Chapter 1
Overview of Brachytherapy Physics Ravinder Nath, Ph.D. Department of Therapeutic Radiology Yale University School of Medicine New Haven, Connecticut Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1 Different Types of Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 Intracavitary Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 Interstitial Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 Intraluminal Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Surface Molds and Plaques for Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Different Types of Brachytherapy Loading Systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Manual “Hot” Loading Techniques . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Manual Afterloading Techniques . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Remote Afterloading Techniques . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Image-Guided Techniques . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4 Different Types of Brachytherapy Based on Dose Rate and Duration of Brachytherapy . . . . . . 4 Different Durations of Brachytherapy Typically Used . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4 Permanent Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4 Temporary Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4 Dose Rates Used in Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5
Introduction Brachytherapy is a special procedure in therapeutic radiology that utilizes the irradiation of a target with radioactive sources placed at short distances from the target. Typically, the sources are implanted in the target tissue directly (interstitial brachytherapy) or are placed at distances of the order of a few millimeters from the target tissue, in body cavities such as the uterus, mouth, etc. (intracavitary brachytherapy), or externally on structures such as the eye, skin, etc., (surface plaques or molds). Brachytherapy generates highly conformal dose distributions in a target volume because radioactive seeds are implanted directly within or in the vicinity of the target tissue. For example, in a typical interstitial brachytherapy implant, 50 to 100 radioactive seeds, each about the size of a rice grain, are implanted in the tumor using image-guided implantation techniques such as ultrasound, computed tomography (CT), or fluoroscopy, which allow the physician to place radioactive seeds precisely at desired locations with minimal invasiveness. For these applications, low-energy, photon-emitting radionuclides such as 125I (27 keV) and 103Pd (21 keV) are preferred because these sources provide adequate coverage of tumor when used in a grid of about 1 cm spacing and produce minimal exposures to distant organs in the patient, to the hospital personnel performing the procedure, and to the family members and friends of the patient after he/she is released from the hospital with the radioactive seeds in place. Most brachytherapy procedures today are performed in one-day surgery suites without the need for hospitalization. These factors and the depth dose characteristics make brachytherapy a very cost effective and patient friendly procedure compared to 3-D conformal radiotherapy (3DCRT) or intensity-modulated radiation therapy (IMRT), which also produce highly conformal dose distributions. A key advantage of 3DCRT or IMRT over brachytherapy is that it is noninvasive. However, both 3DCRT and IMRT are very sensitive to patient localization and setup errors because of high dose gradients at the periphery of the target volume. Therefore, the target must be placed at the right position with a precision of about a millimeter relative to the linear accelerator (linac) daily over a course of 5 to 6 weeks of 3DCRT and IMRT, and the clinical target volume
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expanded by 5 to 10 mm for intrafraction movement. In contrast, brachytherapy requires a single visit to a one-day surgery clinic or 3 to 5 visits in an outpatient clinic for high dose-rate (HDR) intracavitary brachytherapy. Unlike 3DCRT and IMRT, brachytherapy is far more forgiving of localization and target motion errors because the implanted sources of radiation in an interstitial implant move with the target. Thus, brachytherapy solves a critical problem of 3DCRT and IMRT, which is that they have a potential to “miss the target very precisely” unless they are implemented with extreme precision. This high-precision requirement makes 3DCRT and IMRT very expensive, labor intensive, and technically complex. For these practical reasons and other important radiobiological reasons related to continuous low doserate (LDR) irradiation, brachytherapy remains a valuable treatment modality for selected cancers despite the current trend of widespread adoption of 3DCRT and IMRT in radiotherapy. In this chapter, an overview of the rich spectrum of various brachytherapy procedures is presented so that the students have an appreciation of the overall field. Most brachytherapy uses gamma-emitting radionuclides. Many use radionuclides such as 137Cs and 192 Ir, which emit high-energy gammas, that penetrate deeply and also require heavy shielding for radiation protection of the personnel and the patient’s family. Whenever possible, brachytherapy with low-energy gamma emitters, like 125I and 103Pd is preferred because it requires minimal shielding for radiation protection. On the other hand, beta-emitting sources are commonly used as unsealed sources for systemic brachytherapy similar to nuclear medicine diagnostic procedures. Beta particles are absorbed within a few millimeters in tissues. Hence, beta emitters require minimal radiation shielding unless there is spillage of radioactive material in a liquid form and subsequent contamination of work surfaces. Beta emitters commonly used for brachytherapy include 32P, 106Ru, 90Sr, and 90Y. In the last decade, the betaemitting sources also generated a great deal of interest in intravascular brachytherapy for prevention of restenosis after angioplasty because the distances of interest for treatment are about 2 mm, typical wall thickness of a coronary blood vessel. In this application, the sealed sources of 32P and 90Sr were widely used and 90Sr is still being used today. The only neutron emitter, which has been used clinically, is 252Cf. The advantage of neutron emitters is that their interactions with tissues produce heavy charged particles that have much higher linear energy transfer (LET) than the secondary electrons produced by gammas or betas and because of the higher LET, neutrons are theoretically more effective against hypoxic tumors. Neutrons are, however, notoriously difficult to shield and present a far greater challenge in reducing the radiation hazards to personnel. For these reasons, neutrons are rarely used except in a few major medical research centers.
