow) between the Lucite plombage, causing mass effect at the level of the right main bronchus (small white arrow) and (b) an effaced right upper lobe bronchus (black arrow) anterior to the right bronchus intermedius, which is consequently displaced caudally.


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D Patsios, M de Perrot, M-S Tsao and G Weisbrod

marker (AE1/AE3 and EMA), lymphohaematopoietic cell markers (CD45, CD3, CD5, CD30 and CD68), melanocytic marker (S100), but positive staining for vimentin (Figure 4c), CD31 (Figure 4d) and focally for Factor VIII. Tumour cells also stained negative for the mesothelial cell marker calretinin. The histological finding and immunophenotype were consistent with an epithelioid angiosarcoma. The peribronchial nodes were negative for tumour.

Discussion

Figure 3. 16 Lucite balls, removed from the excised right upper lobe.

negative staining of tumour cells for epithelial cell

Early spontaneous complications of plombage procedures include superinfection, migration of material and erosion of adjacent structures and broncho-extrapleural fistulae. In the long term, such complications become rare. Superinfection with Mycobacterium tuberculosis or common pyogenic bacteria has been the most frequent complication. The first sign is an acute onset of swelling of the collapsed space. The increased intracavitary tension leads to migration of the material, mechanical

(a)

(b)

(c)

(d)

Figure 4. (a) Tumour infiltrates the pleura and thickened fibrous tissue that encapsulated the necrotic lung parenchyma (haematoxylin and eosin, 650). (b) Tumour cells appear polygonal and epithelioid, and they infiltrate the blood vessel lumens (haematoxylin and eosin, magnification 6100). (c) Immunohistochemistry show strong positive staining of tumour cells for vimentin (magnification 6200). (d) Many tumour cells also show membranous staining for CD31, a marker of endothelial cells (magnification 6200).

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Case report: Epithelioid angiosarcoma of the lung

osteolysis of adjacent ribs, erosion of adjacent structures and fistula formation [1]. Patients may rarely experience pain because of the increased intracavitary pressure and haemoptysis if adjacent structures are eroded [2]. Malignant change induced by implanted foreign material has been documented in both experimental and clinical surgery [3]. Malignancy may occur rarely in plombage space or adjacent to plombage. Primary nonsmall cell bronchogenic carcinoma has been reported in non-smokers who had undergone Lucite plombage [4]. Foreign body reaction has been described as a contributing factor to the development of neoplasia in plombage cavities. Sarcomatous change such as malignant histiocytofibroma is recognized to be a specific complication of plombage [5]. Chronic infection itself has also been proposed as a predisposing factor for malignant fibrous histiocytoma and angiosarcoma [6–9]. The latter is a rare malignant tumour that affects various organs, mainly the skin, soft tissue, breast, liver, spleen and bone. Cases of angiosarcoma secondary to chronic pyothorax associated with tuberculosis have been described [6–9]. The calculated frequency rate of angiosarcoma associated with chronic pyothorax was 3600-fold higher than that of the general population [7]. Primary angiosarcoma of the pleura presents as a diffuse pleuropulmonary tumour macroscopically and microscopically simulating mesothelioma [7]. A history of pyothorax was involved in all cases of pleural angiosarcoma reported in a series of 26 cases in the literature [8]. However, it has been suggested that angiosarcoma associated with pyothorax seems to show the ordinary pathologic features of soft tissue angiosarcoma with no resemblance to primary pleural angiosarcoma [6]. Primary lung epithelioid angiosarcoma is a rare entity. Although it may present as a solitary lung mass [10], a unique case of epithelioid angiosarcoma arising in the lungs as bilateral multinodular infiltrates presented as pulmonary haemorrhage [11]. Pulmonary artery epithelioid angiosarcoma with pulmonary parenchymal invasion is also rare and has been reported to present with massive haemoptysis [12]. However, the majority of pulmonary angiosarcomas reported are usually due to secondary pulmonary involvement of soft tissue angiosarcomas [13]. The number of patients treated with plombage is attenuating rapidly and fewer patients will present in the


future with late complications. Clinicians and radiologists will therefore have less experience. Epithelioid angiosarcoma should be included in the differential diagnosis, even though a rare entity, when non-smokers who had undergone remote Lucite plombage lung collapse therapy present with haemoptysis and have radiological demonstration of increased mass effect at the site of plombage, as illustrated in the case we present above.

References 1. Massard G, Thomas P, Barsotti P, Rieva P, Giudicelli R, Reboud E, et al. Long-term complications of extraperiosteal plombage. Ann Thorac Surg 1997;64:220–5. 2. Weissberg D, Weissberg D. Late complications of collapse therapy for pulmonary tuberculosis. Chest 2001;120:847–51. 3. Lee YS, Pho RW, Nather A. Malignant fibrous histiocytoma at site of metal implant. Cancer 1984;54:2286–9. 4. Harland RW, Sharma M, Rosenzweig DY. Lung carcinoma in a patient with Lucite sphere plombage thoracoplasty. Chest 1993;103:1295–7. 5. Fauquert P, Saraux A, Guillermit MN, et al. Histiocytome fibreux malin thoracique. A propos d’un cas survenant apre`s pneumothorax therapeutique avec plombage par billes. Semaine des Hopitaux Paris 1989;65:2181–5. [Article in French]. 6. Hattori H. Epithelioid angiosarcoma arising in the tuberculous pyothorax. Report of an autopsy case. Arch Pathol Lab Med 2001;125:1477–9. 7. Aozasa K, Naka N, Tomita Y, Ohsawa M, Kanno H, Uchida A, et al. Angiosarcoma developing from chronic pyothorax. Modern Pathol 1994;7:906–11. 8. Zhang PJ, Livolsi VA, Brooks JJ. Malignant epithelioid vascular tumors of the pleura: report of a series and literature review. Human Pathol 2003;31:29–34. 9. Katsura H, Iuchi K, Matsumura A. Three cases of angiosarcoma complicated with chronic pleuritis. [Article in Japanese; Abstract in English]. Nihon Kokyuki Gakkai Zashi 2004;42:897–902. 10. Mukundan G, Urban BA, Askin FB, Fishman EK. Pulmonary epithelioid haemangioendothelioma: atypical radiologic findings of a rare tumor with pathologic correlation. J Comput Assist Tomogr 2000;24:719–20. 11. Sheppard MN, Hansell DM, Du Bois RM, Nicholson AG. Primary epithelioid angiosarcoma of the lung presenting as pulmonary haemorrhage. Human Pathol 1997;28:383–5. 12. Goldblum JR, Rice TW. Epithelioid angiosarcoma of the pulmonary artery. Human Pathol 1995;26:1275–7. 13. Lund L, Amre R. Epithelioid angiosarcoma involving the lungs. Arch Pathol Lab Med 2005;129:e7–e10.

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CASE REPORT

Response evaluation in gastrointestinal stromal tumours treated with imatinib: misdiagnosis of disease progression on CT due to cystic change in liver metastases 1

K M LINTON,

MBChB, MRCP,

2

M B TAYLOR,

MBChB, MRCP, FRCR

and 1J A RADFORD,

MBChB, MD, FRCP

1

Cancer Research UK Department of Medical Oncology and 2Department of Diagnostic Radiology, Christie Hospital NHS Trust, Wilmslow Road, Withington, Manchester M20 4BX, UK ABSTRACT. Imatinib is a highly effective treatment for patients with metastatic gastrointestinal stromal tumours (GIST). In most instances, response to imatinib treatment is assessed with CT. We present two cases where CT demonstrated the appearance of new low density liver lesions after 8–12 weeks of imatinib treatment. While this finding is consistent with progressive disease due to new lesions appearing at a previously uninvolved site, we hypothesise that the appearance of new liver lesions is in fact due to cystic change within previously occult, solid metastases. These untreated solid metastases were not visible on conventional portal phase CT due to their small size and vascular nature. Our hypothesis is supported by the observation that extrahepatic sites of disease had reduced in size over the same period of imatinib treatment and by the subsequent disease outcomes of these two cases. One patient, who continued imatinib because of significant symptomatic improvement despite the CT findings, remained stable on the same dose of imatinib for 18 months. The other patient, whose disease progressed when imatinib was withdrawn, had a dramatic response to treatment when imatinib was restarted at the same dose 2 years later. It is important that radiologists and oncologists who are involved in the management of GIST recognize that the appearance of new, lowdensity liver lesions on CT may represent a response to treatment. This finding must be correlated with symptomatic response and with tumour sites outside the liver before erroneously withdrawing effective imatinib treatment.

Gastrointestinal stromal tumours (GISTs) are rare tumours of mesenchymal origin, which can occur anywhere in the gastrointestinal tract, but most commonly arise in the stomach or small intestine [1]. GISTs were recently reclassified after discovering that these tumours harbour pathogenetically important mutations of the c-kit proto-oncogene (which encodes the KIT receptor) [2, 3]. Mutated KIT in turn leads to unchecked activation of the KIT tyrosine kinase pathway and downstream upregulation of cell growth and survival signals, resulting in GIST formation [4]. Imatinib mesylate (formerly STI-571, Glivec/Gleevec; Novartis, Basel, Switzerland) is a specifically designed tyrosine kinase receptor inhibitor, which targets, among others, the KIT receptor and interrupts neoplastic growth messages to the cell [5]. Imatinib has revolutionized the management of metastatic or unresectable GIST, which is largely resistant to conventional chemotherapy and radiotherapy, and remains the only effective licensed therapy available to these patients [6]. Assessment of tumour response to treatment with CT is the most common end-point for evaluating the clinical efficacy of imatinib and is the main determinant for continuation or discontinuation of therapy for individual patients. We report two cases of imatinib-treated e40

Received 19 May 2005 Revised 27 July 2005 Accepted 10 August 2005 DOI: 10.1259/bjr/62872118 ’ 2006 The British Institute of Radiology

metastatic and unresectable GIST where CT evaluation of the liver following imatinib showed the appearance of new low-density lesions, which were misinterpreted as disease progression.

Case 1 A 56-year-old man presented with weakness, anorexia, fever and upper abdominal pain due to a large, unresectable, c-kit positive duodenal GIST with synchronous liver metastases. CT showed a 9.3 cm predominantly solid mass involving the third part of the duodenum and the pancreatic head (Figure 1a). There was also a 2.0 cm ring enhancing lesion in the liver and 3 subtle, solid, low-density liver lesions. The largest of the low density lesions measured 1.6 cm and had a density of 95 Hounsfield Units (HU). Imatinib 400 mg daily was started in 2002. Within 1 week of starting treatment, he experienced a dramatic improvement of his tumour symptoms and performance status. A response evaluation CT after 8 weeks of imatinib showed that the mass in the duodenum and pancreas had reduced to 7.4 cm and remained solid (Figure 1c). However, between 40 and 50 new low-density liver The British Journal of Radiology, August 2006

Case report: Misdiagnosis in GIST after imatinib

(a)

(b)

(c)

(d)

Figure 1. Patient 1: contrast enhanced CT prior to treatment with imatinib. (a) The primary gastrointestinal stromal tumour (GIST) (outlined) involves the duodenum and pancreatic head and measures 9.3 cm. (b) No liver metastases are visible at this level, although three subtle lesions were seen elsewhere in the liver. Patient 1: contrast enhanced CT after 8 weeks of treatment with imatinib. (c) The duodenal mass (outlined) has reduced in size to 7.4 cm. (d) Multiple small low-density lesions are now seen within the liver (same level as Figure 1b). These were initially reported as being new metastases, but subsequently were considered as being due to cystic change in pre-existing metastases.

lesions were apparent on contrast-enhanced CT, the largest measuring 1.0 cm. The previously identified low density liver lesions had become cystic (density 28 HU). They had also slightly increased in size and the largest now measured 2.0 cm. The ring enhancing lesion was unchanged. The appearance of multiple new liver lesions on CT (Figure 1b,d) was initially reported as disease progression (PD). However, as he had experienced significant symptomatic benefit and because the primary tumour was smaller, imatinib was continued. A repeat CT after a further 8 weeks of treatment showed no further change and all sites of disease remained stable on the same dose of imatinib for the next 2 years.

Case 2 A 67-year-old woman underwent resection of a jejunal leiomyosarcoma in 1993. In 1999 she developed liver The British Journal of Radiology, August 2006

metastases and underwent left hemi-hepatectomy. Within 6 months of surgery she developed a pelvic mass and omental nodules. Biopsy of the pelvic mass confirmed metastatic sarcoma. Shortly afterwards, a metastasis in the second lumbar vertebra as well as new liver metastases were diagnosed. The disease failed to respond to sequential single-agent chemotherapy with doxorubicin and ifosfamide. Her original biopsy was re-examined and found to be c-kit (CD117) and CD34 positive. The diagnosis was therefore revised to GIST and imatinib 400 mg daily was commenced in 2002. The pre-treatment CT showed an 8.0 cm solid pelvic mass (Figure 2a), 2 cystic liver metastases measuring 4.5 cm and 3.0 cm and small omental nodules. The cystic liver metastases had densities of 23–25 HU and the larger metastasis showed enhancing nodules within its wall. A response evaluation CT after 12 weeks of imatinib treatment showed that the mass in the pelvis had e41

K M Linton, M B Taylor and J A Radford

(a)

(b)

(c)

(d)

Figure 2. (a,b) Contrast enhanced. Patient 2: contrast enhanced CT prior to treatment with imatinib. (a) There is an 8 cm solid pelvic mass (outlined). (b) No liver metastases are visible on this section. Note surgical clips from previous hepatic resection (arrow) and L2 vertebral metastasis (*). Patient 2: contrast enhanced CT after 12 weeks of treatment with imatinib. (c) The pelvic mass has reduced in size to 5.1 cm. (d) (At the same level as Figure 2b) A low lesion (arrow) is one of eight newly visible lesions in the liver.

reduced to 5.1 cm (Figure 2c). The density had fallen slightly, but the lesion remained solid. The omental nodules were also slightly smaller. The cystic liver metastases, however, had slightly enlarged to 5.1 cm and 3.1 cm, respectively, and 8 new low-density liver lesions, measuring between 3 mm and 8 mm, were seen on CT (Figure 2b,d). The larger cystic metastases were unchanged in density, but the enhancing nodules within the wall of the larger lesion had resolved. Since starting imatinib she developed new abdominal pain and anaemia, which required hospital admission. Because of this, and the appearance of new liver lesions on CT, disease progression was diagnosed and imatinib was stopped. A restaging CT performed 18 months later showed enlargement of the pelvic mass to 14.4 cm. There were also several new bone metastases, including a large destructive sternal metastasis. The liver lesions were relatively unchanged from before. At this time, her previous CT examinations were reviewed. In the light of our experience with the first case, it was considered that the development of new lowdensity liver lesions on the first response evaluation CT e42

may also have been due to cystic response within previously occult metastases. Imatinib was therefore restarted at the same dose of 400 mg daily, almost 2 years after her first exposure to the drug. A repeat CT after 8 weeks of treatment showed a dramatic improvement. The pelvic mass had reduced to 9.4 cm and the sternal metastasis was also significantly smaller. The liver lesions were little changed. On this occasion, her tumour symptoms and performance status improved significantly. She has now been followed-up for over 12 months and her disease has continued to respond to imatinib.

Discussion Cystic change is a well recognized feature of certain tumours that are responding to treatment (such as nonseminomatous germ cell tumours) [7]. In their paper of 2002, Chen et al first reported how previously solid metastases from GIST can become cystic on contrastenhanced CT following successful treatment with The British Journal of Radiology, August 2006

Case report: Misdiagnosis in GIST after imatinib

imatinib. They observed that treated metastases had a slightly higher density (in the range of 21–29 HU) than would be expected for simple cysts [8]. Another study examined the serial histological changes in GIST treated with imatinib and confirmed that treatment-induced cystic change is associated with tumour response: compared with their previously solid counterparts, cystic metastases had lower mitotic counts, lower proliferation indices, loss of cellular cohesion and shrinkage of tumour tissue [9]. It is recognized that liver metastases from GIST are frequently hypervascular prior to treatment [10]. These lesions may therefore not be visible on conventional portal venous phase CT, which is the most common CT technique used in oncology patients. In both of our cases only portal venous phase CT was used and small untreated solid liver metastases were not visible. The appearance of new lesions in the liver after imatinib treatment was therefore interpreted as PD. According to conventional RECIST criteria, PD is defined as >20% increase in the sum of the long axis of up to 10 target lesions, or the appearance of new lesions [11]. After continuation of imatinib, we observed no further change in the cystic liver metastases and at extrahepatic tumour sites, either tumour stabilization (case 1) or shrinkage (case 2) occurred. These successful clinical outcomes gave retrospective evidence that initial treatment with imatinib was effective. They also corroborate our observation that the appearance of new liver lesions on CT in the first 8–12 weeks after starting treatment is due to increased lesion conspicuity following cystic change and not due to disease progression. The decision to continue treatment in the first case was based largely on the fact that the patient experienced considerable clinical improvement of tumour symptoms after imatinib treatment. Conversely, treatment was stopped in the second case where a symptomatic benefit was not apparent. Symptomatic improvement as a marker of response may be valuable. However, as our second case illustrates, the interpretation of symptoms that develop as a result of imatinib treatment may be misleading. This patient developed new symptoms of abdominal pain and anaemia soon after commencing imatinib, which were thought at the time to be due to underlying disease progression. GISTs are very vascular tumours and intratumoural haemorrhage following effective imatinib treatment is not uncommon [9]. Therefore, in retrospect, we conclude that in this case the development of pain and anaemia occurred because of imatinib-related haemorrhage into liver or pelvic metastases, and not because of underlying disease progression. Several other side effects of imatinib can overlap with tumour symptoms and should not be misdiagnosed as PD, including nausea, vomiting and oedema formation (especially if associated with the development of ascites) [6]. The development of new symptoms following imatinib must therefore be carefully evaluated to avoid a misdiagnosis of disease progression, especially in the context of apparently corroborative radiological evidence. The radiological techniques used in the initial staging and in the treatment response evaluation of imatinib must be critically evaluated. Functional imaging modalities such as PET (positron emission tomography) have The British Journal of Radiology, August 2006

been shown to provide significantly more accurate assessments of response to imatinib than CT, especially in the early post-treatment stage (first 8 weeks) [12, 13]. However, this modality is not widely available in the non-research setting. Due to the vascular nature of GISTs, it has been suggested that triple phase CT (including non-contrast, arterial phase and portal phase acquisitions) is more sensitive in the initial staging of patients [12]. Our experience has also found that untreated solid liver metastases may not be detectable on single portal venous phase CT, although we have not evaluated dual (arterial and portal venous phase) or triple phase techniques. There is evidence that dynamic gadolinium enhanced MRI may be superior to single phase CT in early disease assessment [10] but, to our knowledge, this technique has not been fully evaluated and has not been compared with the results from triple phase CT. According to strict RECIST criteria, cystic lesions should be regarded as non-measurable [11]. However, this may be difficult in GIST when target lesions, which are initially solid, only become cystic following imatinib exposure. Modified objective criteria using a combination of tumour size and density on CT have therefore been proposed for early response evaluation in GIST [14]. Despite its limitations, CT remains the most common means of evaluating treatment response to imatinib in GIST. Based on the current evidence, including our experience, we propose evaluating dual phase CT (arterial and portal venous phases) and gadolinium enhanced MRI in the initial staging and treatment assessment of GIST, using modified objective criteria to evaluate response to imatinib. Currently, neither NICE (National Institutes for Clinical Excellence, UK) nor NCCN (National Comprehensive Cancer Network, USA) GIST treatment guidelines recognize symptomatic improvement as a marker of response [15, 16]. In light of our experience, we furthermore propose that these guidelines should be amended to include symptomatic improvement as a response criterion, especially when the interpretation of radiological findings is in doubt. Guidelines should also emphasise that response may be associated with the appearance of new cystic liver lesions, especially if single phase CT is used. It is essential that clinicians and radiologists who are involved in the treatment of GIST are aware of the limitations of CT response evaluation in order to avoid inappropriate withdrawal of imatinib treatment. Furthermore, we are obliged to review the radiology of all patients with GIST who may have stopped imatinib prematurely due to a misinterpretation of their response evaluation CTs, as these patients may benefit from the reintroduction of imatinib even after long periods off treatment.

References 1. Lev D, Kariv Y, Issakov J, Merhav H, Berger E, Merimsky O, et al. Gastrointestinal stromal sarcomas. Br J Surg 1999;86:545–9. 2. Hirota S. Gastrointestinal stromal tumors: their origin and cause. Int J Clin Oncol 2001;6:1–5.

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K M Linton, M B Taylor and J A Radford 3. Fletcher CD, Berman JJ, Corless C, Gorstein F, Lasota J, Longley BJ, et al. Diagnosis of gastrointestinal stromal tumors: a consensus approach. Hum Pathol 2002;33: 459–65. 4. Heinrich MC, Rubin BP, Longley BJ, Fletcher JA. Biology and genetic aspects of gastrointestinal stromal tumors: KIT activation and cytogenetic alterations. Hum Pathol 2002;33:484–95. 5. Joensuu H, Roberts PJ, Sarlomo-Rikala M, Andersson LC, Tervahartiala P, Tuveson D, et al. Effect of the tyrosine kinase inhibitor STI571 in a patient with a metastatic gastrointestinal stromal tumor. N Engl J Med 2001;344: 1052–6. 6. van Oosterom AT, Judson IR, Verweij J, Stroobants S, Dumez H, Donato di PE, et al. Update of phase I study of imatinib (STI571) in advanced soft tissue sarcomas and gastrointestinal stromal tumors: a report of the EORTC Soft Tissue and Bone Sarcoma Group. Eur J Cancer 2002;38 Suppl. 5:S83–7. 7. Husband JE, Schwartz LH, Spencer J, Ollivier L, King DM, Johnson R, et al. Evaluation of the response to treatment of solid tumours - a consensus statement of the International Cancer Imaging Society. Br J Cancer 2004;90:2256–60. 8. Chen MY, Bechtold RE, Savage PD. Cystic changes in hepatic metastases from gastrointestinal stromal tumors (GISTs) treated with Gleevec (imatinib mesylate). AJR Am J Roentgenol 2002;179:1059–62. 9. Reichardt P, Schneider U, Stroszczynski C, Pink D, Hohenberger P. Molecular response of gastrointestinal stromal tumour after treatment with tyrosine kinase inhibitor imatinib mesylate. J Clin Pathol 2004;57:215–7.

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10. Sandrasegaran K, Rajesh A, Rushing DA, Rydberg J, Akisik FM, Henley JD. Gastrointestinal stromal tumors: CT and MRI findings. Eur Radiol 2005;15:1407–14. 11. Padhani AR, Ollivier L. The RECIST (Response Evaluation Criteria in Solid Tumors) criteria: implications for diagnostic radiologists. Br J Radiol 2001;74:983–6. 12. Choi H, Charnsangavej C, de Castro FS, Tamm EP, Benjamin RS, Johnson MM, et al. CT evaluation of the response of gastrointestinal stromal tumors after imatinib mesylate treatment: a quantitative analysis correlated with FDG PET findings. AJR Am J Roentgenol 2004;183:1619–28. 13. Antoch G, Kanja J, Bauer S, Kuehl H, Renzing-Koehler K, Schuette J, et al. Comparison of PET, CT, and dual-modality PET/CT imaging for monitoring of imatinib (STI571) therapy in patients with gastrointestinal stromal tumors. J Nucl Med 2004;45:357–65. 14. Choi H. Critical issues in response evaluation on computed tomography: lessons from the gastrointestinal stromal tumor model. Curr Oncol Rep 2005;7:307–11. 15. National Institute for Clinical Excellence. Appraisal Consultation Document: Imatinib for the treatment of unresectable and/or metastatic gastro-intestinal stromal tumours. Available from: http://www.nice.org.uk/ page.aspx?o5117718e [Accessed 31 March 2006]. 16. Demetri GD, Dematteo RP. Optimal management of patients with gastrointestinal stromal tumours (GIST): an expansion on the NCCN sarcoma clinical practice guidelines in oncology. A Continuing Medical Education Presentation recorded at NCCN’s 9th Annual Conference Available from: http://www.nccn.org/professionals/ webcasts_gist.asp [Accessed 31 March 2006].



CASE REPORT

Radiation myelopathy after radioactive iodine therapy for spine metastasis 1

H MURAKAMI, K TOMITA, MD

MD,

1

N KAWAHARA,

MD,

1

T YAHATA,

MD,

2

K YOKOYAMA,

MD,

3

K KOMAI,

MD

and

1

Department of 1Orthopaedic Surgery, 2Nuclear Medicine, and 3Neurology and Neurobiology of Aging, Kanazawa University, Kanazawa, Japan ABSTRACT. A case of radiation myelopathy after radioactive iodine therapy is reported. This is the first report to describe radiation myelopathy after I-131 therapy. A 62-yearold female with spinal metastasis of T10 received I-131 therapy. She presented with radiation myelopathy 34 months after the irradiation. We need to recognize the possibility of this serious complication even in the case of I-131 therapy. There is a risk of radiation myelopathy even after I-131 therapy, especially in cases with spinal cord compression such as this.

Radiation therapy is commonly applied as a primary or adjuvant therapy for malignancies. One of the most serious complications following radiation therapy is myelopathy, particularly in long-term surviving patients. Post-radiation myelopathy occurs when the spinal cord is included within the radiation field in cases of high total radiation doses, or for high radiation doses per fractionation. The myelopathy may result in chronic progressive symptoms months or years after radiotherapy and no treatment has proved satisfactory. Radioactive iodine (I-131) therapy is a type of radiotherapy and has an important role in the treatment of thyroid cancer and its metastasis. However, there are no previous reports describing myelopathy after I-131 therapy. We report on the first case of radiation myelopathy due to I-131 administration in a case of thyroid cancer metastasis.

Received 20 June 2005 Revised 15 August 2005 Accepted 17 August 2005 DOI: 10.1259/bjr/16265478 ’ 2006 The British Institute of Radiology

Upon neurological examination, she had hypothermia and hypoalgesia in her left lower limb, and muscle weakness (3-4/5) and pyramidal signs in her right lower limb. That is, right side injured type Brown-Sequard syndrome was present. All laboratory data were within normal range. Plain X-ray film revealed a right pedicle sign and CT scans demonstrated an osteolytic lesion of T10. MRI demonstrated a tumour lesion with abnormally low signal intensity on T1 weighted images and high signal intensity on the T2 weighted images on the right side of the T10 vertebral body without clear enhancement after intravenous gadolinium (Gd) injection. In addition, it showed an appearance of high signal intensity from T10 to T5 on the T2 weighted images on

Case report A 62-year-old female underwent a total thyroidectomy for follicular thyroid carcinoma and a partial lobectomy for solitary lung metastasis in March 1999. At that time, spine metastasis of T10 was detected (Figure 1) and I-131 therapy was done three times (7.6 GBq each time) from September 1999 to March 2000. The patient received a total of 22.8 GBq of radioiodine, the routine amount for bone metastasis. In January 2003, she felt hypesthesia on her left lower extremity without any apparent trauma. Then, muscle weakness in her right lower extremity appeared and she was transferred to our hospital for further examination. Address correspondence to: Norio Kawahara, Department of Orthopaedic Surgery, Kanazawa University, 13-1 Takaramachi, Kanazawa, 920-8641, Japan.


Figure 1. On axial T2 weighted image taken in 1999, the tumour extended to the right vertebral body, right pedicle, and right lamina of T10.

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H Murakami, N Kawahara, T Yahata et al

(a)

(b)

Figure 2. MRI demonstrating abnormal intensity at T10. (a) Sagittal T2 weighted image showing high signal intensity on the spinal cord. (b) Axial T2 weighted image.

the spinal cord (Figure 2a). On axial images of the MRI, the tumour extended to the right vertebral body, right pedicle and right lamina of T10 (Figure 2b). The size of the tumour at T10 had not changed since I-131 therapy in 1999. A technetium-99m bone scan showed a cold spot at the right side of T10, and positron emission tomography (PET) also demonstrated no increased uptake. It meant that the I-131 therapy had a sufficient effect on the T10 tumour and there were no other metastatic lesions. Although there was no cord compression by the tumour as seen on the MRI, the spinal canal was narrow due to ossification of the yellow ligament (OYL), as seen on the CT (Figure 3a), and the cord was severely compressed by the OYL as seen on the MRI (Figure 3b) at the level of T10/11. At first, we diagnosed this as Brown-Sequard syndrome due to the OYL on T10/11, rather than the metastatic tumour of T10 based on imaging studies. To improve her paralysis, an excision of the OYL was performed after laminectomies of T9, 10 and 11 in March 2003. During the surgery, extensive fibrous scar tissue existed around the dural matter (Figure 4). Simultaneously, curettage of the T10 vertebral tumour from the posterior approach to evaluate the viability of the tumour and spinal instrumentation was performed. Histological examination of the excised material e46

(vertebral tumour) showed fibrous tissue and necrotic tissue (total necrosis), and no tumour cells (Figure 5). Her sensory and motor disturbance improved (4-5/5 from 3-4/5) immediately after the operation. Her gait exercise was started 6 days after surgery and her post-operative course was uneventful for a week. However, she gradually felt muscle weakness (2-3/5) in her right leg without any apparent trauma about 10 days after the surgery. In spite of intravenous corticosteroid injections for a few days, motor paralysis in the right leg progressed and was complete 24 days after the surgery. Furthermore, muscle weakness (2-3/5) in her left leg and sensory disturbance in her right leg, which was not seen before the operation, appeared simultaneously. Then, sensory disturbance of her left leg, which had improved immediately after the surgery, progressed rapidly. She also had difficulty with urination. Her symptoms did not improve at all and her paralysis was completed 35 days after the surgery. During the progression of paralysis, an MRI was performed several times and the follow-up MRI showed adequate decompression from the OYL with no other lesions around the spinal cord such as haematoma or tumour. The high intensity area of the spinal cord on T2 weighted images, which was detected before the surgery, The British Journal of Radiology, August 2006

Case report: Radiation myelopathy

(a)

(b)

Figure 3. OYL at the level of T10/11. (a) CT scan. (b) T2 weighted MRI.

was seen to remain at the same size and level without any changes on the MRI (Figure 6).

Discussion From the patient’s clinical and imaging studies, we diagnosed radiation myelopathy developing after I-131 therapy. Although it is difficult to diagnose radiation myelopathy definitively in the absence of pathological evidence from the spinal cord, three criteria for the diagnosis of radiation myelopathy have been established [1]: 1. The spinal cord must have been included in the radiation field. 2. The neurological deficit must correspond to the cord segment that was irradiated.

Figure 4. Intraoperative picture showing extensive fibrous scar on the dura.


3. Metastases or other primary spinal cord lesions must be ruled out. The clinical and imaging features in this case met the three criteria. This patient received I-131 therapy for a metastatic tumour of T10 from 1999 to 2000. An extensive fibrous scar around the dural matter at T10, which was seen in surgery, is usually seen after external irradiation. It proves that the spinal cord was certainly irradiated at this level (criteria 1). The neurological deficit level of Brown-Sequard syndrome in this case corresponded to T10 tumour level (criteria 2). The postoperative MRI in this patient showed no cord compression by the metastatic tumour or the OYL. There was no intradural invading tumour. Furthermore, her clinical findings, clinical course and imaging findings were completely different from those of other spinal cord lesions such as spinal cord infarction, multiple sclerosis, or myelitis. Damage to the cord during this surgery was

Figure 5. Low-power photomicrograph of the specimen from excised materials of T10 tumour. Haematoxylin and eosin 6120. e47

H Murakami, N Kawahara, T Yahata et al

Figure 6. Post-operative MRI. A sagittal T2 weighted image showing high signal intensity on the spinal cord.

not a cause of myelopathy, since her neurological deficit improved immediately after the surgery. Her paralysis progressed although no other causes of paralysis were apparent (criteria 3). In addition to the three criteria, the MRI findings in this case were also consistent with those of radiation myelopathy. The MRI features of radiation myelopathy are characterized as an area of high signal intensity within the spinal cord detected on T2 weighted images [2–4]. In this case, the high intensity area of the spinal cord on T2 weighted images from T5 to T10, which was not seen when I-131 therapy started in 1999, is compatible with findings of radiation myelopathy. Another reason to diagnose radiation myelopathy, not Brown-Sequard syndrome due to the OYL, is the characteristic neurological findings. Despite the spinal cord compression by the OYL in the same degrees bilaterally, on the tumour side (right side) injured Brown-Sequard syndrome was seen. The definite diagnosis in this case must be radiation myelopathy. There are no previous reports describing radiation myelopathy e48

due to I-131 therapy in the literature. We need to recognize the possibility of this serious complication even in the practice of I-131 therapy. Radiation-induced myelopathy is the most serious complication associated with radiotherapy. Many cases have been reported since Ahlbom [5] first published on radiation myelopathy after external radiotherapy in 1941. The latent period between termination of irradiation and onset of neurological symptoms has varied between 3–5.5 months and 30–41 months, but in the majority of cases was 9–20 months [6]. In this case, it was 34 months. Higher radiation doses, larger doses per fraction and previous exposure to radiation could be associated with a higher probability of developing radiation myelopathy [7, 8]. Based on animal and human data, the 5% likelihood of the complication after external radiotherapy at 5 years and 50% likelihood of the complication at 5 years for a 10 cm length of spinal cord are 57– 61 Gy and 68–73 Gy in 2 Gy fractions, respectively [9]. Radiotherapy is often administered to end-stage patients with metastatic spinal tumours. There would be many cases in which the patients died before the appearance of myelopathy. The incidence of myelopathy in long-term surviving patients is estimated to be higher than the incidence reported in the literature. The pathogenesis of radiation myelopathy is reported to involve damage to the vascular endothelium with secondary permeability disorders of the blood–brain barrier [10–13]. Once radiation myelopathy occurs, there is no known effective treatment for radiation myelopathy and the symptoms, in general, are irreversible. Because it is primarily concentrated in thyroid tissue, I-131 can be used in the treatment of thyroid cancer, and lung or bone metastasis of thyroid cancer. Bone metastases typically have little or no response to I-131 therapy [14]. However, some authors have found that I-131 therapy can be used with curative intent in patients with bone metastases [15]. Thyroid cancer cells that have spread to other parts of the body are killed when they absorb I-131, which has a physical half-life of 8.05 days. It decays by high energy gamma photon and particulate emissions (beta particle). The beta particle will deposit the majority of its energy within 2 mm of its site of origin [16]. The patient reported on here received a total of 22.8 GBq of radioiodine. This amount of radioiodine is equivalent to about 20–30 Gy in external radiotherapy and should be within the safety margin for the spinal cord. However, especially in cases with spinal cord compression such as in this case, there is a risk of radiation myelopathy after I-131 therapy even if it is the amount generally used, since the tumour is attached to the spinal cord.

Acknowledgments The authors thank William C Hutton, DSc for his kind criticism and advice.

References 1. Pallis CA, Louis S, Morgan RL. Radiation myelopathy. Brain 1961;84:460–79. 2. Komachi H, Tsuchiya K, Ikeda M, Koike R, Matsunaga T, Ikeda K. Radiation myelopathy: a clinicopathological


Case report: Radiation myelopathy

3.

4.

5.

6.

7.

8.

study with special reference to correlation between MRI findings and neuropathology. J Neurol Sci 1995;132: 228–32. Maddison P, Southern P, Johnson M. Clinical and MRI discordance in a case of delayed radiation myelopathy. J Neurol Neurosurg Psychiatry 2000;69:563–64. Zweig G, Russell EJ. Radiation myelopathy of the cervical spinal cord: MR findings. AJNR Am J Neuroradiol 1990;11: 1188–90. Ahlbom HE. Results of radiotherapy of hypopharyngeal cancer at radiumhemmet, Stockhlom, 1930 to 1939. Acta Radiol (Stockh.) 1941;22:155–71. Michikawa M, Wada Y, Sano M, Uchihara T, Furukawa T, Shibuya H, et al. Radiation myelopathy: significance of gadolinium-DTPA enhancement in the diagnosis. Neuroradiology 1991;33:286–9. Godwin-Austen RB, Howell DA, Worthington B. Observations on radiation myelopathy. Brain 1975;98: 557–68. Phillips TL, Buschke F. Radiation tolerance of the thoracic spinal cord. AJR Am J Roentgenol 1969;105:659–64.


9. Rubin P, Casarett GW. Direction for clinical radiation pathology – the tolerance dose. Front Radiation Ther Oncol 1972;6:1–16. 10. Fogelholm R, Haltia M, Andersson LC. Radiation myelopathy of cervical spinal cord simulating intramedullary neoplasm. J Neurol Neulosurg Psychiatry 1974;37:1177–80. 11. Goldwein JW. Radiation myelopathy; a review. Med Pediat Oncol 1987;15:89–95. 12. Jellinger K, Strum KW. Delayed radiation myelopathy in man. J Neurol Sci 1971;14:389–408. 13. Schultheiss TE, Stephens LC. Permanent radiation myelopathy. Br J Radiol 1992;65:737–53. 14. Scott AM, Larson SM. Tumor imaging and therapy. Radiol Clin North Am 1993;31:859–79. 15. Petrich T, Widjaja A, Musholt TJ, Hofmann M, Brunkhorst T, Ehrenheim C, et al. Outcome after radioiodine therapy in 107 patients with differentiated thyroid carcinoma and initial bone metastases: side-effects and influence of age. Eur J Nucl Med 2001;28:203–8. 16. Bender JM, Dworkin HJ. Iodine-131 as an oncology agent. J Nucl Med Technol 1993;21:140–50.

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CASE REPORT

Upper pole of a duplex kidney mimicking adrenal incidentaloma in 18F-fluoro-2-deoxy-D-glucose positron emission tomography: a pitfall in diagnosis 1

S-C CHAN,

MD,

1

T-C YEN,

MD, PhD

and 2K-K NG,

MD

Departments of 1Nuclear Medicine and 2Diagnostic Radiology, Chang Gung Memorial Hospital, Linkou Medical Center, Taoyuan, Taiwan ABSTRACT. 18F-fluoro-2-deoxy-D-glucose positron emission tomography (FDG-PET) has proved to be valuable in the diagnosis and management of a variety of malignancies, but is still limited in providing detailed anatomical information. According to the literature, an adrenal incidentaloma with high FDG uptake usually indicates malignancy and requires further investigation. However, accurate localization of the adrenal gland in FDG-PET is difficult without the presence of surrounding wellvisualized organs, such as the kidney or liver. If these organs have a congenital anomaly or are altered due to a previous operation, misdiagnosis can occur. We present a case with right partial duplex kidney accompanied by abnormal urine retention in the upper pole, which was misinterpreted as an adrenal incidentaloma in FDG-PET. A subsequent CT scan revealed a normal right adrenal gland, but a right partial duplex kidney. Fusion of the PET and CT images showed that the right adrenal lesion seen in the PET image corresponded to the upper pole of the duplex kidney.

18 F-fluoro-2-deoxy-D-glucose positron emission tomography (FDG-PET) has been proved to be useful in the diagnosis of adrenal incidentaloma. According to the literature, an FDG-avid adrenal lesion is usually malignant, which requires further examination [1, 2]. However, correct localization of the adrenal lesion depends on the surrounding visualized organs, such as the kidney and liver. Misdiagnosis may be inevitable for cases with congenital anomalies. Duplex kidney is the most common congenital anomaly of the urinary tract, characterized by two different renal pelvic and ureteral systems in a single kidney. Occasionally, it is associated with ureteral obstruction [3]. Herein, we present a case with right partial duplex kidney accompanied by abnormal urine retention in the upper pole, which was misinterpreted as an adrenal incidentaloma in FDG-PET.

Case report A 50-year-old male experienced general weakness, fatigue and weight loss for several months. He had received serial medical examinations in another local hospital, but the results were normal. Thus, he was referred to our hospital for further survey. Under the consideration of occult malignancy, we arranged a whole-body FDG-PET scan for him. The scan revealed Address correspondence to: Koon-Kwan Ng, Department of Diagnostic Radiology, Chang Gung Memorial Hospital, 5, Fushing St, Kweishan, 333 Taoyuan, Taiwan.

e50

Received 12 May 2005 Revised 24 August 2005 Accepted 30 August 2005 DOI: 10.1259/bjr/38262574 ’ 2006 The British Institute of Radiology

a lesion with high FDG uptake in the right suprarenal region (standardized uptake value530.4, Figure 1a). To further confirm its relevance to the right kidney, a delayed image was acquired 2 h post-intravenous injection of a diuretic. In the delayed image, the right suprarenal lesion remained unchanged while most of the urine in the right renal pelvis was excreted (Figure 1b). Based on the delayed image result, the patient was presumed to be suffering from a right adrenal malignancy, and an abdominal CT scan was subsequently performed using a 16 multislice scanner. The coronal CT image showed a normal right adrenal gland without any aberrant lesion around it. However, CT urography revealed a right kidney with two separate pelvic and ureter systems, the upper ureter joining the lower one at the proximal portion, indicating a partial duplex kidney (Figure 2). Fusion of CT and PET images using commercially available software (HermesH; Nuclear Diagnostics AB, Hagersten, Sweden) showed that the right adrenal lesion in PET exactly corresponded to the upper pole of this right duplex kidney (Figure 3). Because PET also showed prolonged urine retention in the upper pole, obstructive uropathy was highly suspected. Although CT urography revealed an apparent interruption of the ureter of the upper pole, it could have been resulted from either ureteral obstruction or normal ureteral peristalsis. Thus, other examinations such as intravenous pyelography were suggested for further evaluation. However, this patient decided to follow-up later after in-depth discussion because there were no associated symptoms or signs with regard to obstructive uropathy. The British Journal of Radiology, August 2006

Case report: Upper pole of a duplex kidney mimicking adrenal incidentaloma in FDG-PET

(a)

(b)

Figure 1. (a) FDG-PET scan revealed a large FDG avid lesion in the right suprarenal area (arrow). (b) A delayed-phase scan performed 2 h later after injection of the diuretic showed unchanged suprarenal lesion whilst the activity over lower half of right kidney has markedly reduced; thus, an adrenal incidentaloma was presumed.

Discussion The reported incidence of adrenal incidentaloma identified in conventional radiographic examinations is approximately 5% [4]. Additional examinations are usually necessary to confirm the nature of the lesion

and to exclude the presence of life-threatening malignant disease. In FDG-PET, a normal right adrenal gland is barely visualized [5], whereas an adrenal lesion with intense FDG uptake usually represents malignancy [1, 2, 5] and warrants further aggressive examination. FDG-PET is still limited in several aspects. For example, FDG uptake is not specific to malignancies. A benign lesion such as local inflammation can also demonstrate significantly increased FDG uptake [6]. Also, FDG-PET does not provide adequate structural information, which is important for correlating the functional areas of FDG uptake to an anatomical reference. Even with the aid of the tomographic technique of PET, it is still not easy to differentiate a malignant lesion from the physiological uptake of FDG in the genitourinary tract or bowels. For instance, misinterpretation of physiological uptake of FDG by brown adipose tissue surrounding the adrenal gland as malignant metastasis has been reported [7]. Therefore, in the instance of a FDG-avid adrenal lesion, additional anatomical information from CT/MRI is crucial in order to avoid misdiagnosis.

Figure 2. CT urography showed a right partial duplex

Figure 3. Fusion of the CT and FDG-PET images at the same

kidney with a small upper pole. Interruption of the upper pole ureter could be due to peristalsis or obstruction (arrow).

level showed the right ‘‘adrenal’’ lesion in PET corresponded to the upper pole of the right duplex kidney in CT (arrow).


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S-C Chan, T-C Yen and K-K Ng

In this presented case, accumulation of urine in the upper pole of a duplex kidney was misinterpreted as an adrenal malignancy. Duplex kidney is the most common congenital abnormality, exhibiting a urographic incidence of 2% [3]. Duplex kidney is characterized by two separate renal pelvic and ureter systems. Occasionally, a duplex kidney may coexist with ureter obstruction, vesicoureteric reflex, ectopic ureteric orifice, and ureterocoele. In rare instances, two distinct ureters, indicative of a duplex kidney, can be identified in FDG-PET images [8]. As demonstrated in this case, urine retention within the upper pole of the duplex kidney can be mistaken for an adrenal lesion in a FDG-PET image because the upper pole of the duplex kidney is relatively small and may be associated with ureter obstruction [9]. Although radiological examinations could not fully explain the nature of the obstruction in this case, evidence of prolonged urine retention noted in PET is a strong indication for obstructive uropathy. The reason for prolonged urine retention may be correlated to subclinical mucosal inflammation of the ureter caused by vesicoureteric reflux. Mucosal inflammation can lead to narrowing of the ureteral lumen, which is not detectable using either CT or CT urography. Misinterpretation can be avoided in such circumstances. Since the upper pole consisted of one third of the duplex kidney, the ‘‘pseudo-adrenal’’ lesion in the PET image was always huge. For a huge adrenal malignancy, whether it is a metastasis or primary adrenal tumour, the corresponding primary malignancy or clinical and laboratory endocrine abnormalities should be obvious. In a case without the aforementioned abnormalities, the chance of adrenal malignancy is low and the possibility of a duplex kidney should be considered. Since not every case with duplex kidney is accompanied with ureter obstruction, administration of a diuretic is still useful for further confirmation.

e52

In conclusion, because duplex kidney is the most common congenital abnormality of the urinary tract and may be undetected throughout the whole life, interpretation of any suspicious adrenal lesion in the FDGPET image should be done with caution. Correlation with other clinical information, corresponding CT/MRI findings, or a delayed-phase PET image after administration of the diuretic may avoid misdiagnosis.

References 1. Kumar R, Xiu Y, Yu JQ, Takalkar A, El-Haddad G, Potenta S, et al. 18F-FDG PET in evaluation of adrenal lesions in patients with lung cancer. J Nucl Med 2004;45:2058–62. 2. Maurea S, Klain M, Mainolfi C, Ziviello M, Salvatore M. The diagnostic role of radionuclide imaging in evaluation of patients with nonhypersecreting adrenal masses. J Nucl Med 2001;42:884–92. 3. Privett JT, Jeans WD, Roylance J. The incidence and importance of renal duplication. Clin Radiol 1976;27:521–30. 4. Caplan RH, Strutt PJ, Wickus GG. Subclinical hormone secretion by incidentally discovered adrenal masses. Arch Surg 1994;129:291–6. 5. Bagheri B, Maurer AH, Cone L, Doss M, Adler L. Characterization of the normal adrenal gland with 18F-FDG PET/CT. J Nucl Med 2004;45:1340–3. 6. Rao SK, Caride VJ, Ponn R, Giakovis E, Lee SH. 18F fluorodeoxyglucose positron emission tomography-positive benign adrenal cortical adenoma: imaging features and pathologic correlation. Clin Nucl Med 2004;29:300–2. 7. Reddy MP, Ramaswamy MR. FDG uptake in brown adipose tissue mimicking an adrenal metastasis: source of false-positive interpretation. Clin Nucl Med 2005;30: 257–8. 8. Balan KK, Balan A. Detection of duplex kidney on a wholebody F-18 FDG positron emission tomographic scan. Clin Nucl Med 2003;28:315. 9. Hartman GW, Hodson CJ. The duplex kidney and related abnormalities. Clin Radiol 1969;20:387–400.



CASE REPORT

Stable disease and improved health-related quality of life (HRQoL) following fractionated low dose 131I-metaiodobenzylguanidine (MIBG) therapy in metastatic paediatric paraganglioma: observation on false ‘‘reverse’’ discordance during pre-therapy work up and its implication for patient selection for high dose targeted therapy S BASU,

MBBS (Hons), DRM, DNB

and N NAIR,

MD

Radiation Medicine Centre, Tata Memorial Hospital Annexe, Jerbai Wadia Road, Parel, Bombay 400 012, India ABSTRACT. The incidence of paraganglioma in the paediatric population is exceedingly rare, accounting for ,0.1% of childhood cancers. We report here the response and toxicity profile in a case of malignant paraganglioma which was treated with what is currently perceived as an unconventional and non-standard approach, using three consecutive low doses of 131I-MIBG (a cumulative dose of 11 647.6 MBq). The patient had a stable disease at the end of 42 months follow-up following the first treatment with 131I-MIBG. Excellent symptomatic and hormonal responses were observed. The only adverse effect was mild nausea in the first 24 h after therapy. In addition to the potentially primary end point of radiological and biochemical response measurement, we, in this paper, endeavoured to look for a quality of life evaluation for this form of treatment. Given the rarity of this condition, the experience gained by this therapeutic approach is intriguing from response and toxicity standpoints and may be extrapolated to malignant pheochromocytoma as well. An apparent ‘‘reverse discordance’’ between 131 I-MIBG scintigraphy and 99Tcm-MDP bone scan encountered during pre-therapy work up is also described with possible explanations. This draws attention to an important clinical issue in selecting patients for high dose targeted therapy.

Whole body 131I-MIBG scintigraphy is a relatively simple yet highly accurate procedure that plays a pivotal role in the clinical algorithm of evaluation of neural crest tumours. The fascinating principle of radiopharmaceutical targeting in tumours has been successfully exploited in the application of MIBG in diagnosis and treatment of neuroendocrine tumours. Data comparing the bone scan and MIBG scan in identifying the skeletal lesions in the neural crest tumours are sparse. Few studies [1–3] addressing this issue, are mainly restricted to the cases of neuroblastoma, wherein greater numbers of skeletal lesions were consistently evident on MIBG scintigraphy. There has not been any convincing study comparing the two scintigraphies in documenting skeletal secondaries in pheochromocytoma or paraganglioma. Common experience, however, suggests that radiolabelled MIBG is particularly useful and superior to conventional bone scan in the localization of malignant pheochromocytoma Address correspondence to: Dr Sandip Basu, Radiation Medicine Centre, Bhabha Atomic Research Centre, Tata Memorial Centre Annexe, Jerbai Wadia Road, Parel, Bombay 400 012, India.


Received 18 November 2004 Revised 16 July 2005 Accepted 2 September 2005 DOI: 10.1259/bjr/27645093 ’ 2006 The British Institute of Radiology

at remote skeletal metastases. The authors, in this article, present an unusual finding in a case of malignant paraganglioma where skeletal scintigraphy demonstrated more metastatic sites than diagnostic 131I-MIBG scintigraphy, all of which were subsequently found to be true positive lesions by the subsequent 131I-MIBG posttreatment scan. This apparent ‘‘reverse discordance’’ in diagnostic scan, subsequently clarified by post-therapy scan, stresses the fact that one needs to exert caution before designating the skeletal lesions positive by bone scan but ‘‘negative’’ on the diagnostic MIBG scan as either non-specific or non 131I-MIBG concentrating, even though the MIBG diagnostic scintigraphy may have shown uptake in some of the bone scan positive sites. The patient was treated thrice at an interval of 12 weeks showing an excellent palliative response, assessed by the follow up scans at different time intervals and Functional Assessment of Cancer Therapy General (FACT-G) quality of life questionnaire. This report probably represents the first effort to prospectively evaluate health-related quality of life (HRQoL) in 131I MIBG therapy. e53

S Basu and N Nair

Case report A 13-year-old female presented with a history of recurrent episodes of sudden severe frontal headache, diaphoresis and palpitation, and documentation of high blood pressure recordings during the episodes. The last episode was 15 days previously when her blood pressure (BP) recording was 240/130 mm Hg. On examination, her basal pulse rate was 90 min21 and BP was 160/ 110 mm Hg. Her fundus showed grade II hypertensive changes. An ultrasound as well as a CT of the abdomen revealed a large extra-adrenal mass measuring 6 cm63 cm64 cm between inferior vena cava and aorta adjacent to the hilum of right kidney. The 24 h urinary vanillylmandelic acid (VMA) was 149 mg g21 of creatinine (normal: 1–9 mg g21 of creatinine). The patient commenced phenoxybenzamine with good preoperative control of hypertension, and the entire mass was excised. The histopathology revealed it to be extraadrenal pheochromcytoma. A post-operative CT scan of the abdomen at 3 months confirmed complete excision of the mass. Post-operatively, her BP was normal and she remained asymptomatic for 1 year. Thereafter, she started experiencing episodes of giddiness, frontal headache and pain at the mid dorsal region, palpitation and occasional visual obscurations. Her BP was 190/ 110 mm Hg and there was tenderness over right 8th and 9th ribs and the dorsal spine. She was put on prazosin hydrochloride (5 mg thrice daily) and atenolol (25 mg once daily) for control of hypertension. She required analgesics (a combination of ibuprofen and paracetamol was prescribed) almost daily for the alleviation of pain. A repeat CT scan of the abdomen revealed an ill defined lesion with an enhancing component measuring 4 cm63 cm62.5 cm to the right of abdominal aorta at the level of T11–L2 vertebrae. A whole body skeletal survey with 16 mCi 99Tcm methylene diphosphonate (MDP) demonstrated abnormal focal uptake of tracer in the following (Figure 1):

N N N N N

Medial part of left supraorbital region T4–T6 vertebrae Left shoulder Right 8th rib Right superolateral part of L3 vertebra

A 1 mCi diagnostic 131I-metaiodobenzylguanidine (MIBG) whole body scan (Figure 2a) 48 h post-injection demonstrated a well defined focal area of increased tracer uptake in:

Tcm-MDP bone scintigraphy showing abnormal focal uptake of tracer in the medial part of left supraorbital region, T4–T6 vertebrae, left shoulder, right 8th rib and right superolateral part of L3 vertebra.

N

excised) with skeletal metastases was made. The 24 h urinary VMA was 19.5 mg g21 of creatinine (normal: 1– 9 mg g21 of creatinine). In view of the persistence of hypertension and pain at the metastatic sites, the patient was considered for 131I-MIBG therapy. As a part of the pre-therapy workup protocol a repeat 131I-MIBG whole body scan was performed, which showed focal tracer uptake only at the region of T4–T6 vertebral lesions (Figure 2b). The other skeletal lesions evident in bone scan did not show any tracer uptake. The abdominal tracer uptake earlier was not evident this time (not shown in Figure 2b) consistent with history of complete removal of the mass. Her complete blood counts and

N

Mid thoracic region (more evident in posterior view); and At an area inferomedial to the right lobe of liver.

These uptakes corresponded to the T4–T6 vertebral lesion seen in bone scintigraphy and the mass lesion evident in CT scan abdomen, respectively. Excision of the para-aortic mass along with right 8th rib biopsy was carried out in the same sitting under general anaesthesia. The histopathology of the mass was consistent with the diagnosis of paraganglioma and the rib biopsy was reported as metastasis from malignant paraganglioma. So, a diagnosis of recurrent paraganglioma (which was e54

Figure 1.

99


Case report: MIBG in metastatic paediatric paraganglioma

(a)

(b)

Figure 2. (a) 1 mCi diagnostic 131I-metaiodobenzylguanidine (MIBG) whole body scan 48 h post-injection demonstrated a well defined focal area of increased tracer uptake in midthoracic region (more evident in posterior view) and at an area inferomedial to the right lobe of liver. (b) Pre-therapy repeat 131I-MIBG posterior chest scan showing focal tracer uptake only at the region of T4–T6 vertebral lesions. Rest of the whole body survey did not reveal any abnormal tracer uptake.

serum urea and creatinine levels were all within normal limits and she was treated with 4025.6 MBq (108.8 mCi) 131 I-MIBG by intravenous infusion over 3.5 h. Proper precautions like thyroid blockade and adequate hydration were utilized for thyroid and urinary bladder protection at each diagnostic scan and therapy. Before discharging, routine whole body imaging was undertaken, which showed focal concentration of 131I-MIBG at all the sites seen on skeletal scintigraphy (Figure 3), i.e.:

N N N N N

In the medial part of left supraorbital region T4–T6 vertebrae Left shoulder Right 8th rib Right superolateral part of L3 vertebra

The intensity of the tracer uptake at the individual sites varied. She was treated with 131I-MIBG two times subsequently at an interval of 12 weeks with a dose of 3552 MBq (96 mCi) and 4070 MBq (110 mCi) and showed considerable symptomatic improvement and became normotensive. Her 24 h urinary VMA level was within normal limits at 3 months after the third therapy. All of the therapies were well tolerated, with mild nausea as the only adverse effect in the first 24 h of treatment. The follow-up haematological profiles were within normal limits on all occasions. The total leukocyte count was 7200 mm23, haemoglobin 11.1 g%, total platelet count 180 000 mm23 at the pre-treatment baseline investigation. The haemograms The British Journal of Radiology, August 2006

carried out weekly for 6 weeks following treatment and from during the pre-therapy workup of subsequent therapies did not reveal any appreciable change in the counts at any instant. The cumulative dose of 131I-MIBG used in this case was 11 647.6 MBq (314.8 mCi). She is being assessed periodically, with the most recent at the end of 42 months after the first treatment. The 131IMIBG whole body scan showed tracer uptake only at the region of T4–T6 vertebral lesions and was of the same size. The bone scan showed focal uptake at the same sites, with no new sites revealed. The lesional uptakes were of the same intensity and dimension as seen in the first scan. She was designated as having ‘‘stable disease’’. Considerable symptomatic improvement was documented, with only three to four paroxysmal hypertensive episodes since the first therapy. Since the last therapy, there was no single episode of headache, diaphoresis, palpitation or apprehension. Atenolol was discontinued and the prazosin dosage was reduced to 2.5 mg thrice daily while she was being evaluated for the 2nd therapy. She was weaned off prazosin while she was being evaluated at 3 months after the 3rd therapy. The only symptom was pain at the dorsal vertebrae and left shoulder consistent with known metastatic foci at those sites. However, pain in none of these sites was severe enough to warrant intake of analgesics. She did not have any symptom related absenteeism from school in the last 2 years and her school performance was excellent. Assessment of HRQoL using FACT-G questionnaire (version 4) saw improvement in all the subscales. This e55

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dorsal vertebrae and the left supraorbital region). She had a reduced performance in several components (Table 1). This was probably, in the majority of the components, related to the catecholamine surge characteristic of neuroendocrine tumours and the improvement was consistent with the gradual fall of urinary VMA level to normal. Her subsequent urinary VMA levels were within normal limits on all occasions.

Discussion

Figure 3. Post 131I-MIBG treatment scan showing focal concentration in the medial part of left supraorbital region, T4–T6 vertebrae, left shoulder, right 8th rib and right superolateral part of L3 vertebra.

27-item ordinal scoring [15] based on a five-point scale of categorical response choices, encompasses the four primary domains of quality of life (i.e. physical, social/ family, emotional and functional well being). Several items have reverse anchors, and hence care was taken to code these items by reversing them while scoring those items. On presentation to us, her debility was related mainly to the hypertension related to the catecholamine surge and pain at the metastatic sites (especially at the e56

The nomenclature for paraganglioma is inconsistent. Sometimes the term is used to designate any extraadrenal tumour of the paraganglion system. In other instances, the term is reserved for non-functional tumours of the paraganglion system, with functional tumours referred to as ‘‘extra-adrenal pheochromocytomas’’. Tumours in the head, neck and paravagal region are usually non-functioning. Tumours around the aorta and sympathetic chain and visceral tumours, such as bladder tumours, usually elaborate catecholamines. Usually extra-adrenal pheochromocytomas (i.e. functional paragangliomas) occur in the abdomen along the great vessels, the most common site being the upper periaortic region from the diaphragm to the lower poles of the kidneys [5, 6]. The second most common site is the organs of Zuckercandl at the origin of the inferior mesenteric artery. The present case imparts two important messages and is discussed below with the possible explanations. The basic principle of bone scanning is dependent on local osteoblastic reaction or the bone remodelling response. Before the introduction of radiolabelled MIBG, bone scanning remained the cornerstone of the evaluation of skeletal secondaries in neuroendocrine tumours because it is a highly sensitive, cost effective and non-invasive means to evaluate the whole body skeleton in a single examination. However, several drawbacks of this modality were addressed after the regular use of diagnostic MIBG scanning. The few comparative studies between the two scintigraphies in detection of skeletal metastases from neural crest tumours are mainly restricted to that from neuroblastoma. The study by Shulkin et al in 77 patients of neuroblastoma found nearly twofold greater number of skeletal lesions evident in MIBG scintigraphy [1]. Gordon et al reported about 60% more lesions with MIBG scintigraphy than with bone scanning [2]. In another comparative study between 123I-MIBG and bone scan, more lesions were found with 123I-MIBG scan and there were several false-positive lesions on bone scintigraphy [3]. The reasons put forward for the depiction of more skeletal metastatic lesions in neuroblastoma with MIBG scintigraphy are the following: (a) as the metastatic lesions usually originate from within the bone marrow, skeletal scintigraphy may underestimate early spread. (b) The propensity of neuroblastoma to metastasise in the metaphysis adjacent to the epiphyseal plates (which themselves are sites of increased skeletal tracer uptake); metastatic involvement may be difficult to appreciate. (c) In symmetric metaphyseal lesions, especially with relatively little alteration in skeletal tracer uptake, some lesions escape detection by skeletal The British Journal of Radiology, August 2006

Case report: MIBG in metastatic paediatric paraganglioma Table 1. Summary of the quality of life data at multiple time points Baseline performance

Lack of energy, trouble meeting needs of life, aches or pain, bothered about side effects of treatment, feeling ill, forced to spend time in bed, satisfaction with family communication about illness, sadness, satisfaction about coping with illness, losing hope, nervousness, anxiety about dying, anxiety about condition getting worse, able to work, able to enjoy work or life, sleep, content with the quality of life at that instant: all these had a reduced performance status prior to treatment At the time of work All the components except pain showed an improvement by at least 2 grades. Pain was graded 3 at up before 2nd therapy this time point 3 months following All except pain were graded either normal or one grade below normal. ‘‘Pain’’ was graded as 2 at this 3rd therapy time point, which showed gradual improvement with time

scintigraphy. On the other hand, few cases where more lesions are detected in bone scans are often found to be false positive due to the scanner’s sensitivity to a wide range of pathological processes [3]. Common experience, as well as research [4], has convincingly proven that radiolabelled MIBG is clearly the scintigraphic procedure of choice in the detection of skeletal metastases at various sites as well as local recurrences in malignant pheochromcytoma. The present case of malignant paraganglioma is unique in several aspects. Here the skeletal scintigraphy picked up more true positive secondaries compared with diagnostic MIBG scintigraphy and thereby demonstrating a ‘‘reverse discordance’’, which is contrary to the common experience. The lesions were evident in the post-therapeutic scan, but of varying intensity. The tracer uptake in the post-treatment scan proves that they were all true positive lesions as well as 131I-MIBG concentrating. The term ‘‘reverse discordance’’ is used to emphasise this unusual finding of bone scan identifying more true positive lesions than the MIBG scintigraphy in a malignant paraganglioma. To the best of our knowledge, no such case has been reported previously. One of the important principles of the current version of ‘‘seed and soil’’ theory, originally hypothesised by Stephen Paget more than a century ago, is that the outcome of metastasis depends on multiple interactions (cross-talk) between metastatic cells with homeostatic mechanisms unique to organ microenvironments influencing the biology of cancer growth, angiogenesis and metastasis in several ways. As a tumour specific radiopharmaceutical, 131I-MIBG uptake is more reliable in estimating tumour activity than bone scan, which is an indirect measure of tumour activity, being mainly dependent on local osteoblastic reaction. In the present vignette, the varying intensity of 131I-MIBG concentration in the post-treatment scan (Figure 3) reflects a combination of tumour differentiation and tumour volume at individual sites, the highest uptake being at the T4–T6 vertebrae. The differential tumour cell proliferation might be due to the different local homeostatic factors that the tumour cells can usurp. The additional foci of metastases on post-therapeutic scan not visible on diagnostic images is consistent with the principle of dose responsive functional imaging [10] with a radiopharmaceutical and is mainly due to two reasons: higher photon flux due to higher activity and delayed imaging allowing better non-target tissue clearance after high dose therapy. Despite the low dose possible due to the relatively unfavourable characteristics of 131I, it continues to be the The British Journal of Radiology, August 2006

most commonly used radiolabel in many institutions, because of its low cost and wide availability. 123I MIBG yields superior image quality due to higher count density resulting from its more favourable characteristics, allowing higher dose administration. It also permits SPECT imaging. However, both are excellent agents for imaging neuroendocrine tumours. In direct comparison between the two isotopes, one study showed the same number of lesions revealed by both, while another showed more lesions identified with 123I-MIBG [7, 8]. The present case opens up an important clinical issue to be addressed: When 123I radiolabel is not available, will a negative diagnostic 131I-MIBG scan, which is often the only yardstick in several centres for selecting patients for high dose 131I-MIBG therapy, actually be false negative as in this instance and thus deny therapy to the deserving patients? Or, should one disregard the results of diagnostic scans in the presence of an elevated biochemical marker (plasma or urinary catecholamines and urinary metabolites of catecholamines or plasma-free normetanephrine and metanephrine) and immediately treat with a high dose of 131I-MIBG and look for lesional uptake in the post-therapy scan? This, however, has the possible drawback of the administration of unnecessary therapy doses in a number of cases, as no other alternative appears available at present. Obviously, this is a dilemma on which a consensus needs to be built and it is hoped that this case presentation will fuel that process. We feel the philosophy, the intent and the protocol of 131 I-MIBG therapy needs to be re-explored in different clinical settings and investigated accordingly. The treating nuclear medicine physician should clearly set before him the goal of treatment in each individual case scenario before he decides upon the individual dose, the interval between therapies and the number of therapies. The contexts differ between tumour types in the neuroendocrine malignancies, mainly due to the varying tumour biology, and even within the same malignancy from case to case. While paragangliomas or pheochromoctomas in general grow slowly, the neuroblastomas usually behave aggressively with more than 50% of the patients presenting with metastatic disease at the time of diagnosis. The treatment of an inoperable neuroblastoma in the neoadjuvant setting is aimed at making the mass operable at the earliest opportunity with an aggressive therapy, while a relatively slow growing multiple metastatic paraganglioma or a pheochromocytoma as in the present case should be treated with a palliative intent with an aim to stabilize the disease. Continued tumour cell proliferation in a fast e57

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growing malignancy at very low dose rates can contribute to the reduced effectiveness of low dose rate radiation. This stems from the fact that tumours with poor repair capabilities usually exhibit less of a dose rate effect than tumours with good repair capabilities. This means, theoretically, a treatment with a curative intent or with an aim to make an inoperable mass operable will require a small number of treatment cycles with a high amount of radiopharmaceutical, whereas a case with multiple metastases from a slow growing tumour should preferably be dealt with repetitive treatments with smaller doses. The current convention for treating neural crest tumours is to adopt a standardized dose ranging from 7.4 GBq (200 mCi) to 11.1 GBq of 131I-MIBG administered by slow intravenous infusion [9]. There is, at present, no clear guideline regarding the individual and cumulative dose to be administered for a particular tumour and therefore depends upon the patient profile, institutional policy and the treating physician. The role of repetitive treatments with lower doses is not well defined in the case of 131I-MIBG therapy. To date, there has not been any head-to-head prospective comparison to investigate whether better responses can be obtained with multiple lower doses than with single larger doses. A number of studies [11–14] addressing the low-dose-rate (LDR) radiobiology of targeted radiopharmaceutical therapy in radioimmunotherapy of lymphoma have been published. The unproportionally high clinical responsiveness sometimes observed after radioimmunoconjugate therapy has been ascribed to various factors, e.g. cytotoxic enhancement or sensitization to protracted courses of low-dose-rate radiation exposure, radiationassociated apoptosis [11, 12] etc. It is well imaginable that the LDR effect can be extrapolated to the treatment of neural crest tumours. In addition, treatment-related toxicity can be reduced with repetitive lower doses. The authors, here, used lower levels of activity of 131IMIBG than previously reported. If dose kg21 of body weight is considered, the use of around 2.5 mCi kg21 of 131 I MIBG each time is much lower compared with the 10–18 mCi kg21 currently used in neuroblastoma therapy. In advanced stage cases, the principal aim of 131IMIBG therapy is symptom palliation and tumour function reduction as well as at tumour arrest and thereby allowing prolonged survival and good quality of life. The excellent patient response and negligible toxicity with this therapeutic approach subserved the above purpose and warrants further study in larger number of patients. Although not used in this case, we conjecture that the interval between the therapies can be reduced using this approach but requires critical examination. As

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total eradication is usually not possible and hence not the intent of therapy in the multiple metastatic advanced cases, this strategy will allow the treating physician to reserve the dose for treatment of future recurrences.

References 1. Shulkin BL, Shapiro B, Hutchinson RJ. I-131-metaiodobenzylguanidine and bone scintigraphy for the detection of neuroblastoma. J Nucl Med 1992;33:1735–40. 2. Gordon I, Peters AM, Gutman A, et al. Skeletal assessment in neuroblastoma--- the pitfalls of iodine-123-MIBG scan. J Nucl Med 1990;31:129–34. 3. Hadj-Djilani NL, Lebtahi NE, Delaloye AB, et al. Diagnosis and follow up of neuroblastoma by means of iodine-123 metaiodobenzylguanidine scintigraphy and bone scan, and the influence of histology. Eur J Nucl Med 1995;22:322–9. 4. Shapiro B, Sisson JC, Lloyd R, et al. Malignant pheochromocytoma: clinical, biochemical and scintigraphic characterization. Clin Endocrinol (Oxf) 1984;20:189–203. 5. Ulchaker JC, Goldfarb DA, Bravo EL, et al. Successful outcomes in pheochromocytoma surgery in the modern era. J Urol 1999;161:764–7. 6. Whalen RK, Althausen AF, Daniels GH. Extra-adrenal pheochromocytoma. J Urol 1992;147:1–10. 7. Gelfand MJ. I-123-MIBG and I-131-MIBG imaging in children with neuroblastoma. J Nucl Med 1996;37:35P. 8. Simon A, Hoefnagel CA, deKraker J. 123I MIBG or 131I MIBG for imaging of neuroblastoma. Eur J Nucl Med 1992;19:589. 9. Hoefnagel CA, Lewington VJ. MIBG therapy. In: Nuclear medicine in clinical diagnosis and treatment. Volume 2. 2nd edition. Churchill Livingstone, 1998. 10. Hickeson P, Charron M, Maris J, et al. Biodistribution of post-therapeutic versus diagnostic 131I-MIBG scans in children with neuroblastoma. Pediatr Blood Cancer 2003;42:268–74. 11. Murtha AD. Radiobiology of low-dose-rate radiation relevant to radioimmunotherapy. Cancer Biother Radiopharm 2000;15:7–14. 12. Macklis RM, Beresford BA, Humm JL. Radiobiologic studies of low-dose-rate 90Y-lymphoma therapy. Cancer 1994;73(3 Suppl.):966–73. 13. Macklis RM, Lin JY, Beresford B, et al. Cellular kinetics, dosimetry, and radiobiology of alpha-particle radioimmunotherapy: induction of apoptosis. Radiat Res 1992;130:220–6. 14. Hernandez MC, Knox SJ. Radiobiology of radioimmunotherapy: targeting CD20 B-cell antigen in non-Hodgkin’s lymphoma. Int J Radiat Oncol Biol Phys 2004;59:1274–87. 15. Cella DF, Tulsky DS, Gray G, et al. The functional assessment of cancer therapy scale: development and validation of the general measure. J Clin Oncol 1993;11:570–9.



CASE REPORT

MRI of a benign chondroid syringoma 1

U KERIMOGLU,

MD,

1

U AYDINGOZ,

MD,

2

O OZKAYA,

MD,

3

A E AKSU,

MD

and 1F B ERGEN,

MD

Departments of 1Radiology, 2Pathology, and 3Plastic and Reconstructive Surgery, Hacettepe University School of Medicine, 06100 Ankara, Turkey ABSTRACT. Chondroid syringoma, also known as mixed tumour of the skin, is a relatively rare, usually benign tumour. A few malignant cases, especially in the lower extremities, have been published, but most of them behave in a benign fashion. A case of a middle-aged woman with surgically proved chondroid syringoma in the pretibial region is presented in this report with MRI and histological findings. Despite rapid growth over a short period of time and a location reportedly associated with malignancy, the histological features were benign.


Chondroid syringoma is a rare, non-melanoma type of skin tumour that originates from the sweat glands, with reported incidence rates varying from 0.01% to 0.1% [1]. It demonstrates sweat gland features with a cartilage-like stroma [1]. Although its histological findings are well described, the radiological features of subcutaneous chondroid syringoma have been reported only on two cases [2, 3]. We report here MRI findings of a case with histologically proven subcutaneous chondroid syringoma.

Case report A 53-year-old woman presented with a nodular subcutaneous mass, which had grown slowly during the last 4 years at the proximal right pretibial region. It was painless and mobile, but had grown rapidly and started to become firm and itchy with a purplish overlying skin in the last 4 months (Figure 1). MRI was performed in a 0.5-T unit (Philips Gyroscan, Best, The Netherlands) with a dedicated extremity (knee) coil. Transverse T2 weighted (repetition time (TR): 4872 ms, echo time (TE): 100 ms, field of view (FOV): 21 cm), sagittal short tau inversion recovery (STIR) (TR: 1400 ms, TE: 30 ms, inversion time (TI): 120 ms, FOV: 22 cm), preand post-contrast sagittal T1 weighted (TR: 500 ms, TE: 20 ms, FOV: 22 cm), and transverse fat-saturated T1 weighted (TR: 525 ms, TE: 19 ms, FOV: 23 cm) MR images (Figure 2) revealed a 3 cm 6 2.7 cm 6 2 cm, well-circumscribed, lobular mass that enhanced diffusely after intravenous gadolinium administration (Figure 2c). The lesion demonstrated low signal intensity with a hyperintense focal nodular area anteriorly on T1 weighted images (Figure 2a), heterogeneous intensity with cystic areas on T2 weighted images and diffusely high signal intensity on STIR images (Figure 2b). No oedema was seen around the lesion. A wide excisional biopsy was performed and, on histological examination, the The British Journal of Radiology, August 2006

tumour was confined to the dermis and the subcutaneous tissue. It comprised multiple lobular structures and circumscribed by a fibrous capsule. The epithelial component was mixed with cartilaginous, fibrous and myxoid stroma. The epithelial cells were arranged as solid masses or formed tubular structures where two rows of cuboidal cells rimmed the tubular luminal surface (Figure 3). There were no atypical mitoses or necrosis. Mitoses were rare. The surgical bed was clear of any tumoural cells.

Discussion Skin tumours are classified as melanoma and nonmelanoma types [2]. Chondroid syringoma, a

Figure 1. An exophytic subcutaneous lesion (arrow) is seen in the lateral aspect of pretibial region.

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U Kerimoglu, U Aydingoz, O Ozkaya et al

(a)

(b)

(c) Figure 2. (a) Sagittal T1 weighted image reveals a well-circumscribed hypointense lesion associated with a hyperintense focus (arrow) in its anterior aspect. (b) Sagittal short tau inversion recovery (STIR) image demonstrates a homogeneously hyperintense mass (arrow) featuring hypointense septa. (c) Transverse T1 weighted image with fat saturation after intravenous contrast administration displays a heterogeneous diffuse enhancement of the lesion (arrow).

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Case report: Benign chondroid syringoma

Figure 3. Photomicrograph shows a nodular neoplasm surrounded by a fibrous capsule (long white arrow), an epithelial component mixed with a basophilic cartilaginous component (short white arrow), and glandular structures (black arrow) (haematoxylin and eosin, 640).

non-melanoma type of skin tumour, is so rare that it is not in the usual list of radiological differential diagnoses of soft tissue tumours and is being diagnosed histologically [1, 2, 4]. Most chondroid syringomas are benign with a slow progression over years [5, 6]. Chondroid syringomas usually present as a solitary, asymptomatic subcutaneous nodule, mainly seen in males. The malignant tumours which are seen mostly in extremities and torso are also slowly growing nodules, but have a higher incidence in females like our patient [4, 7]. They infiltrate regional lymph nodes and metastasise to bones and visceral organs [1]. Such infiltration and metastases were not present in our case. Malignant lesions are usually larger than 3 cm, but there are also published benign cases that measure larger than 4 cm [5]. Rapid growth over a short time, as in our patient, is characteristic for malignant chondroid syringomas [4, 5, 7]. Interestingly, the mass in our patient was histologically benign. Definitive diagnosis of chondroid syringomas is established with histological examination after total excision [1, 5]. Histological diagnosis is usually straightforward, although they can be misdiagnosed as dermoid, sebaceous cysts, neurofibromas and dermatofibromas, and basal cell carcinomas and squamous cell carcinomas [1, 2, 5]. Radiological features of chondroid syringomas are not as suggestive as the histological findings. Both of the two reported cases of chondroid syringomas in the literature, where radiological features were mentioned, had undergone MRI and one had additional radiographs and CT [2, 3]. One of the cases had a superficial chondroid syringoma like our patient [2], while the other had a deep soft tissue mass in the ankle [3]. The former, a palmar tumour, displayed intermediate signal intensity on proton-weighted MR images and high signal intensity with a heterogeneous appearance on T2 weighted images, associated with mainly high signal intensity


interspersed with regions of intermediate signal on STIR sequence [2]; however, there was no mention of intravenous contrast administration for this patient. The latter case presented as a subfascial multilobulated ankle lesion, which demonstrated a hyperintense area displaying hypointense septa on T2 weighted images and inhomogeneous enhancement after intravenous contrast administration [3]. Obviously, these findings are not specific and can be seen in any kind of soft tissue tumour. Although MRI findings in our case resembled some of the features described in these patients, the presence of a focus of increased signal intensity on pre-contrast T1 weighted images was unique in our patient. This focus suggested the presence of melanin or haemorrhage within the lesion and a tentative diagnosis of a soft tissue sarcoma with melanocytic differentiation was made. Histological examination, however, did not disclose the presence of either melanin or a focus of haemorrhage. MRI helped to visualize the borders and the extension of the lesion in our patient. Although the findings did not specify the diagnosis of chondroid syringoma, MRI is very helpful in visualizing the invasion of the tumour and the involvement of the deep structures [2]. Although, in retrospect, our case has presented with the suspicious findings for malignant behaviour such as tumour involvement of the lower extremity (pretibial region) and rapid growth, there was no histological evidence of vascular space invasion and lymph node infiltration. Still, we believe the gender of our patient and the location of the tumour necessitate the follow up of the case for an increased risk of recurrence and metastasis. In conclusion, chondroid syringoma needs to be considered in the differential diagnosis of the lower extremity soft tissue masses despite its rare incidence. The features of our patient suggest that rapid growth over a short time does not necessarily imply a malignant histology. As is the case with many soft tissue tumours, unfortunately, MRI findings are not characteristic.

References 1. Yavuzer R, Basterzi Y, Sari A, Bir F, Sezer C. Chondroid syringoma: a diagnosis more frequent than expected. Dermatol Surg 2003;29:179–81. 2. Nicolaou S, Dubec JJ, Munk PL, O’Connell JX, Lee MJ. Malignant chondroid syringoma of the skin: magnetic resonance imaging features. Australas Radiol 2001;45:240–3. 3. Adachi T, Oda Y, Sakamoto A, Saito T, Tamiya S, Hachitanda Y, et al. Mixed tumor of deep soft tissue. Pathol Intern 2003;53:35–9. 4. Sheikh SS, Pennanen M, Montgomery E. Benign chondroid syringoma: report of a case clinically mimicking a malignant neoplasm. J Surg Oncol 2000;73:228–30. 5. Sungur N, Uysal A, Gumus M, Kocer U. An unusual chondroid syringoma. Dermatol Surg 2003;29:977–9. 6. Terrill RQ, Groves RJ, Cohen BM. Two cases of chondroid syringoma of the hand. J Hand Surg (Am) 1987;12:1094–7. 7. Barnett MD, Wallack MK, Zuretti A, Mesai L, Emery RS, Berson AM. Recurrent malignant chondroid syringoma of the foot: a case report and review of the literature. Am J Clin Oncol 2000;23:227–32.

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CASE REPORT

Endovascular repair of a recurrent aortocaval fistula and anastamotic false aneurysm 1

S PATHAK, MBBS, MRCSEd, 2S W YUSUF, DM, FRCS, 1T N DOYLE, P A E HURST, MS, FRCS and 3C DAVIDSON, MB BChir, FRCP, FESC

FRCR,

1

I A FRANCIS,

FRCS, FRCR,

2

Departments of 1Radiology, 2Vascular Surgery and 3Cardiology, Brighton and Sussex University Hospitals NHS Trust, The Royal Sussex County Hospital, Eastern Road, Brighton BN2 5BE, UK ABSTRACT. Aortocaval fistula (ACF) and false aneurysm are a recognized complication of open abdominal aortic aneurysm (AAA) repair. Untreated they are often fatal. However, open surgical repair of this complication is associated with a high operative mortality and a significant complication rate. Endovascular management using a stentgraft to exclude the false aneurysm and fistula is a technically feasible alternative and confers many advantages over open repair by virtue of its minimally invasive nature. We report the endovascular management of this rare but serious complication of open AAA repair.

Received 3 September 2004 Revised 26 April 2005 Accepted 2 September 2005 DOI: 10.1259/bjr/27172509 ’ 2006 The British Institute of Radiology

Aortocaval fistula (ACF) and false aneurysm are an uncommon complication of open abdominal aortic aneurysm (AAA) repair. Most occur as a complication of aortic graft infection and their prevalence is estimated at less than 1% [1]. Untreated they are often fatal. Most patients are elderly with decreased physiological reserve and multiple co-morbidities, this is reflected in the high complication rate of open repair with a perioperative mortality of up to 90% [2]. Endovascular repair using a stent-graft is able to exclude both the false aneurysm and the abnormal fistula in a single procedure. Endovascular repair of anastomotic aneurysms following open surgery has been reported [3]. We describe the endovascular management of a recurrent ACF with false aneurysm.

ACF. The original graft was an aorto-bifemoral graft with a short body. The stent-graft was designed to cover the segment from immediately below the renal arteries to the graft bifurcation. Intra-arterial access was obtained via the left limb of the aorto-bifemoral graft. A 7 F angiocatheter was introduced via the right brachial artery. The endovascular stent delivery system was introduced over a 0.350 Lunderquist wire and deployed at the site of the false aneurysm and ACF, with its covered segment below the origin of the renal arteries. Completion angiogram showed no further communication of the aorta with the IVC or escape of contrast into the false aneurysm.

Materials and methods A 75-year-old man with a previous history of AAA repair presented with an abdominal mass, fluid retention and shortness of breath. Clinical examination revealed cardiac failure and a pulsatile mass in the abdomen with a ‘‘machinery-like’’ murmur. All peripheral pulses were present. Multidetector CT scan showed a 14 cm false aneurysm adjacent to the proximal graft anastamosis and a fistula between the false aneurysm and the inferior vena cava (IVC) (Figure 1). A transaxillary arteriogram showed contrast extravasating from the posterior aspect of the proximal graft anastamosis into the false aneurysm. A repair was undertaken using an endovascular custom made stent-graft (William Cook, Bjaeverskov, Denmark) to exclude both the false aneurysm and the Address correspondence to: S W Yusuf, Consultant Vascular Surgeon, The Royal Sussex County Hospital, Eastern Road, Brighton BN2 5BE, UK.

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Figure 1. False aneurysm with aortocaval fistula. The British Journal of Radiology, August 2006

Case report: Aortocaval fistula and anastamotic false aneurysm repair

Results The immediate post-procedure period was complicated by clinical signs of consumption coagulopathy, with a falling platelet count and abnormal clotting tests. The patient spent 9 days in ITU but made a satisfactory recovery. A CT scan was performed before discharge which showed no evidence of contrast extravasation and a successful exclusion of the false aneurysm and ACF. The patient was discharged on day 24 on long-term antibiotics. 6 months later, the patient required an axillofemoral bypass graft to deal with a false aneurysm which had developed at the distal anastamosis of the original graft at the right groin; this graft thrombosed at 3 months leading to an above knee amputation. Subsequent CT scan at 18 months demonstrated satisfactory appearances of the stent-graft with continuing exclusion and shrinkage of the false aneurysm (Figure 2). At the same time, a whole body white cell scan did not show any increased uptake of tracer around the stent-graft.

Discussion The endovascular management of the complications of open AAA is fast gaining recognition as a viable alternative to surgical repair. The use of endovascular stents to exclude anastomotic false aneurysms [3] and aortoenteric fistulae is well documented [4], but to our knowledge there is only one report which describes the emergency endovascular management of recurrent ACF [5]. ACF and false aneurysms are a rare complication of aortic bypass and are commonly caused by aortic graft anastomotic breakdown as a result of mechanical stresses and infection. Traditionally, treatment has been by open surgical repair involving graft removal with in situ replacement with femoral veins or extra-anatomical bypass. However, most of these patients are elderly with multiple co-morbidities. This is reflected in the poor outcome and high complication rate of open repair which is associated with an operative mortality rate of 25–90% [2]. The minimally invasive nature of an endovascular approach places less demand on the patients’ physiological reserve and achieves successful exclusion of the false aneurysm and fistula in a single procedure. Care must be taken, because of the introduction of a synthetic graft into a potentially infected area


Figure 2. Exclusion of false aneurysm and aortocaval fistula with the stent-graft.

and suppression of infection is of the utmost importance in maintaining the long term viability of the new graft. Nevertheless, endovascular repair of this condition provides a less invasive alternative to open surgical management.

References 1. Bednarkiewicz M, Pretre R, Kalangos A, Khatchatourian G, Bruschweiler I, Faidutti B. Aortocaval fistula associated with abdominal aortic aneurysm: a diagnostic challenge. Ann Vasc Surg 1997;11:464–6. 2. Orton DF, LeVeen RF, Saigh JA, Culp WC, Fidler JL, Lynch TJ, et al. Aortic prosthetic graft infections: radiologic manifestations and implications for management. Radiographics 2000;20:977–93. 3. Liewald F, Kapfer X, Gorich J, Halter G, Tomczak R, Scharrer-Pamler R. Endograft treatment of anastomotic aneurysms following conventional open surgery for infrarenal aortic aneurysms. Eur J Vasc Endovasc Surg 2001;21: 46–50. 4. Burks JA Jr, Faries PL, Gravereaux EC, Hollier LH, Marin ML. Endovascular repair of bleeding aortoenteric fistulas: a 5-year experience. J Vasc Surg 2001;34:1055–9. 5. Gandini R, Ippoliti A, Pampana E, Marchetti AA, Pistolese GR, Simonetti G. Emergency endograft placement for recurrent aortocaval fistula after conventional AAA repair. J Endovasc Ther 2002;9:208–11.

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CASE REPORT

Intraosseous haemangioma of the proximal femur: imaging findings 1

A CHAWLA,

DMRD,

2

M SINGRAKHIA,

MS,

1

M MAHESHWARI,

MD,

2

N MODI,

MS

and 1H PARMAR,

MD

Departments of 1Radiology and 2Orthopaedics, King Edward VII Memorial Hospital, Parel, Mumbai - 400012, India ABSTRACT. Haemangiomas of bone are uncommon lesions, accounting for approximately 1% of all primary bone tumours. The most frequent sites of involvement are the calvaria and the vertebral column. When haemangiomas involve long tubular bones, they are usually found in the diaphysis or metadiaphysis. Juxta-articular or epiphyseal location for a long bone haemangioma is rare. We present the imaging findings in a case of a histopathologically proven juxta-articular intraosseous haemangioma of the proximal femur. We believe ours is the first report of a haemangioma involving the proximal end of the femur.

Case report A 38-year-old man presented with gradually increasing right hip pain over 4 months. There was no evidence of previous trauma or fever. There was no history of other medical or surgical illness. On examination, there was restricted hip joint movements and tenderness. Baseline haematological examination was normal. A plain radiograph of the hip showed a 4 cm 6 3 cm osteolytic lesion involving the right femoral head, extending into the femoral neck. The lesion had a narrow zone of transition with an unremarkable matrix with focal expansion on its lateral aspect (Figure 1). The soft tissues, contralateral hip joint and remaining pelvis were normal. MRI was performed on a 0.2 Tesla open MR system (Magnetom Open Viva; Siemens Medical Systems, Erlangen, Germany), with images obtained in axial, coronal and sagittal planes. Spin echo (SE) images were obtained with repetition time/time to echo/ number of acquisition (TR/TE/NA) of 528/26/2 for T1 weighted images (WI) and 2840/102/3 for T2 WI with 4 mm slice thickness and 1 mm intersection gap. On T1 WI, the lesion was isointense (to muscle) (Figure 2a). It was markedly hyperintense on T2 WI and short tau inversion recovery (STIR) images (Figure 2b). There were no fluid–fluid levels within the lesion. Minimal joint effusion and bone marrow oedema in the right acetabulum were also noted (Figure 2b). A bone scan revealed a solitary focus of mild increased uptake in the right femoral head. Based on the patient’s age and imaging features, a differential diagnosis of a giant cell tumour (GCT), aneurysmal bone cyst (ABC) and plasmacytoma were considered. While excision biopsy could have been performed for this lesion, we decided to perform excision of the tumour Address correspondence to: Dr Mohit Maheshwari, 10, Ava Road, #09-06 Ava Towers, Singapore 329949.

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Received 10 May 2004 Revised 9 May 2005 Accepted 2 September 2005 DOI: 10.1259/bjr/53131368 ’ 2006 The British Institute of Radiology

followed by an uncemented total hip arthoplasty. On gross examination, the mass appeared as a vascular irregular growth arising from the superolateral quadrant of the femoral head. Frozen section pathology showed regular patterns of mononuclear cells with multiple vessels, suggestive of a benign bone tumour. Histopathology revealed multiple thin-walled vascular spaces with endothelial lining of cells and stuffed with red blood cells. Specks of calcification and border sclerosis were also noted (Figure 3). There was no evidence of malignancy. The final diagnosis was of a haemangioma of the head of femur. On 6 month followup, the patient was asymptomatic with no radiological evidence of recurrence.

Figure 1. Frontal radiograph of both hip joints demonstrating an osteolytic lesion in the head of the right femur with a narrow zone of transition. The lesion is causing minor expansion laterally. The British Journal of Radiology, August 2006

Case report: Intraosseous haemangioma of the proximal femur

(a)

(b)

Figure 2. (a) T1 weighted imaging axial image at the level of the femoral head shows an isointense (to muscle) lobulated mass lesion involving the femoral head. (b) Short tau inversion recovery (STIR) coronal image at the level of the femoral head shows this lesion to be markedly hyperintense. The right acetabulum also shows bone marrow oedema.

Discussion

Figure 3. Histopathology of the excised specimen shows collections of fine thin walled vessels stuffed with red blood cells. Specks of calcification and borders of sclerosis are also seen.


Osseous haemangioma accounts for less than 1% of all primary bone tumours [1–3] and most frequently involve the bones of the axial skeleton. The peripheral, long, tubular bones and the ribs are the least affected sites [1, 2, 5]. Haemangiomas of the tubular bones can be broadly divided as medullary, periosteal and intracortical. The medullary type is further subdivided into diaphyseal (48%), metadiaphyseal (30%), metaphyseal (12%), metaepiphyseal (4%), epimetadiaphyseal (3%) and epiphyseal (1%) [6]. Epiphyseal or juxtarticular location is rare with only two cases reported in the literature; neither of these involved the femur [4, 5]. Haemangiomas involving the axial skeleton are commonly asymptomatic [3] and may remain undetected for long time. They are often discovered incidentally during a radiographic study [3], whereas haemangiomas of the appendicular skeleton are often symptomatic. In one series, peripheral haemangioma were symptomatic in 91% of cases, out of which 37% had pain, 7% had swelling and 7% had restriction of function due to pathological fracture [6]. The rate of pathological fracture depends upon the type of bone involved and whether this bone is weight-bearing. In our case the tumour was located in the weight-bearing posterolateral aspect of the femoral head and was at risk of pathological fracture. The histopathological diagnosis of a haemangioma is not difficult in an en bloc specimen. However, diagnosis from biopsy or tissue curettage is challenging to the pathologist. This is due to the destructive nature of these procedures, disrupting thin-walled blood vessels and resulting in histological sections showing non-diagnostic empty spaces with scattered bone trabeculae. This causes hindrance in the pre-operative diagnosis of haemangioma by any invasive procedure [6]. Histologically, haemangiomas can be classified as cavernous, capillary, venous or mixed, depending on the type of vascular involvement [6]. Cavernous haemangioma is the most e65

A Chawla, M Singrakhia, M Maheshwari et al

common type in the peripheral bones and accounts for up to 50% of all cases reported [6]. They are predominant in the medullary and the intracortical portion of the bone and are not yet reported in the subarticular epiphyseal bone. Pure capillary haemangioma accounts for 10% of all these types as reported in the literature [7, 8]. The other varieties are rare. Intraosseous haemangiomas may have a variety of radiological appearances. Due to the lack of consistent features and rarity, it is often difficult to radiologically diagnose these lesions pre-operatively [6]. On radiographs, haemangiomas may show a coarse loculated, sunburst, moth-eaten, and soap bubble appearance due to expansive proliferation of engorged vessels and thickened, remodelled bone trabeculae [7]. An expansile osteolytic appearance is their least common presentation [9]. On MRI, haemangiomas show a variable appearance and may demonstrate low, intermediate or high signal intensity on T1 WI. Haemangiomas may sometimes appear hyperintense on T1 WI because of their fat content. They are usually hyperintense on T2 WI or STIR [10]. Hyperintensity seen on T2 WI is due to free water in stagnant blood within the haemangioma. On STIR images, these lesions appear markedly intense due to suppression of normal bone marrow. Haemangiomas can show marked to minimal or no enhancement after contrast administration [11]. In our case, the lesion was hypointense on T1 WI and markedly hyperintense on T2 WI and STIR images. The differential diagnosis for a long bone haemangioma in juxta-articular location includes GCT, ABC and plasmacytoma. Other rare possibilities are metastasis and a brown tumour. GCTs predominately arise in long tubular bones [12], with their epicentre in the epiphysis. The radiographic signs of GCT include a geographical radiolucency without internal mineralization, thinning of the cortex and eccentric location in the epiphysis of long bones. ABCs are common in a younger age group with 80% occurring in patients under 20 years of age [10]. Most lesions are metaphyseal; however, they can extend to the epiphyseal end after the growth plate is closed [13]. On MRI, ABCs tend to be well circumscribed but heterogeneous. They often show fluid–fluid levels [14, 15]. Plasmacytomas have a typical radiographic appearance of a geographical radiolucent lesion, which is highly expansile with a soap bubble internal architecture [13]. Metastases are commonly multiple, with solitary metastatic lesions occurring in only 10% of patients [16]. An osteolytic appearance is the most common radiographic feature, present in 75% of all metastatic lesions [13]. Brown tumours are usually central, slightly expansile lightly septated lesions, mimicking destructive neoplasms. They are associated with other features of hyperparathyroidism [13]. In our case the lesion was solitary, with no clinical suspicion suggesting metastasis or a brown tumour.

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In conclusion, haemangiomas can present with bizarre imaging appearances and should be included in the differential diagnosis of an osteolytic and expansile lesion of the epiphysis, especially in asymptomatic or mildly symptomatic patients. However, the diagnosis can only be confirmed by histopathological examination. If correctly diagnosed, such tumours in a non-weight bearing region can be managed conservatively, as with spinal haemangiomas [17]. However, symptomatic or aggressive lesions should be treated as a GCT or an ABC.

References 1. Mulder JD, Kroon HM, Schutte HE, Taconis WK. Radiologic atlas of bone tumors. Amsterdam: Elsevier, 1993:506–7. 2. Unni KK. Dahlin’s bone tumors: general aspects and data on 11,087 cases, 5th edn. Philadelphia, PA: LippincottRaven,1996:306–7. 3. Ogose A, Hotta T, Mortio T, Takizawa T, Ohsawa H, Hirata Y. Solitary osseous hemangioma outside the spinal and craniofacial bones. Arch Orthop Trauma Surg 2000;120: 262–6. 4. Mirra JM. Bone tumors: clinical, radiologic, and pathologic correlations. Philadelphia, PA: Lea & Febiger, 1989:1338–81. 5. Yamamoto T, Kurosaka M, Mizuno K. Juxta-articular hemangioma of long bone. Skeletal Radiol 2000;29:535–7. 6. Kaleem Z, Kyriakos M, Totty WG. Solitary skeletal hemangioma of the extremities. Skeletal Radiol 2000;29: 502–13. 7. Dorfman HD, Czerniak B. Vascular lesions. In: Bone tumors. St. Louis, MO: Mosby, 1998:729–814. 8. Wold LE, Swee RG, Sim FH. Vascular lesions of bone. Pathol Ann (Pt 2) 1985;20:101–37. 9. Matsumoto K, Ishizawa M, Okabe H, Taniguchi I. Hemangioma of bone arising in the ulna: imaging findings with emphasis on MR. Skeletal Radiol 2000;29:231–4. 10. Stoller DW. Bone and soft tissue tumours: In: MRI in orthopedics and sports medicine, 2nd edn. Lippincot Williams and Wilkins, 1997;1231–339. 11. Sung MS, Kim YS, Resnick D. Epithelioid hemangioma of bone. Skeletal Radiol 2000;29:530–4. 12. Breitenseher M, Dominkus M, Scharitzer M, Lechner M, Trieb K, Imhof H, et al. Diagnostic imaging of giant cell tumors. Radiologe 2001;41:568–76. 13. Yochum TR, Rowe LJ. Tumours and tumour like processes. In: Essentials of skeletal radiology, 2nd edn. Williams and Wilkins, 1996;975–1192. 14. Munk PL, Helms CA, Holt RG, Johnston J, Steinbach L, Neumann C. MR imaging of aneurysmal bone cysts. AJR Am J Roentgenol 1989;153:99–101. 15. Beltran J, Simon DC, Levy M, Herman L, Weis L, Mueller CF. Aneurysmal bone cysts: MR imaging at 1.5 T. Radiology 1986;158:689–90. 16. Willis RA. The spread of tumours in human body, London: Butterworths, 1973. 17. Galibert P, Deramond H, Rosat P, Le Gars D. Preliminary note on the treatment of vertebral angioma by percutaneous acrylic vertebroplasty. Neurochirurgie 1987;33:166–8.



CASE REPORT

Unusual CT/MR features of putative ligamentum flavum ossification in a North African woman 1

I JAFFAN, MD, 1B ABU-SERIEH, MD, 2T DUPREZ, MD, 2G COSNARD, MD and 1C RAFTOPOULOS, MD, PhD

Departments of 1Neurosurgery and 2Radiology, Saint-Luc Hospital, Universite´ Catholique de Louvain, Brussels, Belgium ABSTRACT. To display the unusual spinal CT and MR findings in a 48-year-old North African woman presenting with two adjacent intracanalar mineralized epidural outgrowths exhibiting mature bone organization with ‘‘cortical’’ and ‘‘trabecular’’ areas and pseudoarthritic changes at their interface. An unusual form of ligamentum flavum ossification (LFO) was speculated, of which features are discussed under the light of the available literature.


Ligamentum flavum ossification (LFO) is an acquired degenerative disease mainly affecting 20% of Asian individuals older than 65 years [1–4]. To our knowledge, there are around 31 reported cases in the non-Asian population [3, 4]. The pathogenesis of this form of ectopic ossification remains under debate. Some authors have correlated it to systemic diseases such as diabetes mellitus and hypoparathyroidism [5, 6]. Others have correlated LFO to the elevated plasma concentration of fibronectin (a proactive protein in endochondral ossification), independently of the patients’ endocrinological profiles [6]. On the other hand, some reports highlighted the importance of the mechanical factors in LFO development [7]. Classically, up to now, the radiological reports have shown LFO as intracanalar bone efflorescences made of cortical bone without spongy trabeculae [1–4]. No pseudoarthritic changes have been reported in this peculiar degenerative condition.

Case report A white North African 48-year-old woman with type II (non insulin-dependant) diabetes mellitus presented with a 2-year history of recurrent episodes of low back pain and sciatica resolving with oral analgesic treatment. She now complained of sensory disturbances and motor weakness in both lower limbs, predominantly on the right side. Neurological examination revealed spastic paraparesis with brisk but symmetric osteotendinous reflexes. Impairment of light touch and pain perception was observed at and below the T10 dermatome. Strength Address correspondence to: Christian Raftopoulos, Department of Neurosurgery, Saint Luc University Hospital, Universite´ Catholique de Louvain, Avenue Hippocrate 10, 1200 Brussels, Belgium.


testing was graded III/V in the right leg, and III+/V in the left one. Spiral CT and MR images disclosed the presence of two epidural abnormal beak-like excrescences at the dorsal aspect of T9–T10, which dramatically impinged on the spinal cord (Figures 1a and 2). Both abnormal outgrowths had remarkable mature bone organization with a peripheral margin of cortical bone, and a central area of trabecular bone, just as though they were supernumerary ossicles (Figure 1a,b). Moreover, pseudoarthritic changes were present at their interface with irregular thickening of the cortical margin (Figure 1a,d) and with the unexpected presence of a ‘‘degenerative’’ air-filled geode (Figure 1c). MR examination demonstrated similar textural organization with the additional relevant feature of the presence of fatty bone marrow within the trabeculae of the spongy component (Figure 2). The patient underwent selective posterior decompressive laminectomy at T9 and complete removal of the two bony outgrowths. The resected material was intimately adherent to the dura-mater. Histopathological examination of the specimens confirmed the presence of bone tissue, fibrocartilagineous tissue and fragments of ligamentum flavum without any tumorous component. The patient completely recovered within a few months after surgery.

Discussion LFO is a well-featured entity in the radiological literature, appearing as bilateral bulging-mass made of ectopic compact bone [1–4]. The most frequently reported location of LFO is the lower third of the thoracic spine [2, 8]. A few reports have described anatomical variants involving either unilateral ossification, or e67

I Jaffan, B Abu-Serieh, T Duprez et al

(a)

(b)

(c)

(d)

Figure 1. Spiral CT work-up. (a) Mid-sagittal reformatted image at the T9–T10 level showing intracanalar protrusion of two supernumerary bone outgrowths originating from the posterior elements. Observe cortical/spongy differentiation of the abnormal ‘‘bones’’, and pseudoarthritic changes at their interface, with focal cortical thickening and ‘‘degenerative’’ geode filled with air at the caudal aspect of the ‘‘pseudoarthrosis’’ (arrow). (b) Axial transverse reformatted image through the T9 level showing perfect cortical/spongy differentiation of abnormal ossicles together with calcifications of the ligamenta flava from which they are originating. (c) Axial transverse reformatted image through the T10 level showing pseudoarthritic thickening of the inferior border of the ‘‘pseudoarthosis’’, and the same air-filled geode as in Figure 1a (arrow). (d) Frontal reformatted image, well disclosing the pseudoarthritic interface between the two supernumerary ossicles. Their origin from posterior elements is obvious.

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Case report: Unusual CT/MR features of ligamentum flavum ossification

(a)

(b)

Figure 2. MR work-up. (a) Mid-sagittal T2 weighted image showing spinal cord compression at the T9–T10 level, with abnormal hyperintensity within cord parenchyma. (b) Mid-sagittal T1 weighted image showing hypointensity on the margins of the ‘‘pseudoarthrosis’’ due to reactive hyperostosis (arrowheads), and fatty bone marrow within central areas of the protruded ossicles (arrows). Normal epidural fat is present up and below the level of the intracanalar abnormalities (double ballarrowheads). Compare similar slice location on CT image (Figure 1a).


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unusual locations [9–12]. The observation reported herein highlighted two additional unreported features in the entity. The first was the perfect cortical/spongy differentiation, leading to a pattern of mature but supernumerary bone-like outgrowths showing thin cortical margin and central spongy bone trabeculae containing fatty marrow. This was well seen on both CT and MR images in a synergistic way. The former showed an intracanalar well-differentiated ‘‘bone’’ with cortical and trabecular areas (Figure 1a,b) and the latter showed the presence of hyperintense material corresponding to fatty marrow within trabeculae on the unenhanced T1 weighted images (Figure 2). The second peculiarity was the presence of pseudodegenerative changes at the interface between the two separated bony outgrowths, leading to a pseudoarthritic pattern of so-called neo-joint or neoarticulation. Changes included irregular thickening of the cortical margins at each side of contact areas and the presence of an air-filled pseudo-geode at the inferior border of the neo-joint (Figure 1a,d). The synergy between CT and MR information was also demonstrated for the degenerative changes since a thick hyperostotic/fibrotic border was seen as a hypointense marginal band-like area on T1 weighted MR images which extended far beyond the hyperdense thickening seen on CT reformats (Figure 2b). We consider that hyperdensity on CT images featured true reactive hyperostosis and that wider hypointensity on MR images represented the combination of hyperostosis together with fibrotic changes within the bone marrow adjacent to the pseudoarthritic interface. The described features strongly suggest a very unusual appearance of LFO rather than alternative hypotheses such as calcified meningioma, developmental abnormality or ossified haematoma. Calcified meningioma may exceptionally mimic the LFO [13]. However, it seemed highly unlikely in the presence of well-differentiated bone outgrowth. Besides, the histopathological examination did not reveal any meningeal tissue. A developmental abnormality such as this has never been described, and the absence of corresponding defect within adjacent vertebrae lowers the probability of such hypothesis. Ossified haematomas are poorly probable in the absence of acute episode of traumatic cord compression and in the presence of such homogeneous and symmetrical ossification. In conclusion, the herein reported case revealed two undescribed features in a putative case of LFO, i.e. the

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mature bone organization and the pseudoarthritic changes at the interface of two contiguous outgrowths.

References 1. Sugimura H, Kakitsubata Y, Suzuki Y, Kakitsubata S, Tamura S, Uwada O, et al. MRI of ossification of ligamentum flavum. J Comput Assist Tomogr 1992;16:73–6. 2. Miyasaka K, Kaneda K, Sato S, Iwasaki Y, Abe S, Takei H, et al. Myelopathy due to ossification or calcification of the ligamentum flavum: radiologic and histologic evaluations. AJNR Am J Neuroradiol 1983;4:629–32. 3. Ben Hamouda K, Jemel H, Haouet S, Khaldi M. Thoracic myelopathy caused by ossification of the ligamentum flavum: a report of 18 cases. J Neurosurg 2003;99:157–61. 4. Kruse JJ, Awasthi D, Harris M, Waguespack A. Ossification of the ligamentum flavum as a cause of myelopathy in North America: report of three cases. J Spinal Disord 2000;13:22–5. 5. Okazaki T, Takuwa Y, Yamamoto M, Matsumoto T, Igarashi T, Kurokawa T, et al. Ossification of the paravertebral ligaments: a frequent complication of hypoparathyroidism. Metabolism 1984;33:710–3. 6. Miyamoto S, Yonenobu K, Ono K. Elevated plasma fibronectin concentrations in patients with ossification of the posterior longitudinal ligament and ossification of the ligamentum flavum. Spine 1993;18:2267–70. 7. Maigne JY, Ayral X, Guerin-Surville H. Frequency and size of ossifications in the caudal attachments of the ligamentum flavum of the thoracic spine. Role of rotatory strains in their development. An anatomic study of 121 spines. Surg Radiol Anat 1992;14:119–24. 8. Okada K, Oka S, Tohge K, Ono K, Yonenobu K, Hosoya T. Thoracic myelopathy caused by ossification of the ligamentum flavum. Clinicopathologic study and surgical treatment. Spine 1991;16:280–7. 9. Mizuno J, Nakagawa H. Unilateral ossification of the ligamentum flavum in the cervical spine with atypical radiological appearance. J Clin Neurosci 2002;9:462–4. 10. Kubota M, Baba I, Sumida T. Myelopathy due to ossification of the ligamentum flavum of the cervical spine. A report of two cases. Spine 1981;6:553–9. 11. Kobayashi S, Okada K, Onoda K, Horikoshi S. Ossification of the cervical ligamentum flavum. Surg Neurol 1991;35:234–8. 12. van Oostenbrugge RJ, Herpers MJ, de Kruijk JR. Spinal cord compression caused by unusual location and extension of ossified ligamenta flava in a Caucasian male. A case report and literature review. Spine 1999;24:486–8. 13. Gamache FW Jr, Wang JC, Deck M, Heise C. Unusual appearance of an en plaque meningioma of the cervical spinal canal. A case report and literature review. Spine 2001;26:E87–9.



CASE REPORT

Double aortic arch with atresia, tapering and aneurysm of the left arch 1

H ITO, MD, PhD, 2A KONISHI, MD, PhD, 2T KON-NAI, MD, PhD, 3T ISHIBASHI, MD, PhD and 3S TAKAHASHI, MD, PhD Departments of 1Diagnostic Radiology and 2Cardiovascular Surgery, Miyagi Cardiovascular and Respiratory Center, Semine, Miyagi and 3Department of Diagnostic Radiology, Graduate School of Medicine, Tohoku University, Sendai, Japan ABSTRACT. An adult male underwent chest radiography for a health check-up. This disclosed both thoracic vascular anomalies and a small nodular shadow in the left side of the superior mediastinum. Axial MRI and three-dimensional volume-rendering MR angiography revealed both a double aortic arch with left atretic arch proximal to the left common carotid artery (subtype 4), and also tapering and aneurysm of the left arch distal to the left subclavian artery. To the authors’ knowledge, this report describes the first case of subtype 4 of atretic double arch with left arch atresia. Such thoracic vascular anomalies have been a theoretical possibility, but no cases have been reported to date.

Double aortic arch is formed from the splitting of the ascending aorta into two limbs that pass to either side of the trachea and oesophagus, and join as a single descending aorta. Anatomic variations of the double aortic arch relate to the relative size of the two arches, partial atresia of one arch and the side of the ductus and descending aorta. Partial obstruction of an arch in an anomalous double aortic arch usually occurs in the left arch; the atretic arch persists as a fibrotic cord [1, 2]. MRI can reveal the vascular anatomic details of congenital anomalies of the aortic arch including a double aortic arch [1, 3]. We report an adult male who has both a double aortic arch with atretic left arch (subtype 4), and also tapering and aneurysm of the left arch, and discuss the differential diagnoses of thoracic vascular anomalies of the present case.

Received 20 July 2004 Revised 17 March 2005 Accepted 19 September 2005 DOI: 10.1259/bjr/28729845 ’ 2006 The British Institute of Radiology

The Netherlands). T1 weighted spin-echo (SE) axial, coronal images were obtained (repetition time (TR)/echo time (TE) 1091/16, slice thickness 5 mm) (Figure 2). Gadolinium-enhanced MR angiography of the thoracic vessels was also performed with spoiled gradient echo

Case report A 39-year-old man underwent chest radiography during a regular health check-up. It revealed a thoracic vascular anomaly and nodular shadow in the left side of the superior mediastinum (Figure 1). He had been healthy since birth with no history of cardiac disease or hypertension. He had no significant history of shortness of breath or dysphagia. Echocardiography revealed no intracardiac defects. Laboratory data such as leukocyte count, serum creactive protein and immunoreactive protein, such as rheumatic factor, were all within normal limits. He underwent MRI examination on a 1.5-Tesla system (Gyroscan Intera 1.5 T Power; Philips Medical Systems, The British Journal of Radiology, August 2006

Figure 1. Chest radiograph indicates an abnormal thoracic aortic arch with right-sided tracheal impression (large arrow). In addition, it demonstrates a small nodular shadow (small arrow) the left of the superior mediastinum.

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(a)

(b)

Figure 2. (a) Axial T1 weighted spin-echo image shows a ‘‘four artery sign’’, which comprises the anterior common carotid (large black arrows) and posterior subclavian (small black arrows) arteries in the symmetrical position. Moreover, the small left arch (small white arrow) transforms into an aneurysm (large white arrow) via the narrow lumen. (T, trachea; E, oesophagus). (b) Axial T1 weighted spin-echo image 2.5 cm inferior to (a) demonstrates proximal right aortic arch (RA) with a retro-oesophageal segment, anterior common carotid arteries (large black arrows), posterior right superior vena cava (large white arrow) and posteroinferior distal aortic arch (small black arrow). The proximal and middle aortic arch compresses the right lateral and posterior wall of the trachea (T) with resultant narrowing of the lumen. Displaced azygos vein (large white arrow), which separates the azygos lobe from right upper lobe, drains into the superior vena cava. Double white arrows show the position of the oesophagus.

sequence. The following sequence parameters were used: TR/TE 4.8/1.5, flip angle 30 ˚, slice thickness 1.4 mm, pitch 0.5, matrix 205 6 256. Three-dimensional (3D) volume-rendering MR angiography was reconstructed using 3D software (Zio M900 Quadra; Amin, Tokyo, Japan) (Figure 3). The MR images demonstrated a double aortic arch. The first branch arising from the right aortic arch was the right common carotid artery, followed by the right subclavian artery and the left innominate artery. The left innominate artery derived from the distal portion of the aortic arch; it coursed upward and showed a gradual tapering of the lumen, then transforming into an aneurysm. Thereafter, the innominate artery distal to

(a)

the aneurysm descended anteriorly and inwardly, taking a course parallel to the right arch and dividing into an upward angled left subclavian artery. It then continued descending further in the same direction, before subsequently bending abruptly upward into the left common carotid artery.

Discussion Axial MRI and 3D volume-rendering MR angiography demonstrate that the right common carotid artery is the first branch and that the right subclavian artery is the second. They originate from the right-sided arch.

(b)

(c)

Figure 3. (a) Anteroposterior image of 3D volume-rendering MR angiography shows the right-sided aortic arch (RA), which divides the first two branches of the right common carotid (RC) and the right subclavian artery (RS). Then, the left subclavian (LS) and left common carotid arteries (LC) are derived from the third branch of the left innominate artery (arrow), which is eventually considered the small left arch (PA, pulmonary artery). (b) Deep left anterior oblique image of 3D volume-rendering MR angiography shows that the left subclavian (LS) and left common carotid artery (LC) are derived from the left arch (black arrow) locating proximal to the aneurysm (white arrow) (RC, right common carotid artery; RS, right subclavian artery; PA, pulmonary artery). (c) Deep right posterior oblique image of 3D volume-rendering MR angiography shows tapering and aneurysm (arrow) of the left arch (RA, right arch; PA, pulmonary artery; RC, right common carotid artery; RS, right subclavian artery; LC, left common carotid artery; LS, left subclavian artery).

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Case report: Double aortic arch with left atretic arch (subtype 4)

Because the right-sided arch is associated with a left-side descending thoracic aorta, the aortic arch passes behind the oesophagus. Therefore, the third branch of the left innominate artery originating from the distal part of the aortic arch would not be expected to further indent the oesophagus. Indeed, it courses upward without crossing the midline. Moreover, the third branch distal to the aneurysm is parallel to the right-sided arch: it divides into an upward left subclavian artery, and then into an upward left common carotid artery. Although the segment distal to the left common carotid artery is not visible, the left common carotid artery and the right-sided aortic arch are in close approximation. We infer that this results from an atretic segment of the left arch. From our patient, we have concluded that the thoracic vascular anomalies of the present case are manifested by atresia of the left arch during embryonic development. Therefore, the third branch, which is described as the left innominate artery, is identical to the left arch. The most likely diagnosis of the present case is a double aortic arch with atretic left arch proximal to the left common carotid artery [1, 4]. A double aortic arch is the most common cause of a symptomatic vascular ring in infants and young children [1]. However, patients with a vascular ring in association with a double aortic arch rarely have accompanying symptoms even in adulthood, as in our patient [5, 6]. Barium oesophagography may suggest a vascular ring, but does not allow definition of the exact configuration of the anomaly. Four subtypes of double aortic arch with left arch atresia exist, according to the site of atresia of the left arch [1]. The hypothetical double aortic arch (Figure 4a) was proposed by Edwards [7]. The zones of the left arch are lettered from back to front as A–D, as first described by Garti et al [9]. A double aortic arch with left arch atresia is classified into subtype 1 if atresia occurs between the ligamentum arteriosum or left ductus and the descending aorta (Zone A). Atresia between the left subclavian artery and the ligamentum arteriosum or left ductus (Zone B), or between the left common carotid artery and the left subclavian artery (Zone C), respectively, forms subtype 2 or 3. Subtype 4 is designated by atresia proximal to the left common carotid artery (Zone D). In the present case, subtype 3 is first excluded from the differential diagnosis. In subtypes 1 and 2, the left innominate artery arises as the first branch from the ascending aorta. For that reason, these two subtypes are excluded. In subtype 4, the left innominate artery would arise as the third branch from descending aorta: this latter subtype is the most likely diagnosis in our case. No such cases have been documented to date [1]. Atresia of the left arch (Zone D) with tapering and aneurysm (Zones A and B) is illustrated in Figure 4b. Most aortic arch anomalies with an atretic left arch (Zone D) tend not to be associated with complex congenital heart disease [1, 2, 5]. This rule applies to the present case. This observation is also in agreement with descriptions in other cases of isolated left subclavian and left carotid arteries where there is atresia of zone D in association also with zones A, B and C [10–12]. The British Journal of Radiology, August 2006

(a)

(b) Figure 4. (a) The hypothetical double aortic arch system proposed by Edwards, showing zones of the left arch labelled A–D (RC, right common carotid artery; RS, right subclavian artery; LC, left common carotid artery; LS, left subclavian artery; AA, ascending aorta; PA, pulmonary artery; rda, right ductus arteriosus; lda, left ductus arteriosus). (b) Atresia of left arch (LAA) along the zone D in hypothetical double aortic arch demonstrates subtype 4 of the double aortic arch with left atretic arch. Moreover, tapering (Tap) and aneurysm (An) are probably located, respectively, in the A and B zones, but the correct origin of the left ductus arteriosus (lda) from the left arch is uncertain (RC, right common carotid artery; RS, right subclavian artery; LC, left common carotid artery; LS, left subclavian artery; AA, ascending aorta; PA, pulmonary artery; rda, right ductus arteriosus; lda, left ductus arteriosus). Illustrations are modified from those by B Singh et al [8].

However, the pattern of branching in our patient is identical to that seen in right aortic arch with an aberrant left innominate artery and ligamentum arteriosum. Therefore, a right aortic arch with a left aberrant innominate artery can not be excluded from the differential diagnosis [1, 4], whereas a left aberrant innominate artery in association with a right aortic arch usually derives from a right-sided aortic arch and courses across the midline behind the oesophagus [4]. e73

H Ito, A Konisho, T Kon-nai et al

An additional nodular radio-opacity in the superior mediastinum is commonly seen in patients with thoracic vascular anomalies [13, 14]. In the present case, the nodular shadow adjacent to the left side of the superior mediatinum is caused by the aneurysm of the left aortic arch, distal to the origin of the left subclavian artery. Arterial tapering and aneurysm are irreversible changes. They occur in association with various aetiologies such as degenerative disease, trauma, infective or inflammatory aortitis, connective tissue disorders and congenital abnormalities [15]. Laboratory data and the clinical history in our patient were normal. We suggest that the arterial changes we have reported are congenital. In summary, we report an extremely rare case of not only a double aortic arch with an atretic left arch (subtype 4), but also anomalous tapering and aneurysm of the left arch distal to the origin of the left subclavian artery. To our knowledge, the present report describes the first case with subtype 4 of a double aortic arch with a left atretic arch, although such a thoracic vascular anomaly has previously been suggested as theoretically possible.

Acknowledgments We thank Mr Takahiro Chiba (Radiological Technologist of Kurihara Chuo Hospital, Tsukidate, Miyagi, Japan) for his skilful assistance with the MRI examination.

References 1. Jaffe RB. Radiographic manifestations of congenital anomalies of the aortic arch. Radiol Clin North Am 1991;29:319–34. 2. Frederic Moes CA, Freedom RM. Rare types of aortic arch anomalies. Pediatr Cardiol 1993;14:93–101.

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3. Kersting-Sommerhoff BA, Sechtem UP, Fisher MR, Higgins CB. MR imaging of congenital anomalies of the aortic arch. AJR Am J Roentgenol 1987;149:9–13. 4. Midiri M, Finazzo M, Pilato M, Lagalla R, De Maria M. Right aortic arch with aberrant left innominate artery: MR imaging findings. Eur Radiol 1999;9:311–5. 5. Ruckman RN. Anomalies of the aortic arch complex. In: Adams FH, Emmanouilides GC, Riemenschneider TA, editors. Moss’ heart disease in infants, children and adolescents. Philadelphia, PA: Williams & Wilkins, 1989:255–72. 6. Simoneaux SF, Bank ER, Webber JB, Parks WJ. MR imaging of the pediatric airway. Radiographics 1995;15:287–98. 7. Edwards JE. Anomalies of the derivatives of the aortic arch system. Med Clin N Am 1948;32:925–45. 8. Singh B, Satyapal KS, Moodley J, Rajaruthnam P. Right aortic arch with isolated left brachiocephalic artery. Clin Anat 2001;14:47–51. 9. Garti IJ, Aygen MM, Vidne B, Levy MJ. Right aortic arch with mirror-image branching causing vascular ring: a new classification of the right aortic arch patterns. Br J Radiol 1973;46:115–9. 10. Boren EL Jr, Matchett WJ, Gagne PJ, McFarland DR. Isolation of the left innominate artery in an elderly patient without congenital heart disease. Cardiovasc Interv Radiol 2000;23:63–5. 11. Crump WD, Dische MR, Anthony CL. Right aortic arch, isolated left common carotid and left subclavian arteries, and subclavian steal syndrome: a variant to polysplenia syndrome. Hum Pathol 1981;12:936–8. 12. Bialowas J, Hreczecha J, Grzybiak M. Right-sided aortic arch. Folia Morphol (Warsz) 2000;59:211–6. 13. Shahian DM, Javid H, Faber LP, Kittle CF, Matthew GR. Lesions of the thoracic aorta and its arch branches simulating neoplasm. J Thorac Cardiovasc Surg 1981;81:251–63. 14. Ohgi S, Ito H, Isogami K, Kanno T. Cardiovascular imaging in-a-month. Chest roentgenogram mimicking double aortic arch in a 30-year-old female with effort dyspnea and dysphagia. J Cardiol 2004;43:147–50. 15. Ko S-F, Wan Y-L, Ng S-H, et al. MRI of thoracic vascular lesions with emphasis on two-dimensional time-of-flight MR angiography. Br J Radiol 1999;72:613–20.



CASE REPORT

Central line pump infusion and large volume mediastinal contrast extravasation in CT P O’SULLIVAN,

MRCPI, FFR, RCSI,

M BROWN,

MD,

B HARTNETT,

BSc, RT(R)

and J R MAYO,

MD

Department of Radiology, Vancouver General Hospital, University of British Columbia, Vancouver, BC, Canada ABSTRACT. The use of multidetector CT scanners for CT angiography requires rapid injection of radiographic contrast media. Central venous catheters are now widely used for this purpose. Several complications may occur while using central venous access for rapid, large volume contrast injection such as catheter rupture and contrast extravasation. We describe a case in which inadvertent malposition of a central venous catheter led to a high volume extravasation of contrast in the mediastinum in a trauma patient.

Received 14 July 2005 Revised 8 September 2005 Accepted 19 September 2005 DOI: 10.1259/bjr/30451779 ’ 2006 The British Institute of Radiology

The use of multidetector CT scanners for CT angiography and CT brain perfusion studies requires rapid injection (2–5 ml s21) of radiographic contrast media. Peripheral intravenous access is preferred. However, in patients in whom this is unavailable, the use of central venous catheters for rapid power injection of contrast media may be considered [1, 2]. Possible complications that may occur while using central catheters for high flow rate injection include catheter rupture, contrast extravasation and delayed catheter malfunction [3]. We describe a case in which inadvertent malposition of the central venous catheter led to a high volume (100 ml) extravasation of contrast within the mediastinum in a young male trauma patient.

against the lateral wall of the superior vena cava (SVC). This catheter was pulled back 2 cm to avoid possible SVC trauma [4–6] (Figure 1). The patient deteriorated over the next 24 h, with a drop in haemoglobin from 11.2 g dl21 to 8.0 g dl21. A contrast enhanced CT scan of the thorax was ordered to evaluate for undiagnosed chest traumatic injury. Upon arrival in the CT suite, the attending ICU nurse flushed, without difficulty, the proximal port of the central catheter with normal saline. She then connected the contrast power injection pump to this port of the multilumen central catheter. A trauma protocol chest CT scan was performed using an 8 channel multidetector row CT scanner (GE Light Speed QX1 Ultra; GE Medical Systems, Milwaukee, WI) with an injection of 100 cm3 of

Case report A 24-year-old male patient was admitted to the emergency department after a high-speed motor vehicle accident. His vital signs were stable, and clinically lower and upper limb fractures were suspected. Initial chest, spine, pelvis and lower limb radiographs showed bilateral lower limb fractures, knee dislocations and hand fractures. The chest radiograph was interpreted as normal. A CT scan of the head, abdomen and pelvis did not reveal any significant abnormality. The patient was admitted to the intensive care unit for ventilatory support. Upon admission to ICU, a left sided, triple lumen, 16 cm, 7-French Arrow-Howes (Arrow International, Reading, PA) [3] central venous catheter was placed in the left subclavian vein by a staff anaesthetist. Peripheral access was avoided due to the multiple upper limb injuries. The postplacement chest radiograph demonstrated the line tip Address correspondence to: Dr Paul O’Sullivan, 899 West 12 Avenue, Vancouver, BC, Canada, V6J 1V9. Email: paul.osullivan @vch.ca.


Figure 1. Following withdrawal of the subclavian line by 2 cm, the tip (arrow) of the multilumen catheter appears to be in a satisfactory position within the left brachiocephalic vein. e75

P O’Sullivan, M Brown, B Hartnett and J R Mayo

(a)

(b)

(c)

Figure 2. Transverse CT section at the level of (a) the manubrium and (b) the carina, demonstrating a large volume of extravasated contrast within soft tissues of the anterior mediastinum. (c) Sagittal image shows extensive anterior mediastinal extravasation.

intravenous contrast media (Optiray/Ioversol 320) at 2.5 ml s21. The flow rate was reduced from 3 ml s21 (our usual for trauma protocol) to 2.5 ml s21 (the maximum flow rate we allow via central access). Contiguous 1.25 mm sections were reconstructed using the standard algorithm and showed a 4 cm oval shaped area of high attenuation within the anterior mediastinum (Figure 2), in keeping with a large volume of intravenous contrast media extravasation. As the patient was intubated and under sedation, there were no abnormal symptoms reported. The patient’s vital signs remained stable during this time. A repeat non-contrast chest CT performed 5 h later showed diffuse spread of the extravasated contrast media throughout the mediastinum, pleural spaces (Figure 3) and into the soft tissues of the neck. The central line was left in place for the next 24 h, although the ICU staff did not use it. The following day, a hand-injected venogram was performed to assess line position, showing the proximal port to be partially outside the vein (Figure 4). The line was subsequently removed without further complication. The line was noted to be fully intact, with no evidence of rupture after removal.

Figure 3. Transverse CT section at the level of the right pulmonary artery shows contrast media outlining the pericardium (arrows) and diffusing into the left and right pleural spaces. e76

The patient had an uneventful recovery from the mediastinal extravasation. He received another chest CT scan 1 week later to exclude pulmonary embolism. No embolus was identified, and the previously identified extravasated contrast media in the mediastinum was seen to have completely resorbed.

Discussion The use of central venous catheters for rapid power injection of contrast is controversial, with manufacturers cautioning their use in this role and with clinicians reporting rare complications including catheter rupture,

Figure 4. Hand injection of the proximal port of the multilumen catheter shows extravasation from the proximal side port (arrow) confirming the malposition of the line. Superior vena cava (SVC) is also shown. The British Journal of Radiology, August 2006

Case report: CT pump infusion and mediastinal contrast extravasation

malfunction, cardiac arrhythmias, contrast extravasation and infection [3]. We report an unusual cause of contrast extravasation associated with injection of the proximal port of an otherwise correctly functioning multilumen central catheter. The large volume contrast extravasation that occurred in our case was due to the combination of factors including the large patient size, repositioning of the catheter to avoid perforation of the vein wall and injection of the lumen exiting 5 cm proximal to the end of the catheter. To avoid this complication in multilumen catheters, staff should be aware of the position of ports relative to the catheter tip and recognize that ports may appear to function normally using hand injections of saline but prove to be malpositioned when a power injector is used. Ideally, contrast media should only be administered through the distal port. If there is doubt as to the adequacy of line placement, a limited volume noncontrast CT scan of the line can be performed to confirm line position prior to the large volume power contrast media injection.


References 1. Herts BR, O’Malley CM, Wirth SL, Lieber ML, Pohlman B. Power injection of contrast media using central venous catheters: feasibility, safety and efficacy. AJR Am J Roentgenol 2001;176:447–53. 2. Herts BR, Cohen MA, McInroy B, Davros WJ, Zepp RC, Einstein D. Power injection of intravenous contrast through central venous catheters for CT: in vitro evaluation. Radiology 1996;200:731–5. 3. Sanelli PC, Deshmukh M, Ougorets I, Caiati R, Heier LA. Safety and feasibility of using a central venous catheter for rapid contrast injection rates. AJR Am J Roentgenol 2004;183:1829–34. 4. Tocino IM, Watanabe A. Impending catheter perforation of superior vena cava: radiographic recognition. AJR Am J Roentgenol 1986;146:487–90. 5. Dailey RH. Late vascular perforations by CVP catheter tips. J Emerg Med 1988;6:137–40. 6. Trigaux JP, Goncette L, Van Beers B, de Wispelaere JF, Pringot J. Radiologic findings of normal and compromised thoracic venous catheters. J Thorac Imaging 1994;9: 246–54.

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CASE REPORT

High resolution MRI in the detection of an intrathymic parathyroid adenoma 1

G ABIKHZER,

MDCM,

2

M LEVENTAL,

MDCM, FRCPC

and 3C RUSH,

MDCM, FRCPC

1

Department of Nuclear Medicine, McGill University and Departments of 2Radiology and 3Nuclear Medicine, Sir Mortimer B. Davis - Jewish General Hospital, 3755 Coˆte Ste. Catherine Road, Montreal, Quebec, H3T 1E2, Canada

ABSTRACT. The most common cause of primary hyperparathyroidism is parathyroid adenoma. Of these, up to 20% have been reported to be in ectopic locations. Multiple imaging techniques have been utilized in the detection of ectopic adenoma and are discussed. We report on a case of an intrathymic parathyroid adenoma that was detected through the novel use of a breast coil in performing a high resolution MRI. The accurate localization permitted minimally invasive surgery, obviating morbidity associated with a sternotomy.

Received 11 February 2005 Revised 31 August 2005 Accepted 6 September 2005 DOI: 10.1259/bjr/26663397 ’ 2006 The British Institute of Radiology

Primary hyperparathyroidism is an important cause of morbidity and has been shown to increase the incidence of cardiovascular abnormalities, diabetes mellitus and urogenital diseases, as well as to increase the risk of mortality [1]. Although 54% of patients are asymptomatic at diagnosis, 58% of all patients exhibit some signs of bone disease [2]. Of those who are symptomatic, the most common symptoms include fatigue and mental status changes, renal calculi and gastric ulcers. We recently encountered a patient with primary hyperparathyroidism in whom a Technetium (Tc) 99m Sestamibi scan (MIBI) revealed increased uptake in the superior mediastinum. CT imaging failed to reveal the exact location of the lesion while standard MRI suggested an anterior mediastinal location. We report on the novel use of a breast coil in performing a high resolution MRI which accurately localized the lesion, allowing minimally invasive surgery.

Case report A 38-year-old woman presented with a palpable thyroid nodule. Investigation revealed a papillary carcinoma of the thyroid. Patient underwent total thyroidectomy with preservation of the parathyroid (PTH) glands. Routine studies revealed hypercalcaemia of 2.87 (2.12–2.62) mmol l21 and PTH level of 86 (10–65) pg ml21. MIBI scan (Millenium VG; GE Medical Systems, Milwaukee, WI; Cardiolite; Bristol-MyersSquibb, New York, NY) revealed no abnormal activity in the neck, but a 2 cm focus of intense uptake in the superior mediastinum (Figure 1). The increased uptake

Address correspondence to: Dr Mark Levental.

e78

was felt to represent either a parathyroid adenoma or a pathologic lymph node. In order to further localize the lesion pre-operatively, the patient underwent uninfused CT of the thorax, which failed to reveal any abnormality. A standard thoracic MRI (Intera; Philips Medical Systems, Eindhoven, The Netherlands) suggested a possible lesion in the anterior mediastinum (Figure 2). A final attempt was made to accurately localize the lesion pre-operatively by performing a high resolution MRI of the anterior mediastinum using a breast coil (Bilateral breast coil; Phillips Medical Systems). T1 weighted, T2 weighted, fat saturated T2 weighted and post-gadolinium infused T1 weighted fat saturated images were obtained. These revealed a sharply defined 9 mm65 mm focus, isointense on T1 weighting, hyperintense on T2 weighing fat saturated images (Figure 3) and which showed enhancement with contrast administration. The lesion was in keeping with a parathyroid adenoma and was accurately localized in the anterior mediastinum fat in the area corresponding to residual thymic tissue. The lesion appeared to be 4.5 cm inferior to the sternal notch, abutting the anterior wall of the ascending aorta. This allowed the Cardiothoracic Surgeon to perform a small window overlying the third coastal cartilage, which was removed together with a small segment of the adjacent rib, permitting the resection of the lesion in toto with immediate intraoperative decline of PTH levels. Pathological diagnosis of fat depleted parathyroid admixed with fragments of thymus was made intraoperatively. The surgery lasted 85 min. Pre-operative PTH levels were 174 pg ml21 and Ca levels were up to 3.05 mmol l21. The levels subsequently declined to 17 pg ml21 and 2.05 mmol l21, respectively, within 12 h post-surgery. The patient made an uneventful The British Journal of Radiology, September 2006

Case report: MRI detection of intrathymic parathyroid adenoma

Figure 1. Anterior planar 99Tcm Sestamibi image of the neck and thorax taken 10 min post-tracer injection. There is no focal tracer uptake in the neck region. There is an intense focus of abnormal tracer uptake projecting in the plane of the superior mediastinum consistent with an ectopic parathyroid adenoma.

recovery and was discharged on post-operative day 1. She remains clinically and biochemically normal.

Discussion Primary hyperparathyroidism is the most common cause of hypercalcaemia with an annual incidence of 27.7

Figure 2. Axial fat saturated T2 weighted (repetition time 1714, echo time 80, slice thickness 4 mm, skip 1 mm, matrix 1596256) image through mid thorax with respiratory and cardiac gating using the phased array coil. There is a subcentimetric focus of hyperintensity within the residual atrophic thymic tissue, lying just anterior to the ascending aorta. There is some failure of fat suppression and/or slow flow in veins in the subcarinal mediastinum. The British Journal of Radiology, September 2006

Figure 3. Axial fat saturated T2 weighted (repetition time 5043, echo time 90, slice thickness 3 mm, skip 0 mm, matrix 2566512) image through mid thorax using the breast coil. There is a 9 mm very clearly defined focus of hyperintensity in the residual thymus tissue of the anterior mediastinum, representing ectopic parathyroid adenoma.

per 100 000 [3]. Of these cases, 87% are attributable to a solitary adenoma, 3% to multiple adenomata, 9% to hyperplasia, 1% to parathyroid carcinoma [4]. 5% of adenomata have an ectopic location. Sites of ectopic localization include thymus, tracheo-oesophageal groove, carotid sheath, intrathyroidal and paraoesophageal locations [2]. This is explained by the common embryological origin of both the inferior parathyroid gland and thymus from the third pharyngeal pouch, both migrating caudally to their respective locations. [5]. Historically, 95% of cases primary hyperparathyroidism were diagnosed and cured by a bilateral cervical exploration. The remaining 5% of cases required additional surgery with associated cost and morbidity. Subsequent surgeries without pre-operative localization studies had a 38% failure rate [6]. Pre-operative localization has allowed for minimally invasive surgery to be performed with decreased mean operative time [4] and a 1.2% complication rate, compared with 3% when no imaging is used [7]. Current imaging strategies for detection of parathyroid adenomata include ultrasound, 99Tcm-MIBI, CT and MRI. Ultrasound has a 56% [8] to 74% [9] success rate for localization of orthotopic adenomata, but has not been shown to be useful in locating ectopic glands in the mediastinum [10]. Many studies have shown the cost effectiveness of preoperative localization with MIBI in all hyperparathyroid patients [2, 4]. It has a sensitivity of 75% [8] to 92% [2] with a positive predictive value of 91% [2], permits a total body scan, but suffers from poor spatial resolution and 9% false positive rate mostly due to concomitant thyroid abnormalities [2]. The poor anatomic information regarding the relationship of the adenoma with surrounding structures renders minimally invasive surgery difficult [9], a procedure used by 59% of endocrine surgeons according to a recent survey [11]. CT has 40% [12] to 86% sensitivity, depending on the size of the lesion, and is capable of detecting mediastinal e79

G Abikhzer, M Levental and C Rush

adenomata [13]. MRI is preferable to CT scanning due to the greater sensitivity, lack of ionizing radiation, multiplanar capability and superior tissue contrast. On T1 weighted images an adenoma appears as low signal intensity, isointense relative to muscle. High signal intensity is most commonly observed on T2 weighted images and lesions are particularly visible when fat saturation is applied. Hyperfunctioning parathyroid glands characteristically show intense contrast enhancement on T1 weighted images. MRI has shown promise in patients with persistent or recurrent hyperparathyroidism after previous surgery or in patients where an initial MIBI scan revealed an ectopic gland. MRI has an overall sensitivity of 77% [14] to 82% and positive predictive value of 89% [15]. MRI sensitivity in correctly locating mediastinal ectopic adenomata is 88% and may reach 100% sensitivity in cases of intrathymic ectopic adenomata [15]. MRI has a 1.6% rate of false positive results, typically due to concomitant thyroid disease or enlarged lymph nodes [14], the latter mimicking the appearance of an adenoma. The combination of a MIBI scan and MRI can combine a functional imaging technique with a precise anatomical scan and has been shown to increase sensitivity to 94% with a positive predictive value of 98% in patients with persistent hyperparathyroidism [15]. In our patient, standard thoracic MRI performed with fat saturated T2 weighted imaging suggested a possible lesion in the anterior mediastinum. We attempted, however, to obtain a higher resolution assessment of the anterior mediastinum in order to provide a detailed roadmap for minimally invasive surgery. Utilizing a breast coil, 2566512 matrix and volume shimming yielded homogeneous, high resolution fat saturated imaging of the anterior mediastinum and revealed a ,1 cm lesion very accurately whereas traditionally, MRI has had difficulty in locating lesions less than 1.5 cm in diameter [10]. The advantage of using a breast coil in a woman is that it allows for close contact of the receiving antenna with the chest wall, whereas in a typical thoracic MRI using a phased array body coil breast tissue displaces the coil anteriorly and may interfere with homogeneous fat saturated image formation. Given the morbidity of open sternotomy and surgical exploration, the ability to precisely localize this small lesion allowed a very limited, minimally invasive surgical intervention with a short operating time and improved cosmetic result. The patient was discharged 1 day post-surgery in contradistinction to the average hospital stay of 14 days when sternotomy is used for excision of ectopic mediastinal adenomata [16]. Our case report supports the conclusion of previous studies in the literature, that pre-operative imaging of hyperparathyroidism allows minimally invasive surgical intervention [2, 4]. We propose that in cases of anterior mediastinal lesions either not detected or not clearly delineated by routine thoracic MRI, that the study be performed with high resolution surface coil to increase chances of detecting a small lesion.

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References 1. Nilsson IL, Yin L, Lundgren E, Rastad J, Ekbom A. Clinical presentation of primary hyperparathyroidism in Europe-nationwide cohort analysis on mortality from nonmalignant causes. J Bone Miner Res 2002;17 Suppl. 2:N68–74. 2. Pellitteri PK. Directed parathyroid exploration: evolution and evaluation of this approach in a single-institution review of 346 patients. Laryngoscope 2003;113:1857–69. 3. Heath H 3rd, Hodgson SF, Kennedy MA. Primary hyperparathyroidism. Incidence, morbidity, and potential economic impact in a community. N Engl J Med 1980;302:189–93. 4. Denham DW, Norman J. Cost-effectiveness of preoperative sestamibi scan for primary hyperparathyroidism is dependent solely upon the surgeon’s choice of operative procedure. J Am Coll Surg 1998;186:293–305. 5. Weller GL. Development of the thyroid, parathyroid and thymus glands in man. Contrib Embryol 1933;24:93–143. 6. Satava RM Jr, Beahrs OH, Scholz DA. Success rate of cervical exploration for hyperparathyroidism. Arch Surg 1975;110:625–8. 7. Udelsman R. Six hundred fifty-six consecutive explorations for primary hyperparathyroidism. Ann Surg 2002;235:665–70. 8. Quiros RM, Alioto J, Wilhelm SM, Ali A, Prinz RA. An algorithm to maximize use of minimally invasive parathyroidectomy. Arch Surg 2004;139:501–6. 9. van Dalen A, Smit CP, van Vroonhoven TJ, Burger H, de Lange EE. Minimally invasive surgery for solitary parathyroid adenomata in patients with primary hyperparathyroidism: role of US with supplemental CT. Radiology 2001;220:631–9. 10. Kang YS, Rosen K, Clark OH, Higgins CB. Localization of abnormal parathyroid glands of the mediastinum with MR imaging. Radiology 1993;189:137–41. 11. Sackett WR, Barraclough B, Reeve TS, Delbridge LW. Worldwide trends in the surgical treatment of primary hyperparathyroidism in the era of minimally invasive parathyroidectomy. Arch Surg 2002;137:1055–9. 12. Ishibashi M, Nishida H, Hiromatsu Y, Kojima K, Uchida M, Hayabuchi N. Localization of ectopic parathyroid glands using technetium-99m sestamibi imaging: comparison with magnetic resonance and computed tomographic imaging. Eur J Nucl Med 1997;24:197–201. 13. Gross ND, Weissman JL, Veenker E, Cohen JI. The diagnostic utility of computed tomography for preoperative localization in surgery for hyperparathyroidism. Laryngoscope 2004;114:227–31. 14. McDermott VG, Fernandez RJ, Meakem TJ 3rd, Stolpen AH, Spritzer CE, Gefter WB. Preoperative MR imaging in hyperparathyroidism: results and factors affecting parathyroid detection. AJR Am J Roentgenol 1996;166: 705–10. 15. Gotway MB, Reddy GP, Webb WR, Morita ET, Clark OH, Higgins CB. Comparison between MR imaging and 99mTc MIBI scintigraphy in the evaluation of recurrent of persistent hyperparathyroidism. Radiology 2001;218:783–90. 16. Cupisti K, Dotzenrath C, Simon D, Roher HD, Goretzki PE. Therapy of suspected intrathoracic parathyroid adenomata. Experiences using open transthoracic approach and videoassisted thoracoscopic surgery. Langenbecks Arch Surg 2002;386:488–93.



CASE REPORT

A vascular ring variant: an unusual case of vocal cord palsy due to an anomalous left carotid artery arising from a retrotracheal arch of the aorta 1

A R JOSHI,

MD,

1

A GARG,

DNB,

1

B VHANMANE,

MB, BS,

1

S MERCHANT,

MD

and 2N NERURKAR,

MS

Departments of 1Radiology and 2Otolaryngology, Lokmanya Tilak Medical College and Municipal General Hospital, Mumbai, India ABSTRACT. We present the case of an unusual vascular ring in a 23-year-old male patient who presented for CT examination of the chest for diagnostic work up for unilateral vocal cord palsy. Contrast-enhanced CT revealed the left common carotid artery to be arising from the ascending part of a retrotracheal arch of aorta and traversing anterior to the trachea in the retromanubrial space.

Received 23 May 2005 Revised 12 August 2005 Accepted 30 September 2005 DOI: 10.1259/bjr/25504514 ’ 2006 The British Institute of Radiology

Vascular rings are uncommon congenital anomalies that occur early in the development of the aortic arch and great vessels. The primary symptomatology associated with these anomalies relates to the structures that are encircled by the ring: the trachea and oesophagus. Rarely may it present with the more common problem of vocal cord paralysis, which is itself a sign of disease and not a diagnosis.

Case report A 23-year-old male patient was referred to our department for contrast-enhanced CT of chest as part of the diagnostic assessment of unilateral vocal cord palsy. He had presented with hoarseness of voice and inability to speak coherently since childhood. Flexible fibre optic laryngoscopy under local anaesthesia revealed reduced left vocal fold closure and decreased sensation in the pharynx. Plain and contrast enhanced multislice CT (Volume Zoom, Siemens, Germany) was performed from the base of skull to mid chest (arch of the aorta). The chest CT (lung windows) was unremarkable. On mediastinal window settings, the arch of the aorta was in an anomalous location, being

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