Different Types of Brachytherapy Intracavitary Brachytherapy Intracavitary techniques involve placing radioactive sources into custom-designed applicators, which are placed in body cavities. This is commonly used for the treatment of gynecological tumors where the radioactive material can be placed in the uterine cavity and vagina. This is also the most common brachytherapy procedure done worldwide.
Interstitial Brachytherapy Interstitial brachytherapy involves placing the sealed radioactive sources within tissues. Thus, most nongynecological implants are performed using interstitial techniques. This is the most common brachytherapy procedure performed in the United States. The most common application is permanent implantation for treatment of prostate cancer. Less common are interstitial implants such as the Syed implants for gynecological tumors, which are more challenging to execute because large variations in
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size, shape, and location prevent them from being visualized and implanted with a uniformly applicable algorithm, like prostates.
Intraluminal Brachytherapy This is really a subclass of intracavitary brachytherapy in which the radioactive sources are inserted in the lumen of a vessel such as the blood vessel, bronchus, esophagus, or bile duct.
Surface Molds and Plaques for Brachytherapy In this technique, the radioactive sources are placed in custom-designed molds or plaques, which are then placed on the surface of the target tissue rather than being placed inside the target tissue. These techniques are not very commonly used.
Different Types of Brachytherapy Loading Systems Over the years, many techniques have been developed for handling and loading radioactive materials in order to reduce the radiation exposure of the personnel. These techniques are based upon the three cardinal principles of radiation protection, reducing time of exposure, increasing the distance from the radioactive material, and inserting shielding between the user and the radioactive material. Most of these techniques fall in one of the following categories.
Manual “Hot” Loading Techniques When brachytherapy was first introduced over a century ago, the radiotherapy sources were manually introduced into the tumor, hence subjecting the physicians and other operating room personnel to unwanted radiation exposure and the subsequent adverse effects of radiation. Direct “hot” loading is almost never used today because of its associated radiation hazards that can now be reduced by the adoption of alternative techniques.
Manual Afterloading Techniques Since 1950s or so, most brachytherapy procedures have been performed using afterloading techniques whereby hollow needles, catheters, or applicators are first inserted into the target volume. The applicators are usually inserted in the operating room. Once the position of the applicators is confirmed, the radioactive material is introduced manually into the applicator. This procedure is usually performed in the patient’s room rather than the operating room. This afterloading procedure improves placement accuracy because the clinicians can take their time in optimizing the positioning of the applicator without incurring excessive radiation exposure.
Remote Afterloading Techniques Although the manual afterloading techniques mentioned above reduce radiation exposure for the physician and the operating room staff, the nursing staff and other caregivers as well as visitors continue to be exposed to a small amount of radiation exposure. Even this small exposure can be virtually eliminated by the use of remote-controlled afterloading, in which the radioactive material is loaded into and out of the applicator by microprocessor controls positioned and operated remotely by the caregivers in an adja-
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cent room. Advances in the developments of remote afterloading techniques, especially remote HDR brachytherapy, have changed profoundly the clinical practice of brachytherapy, so that HDR has become a very popular form of brachytherapy, especially in developed countries.
Image-Guided Techniques Before the use of computerized treatment planning systems became popular, clinicians relied on generalized systems of rules (i.e., the Paterson-Parker System, the Quimby system, the Memorial system, the Paris system, the Manchester system, etc.) for pre-implant or intraoperative planning. All of these early systems depended on tables and nomograms to select number, location, and activities of radioactive sources in an implant to achieve desired dose coverage of an estimated target volume. While traditional systems are the culmination of years of clinical experience and still provide a good point of comparison with modern computerized treatment planning, the description in traditional systems of the actual dose distribution and the underlying anatomy is at best approximate and qualitative. In the last decade, two major advances have been made in brachytherapy treatment planning. One is the increased use of imagebased treatment planning and dose evaluation and the other is the development of computer-assisted dose optimization algorithms for brachytherapy. Developments in three-dimensional tomographic image-based source localization techniques have made it possible to calculate and display dose distributions directly on top of the tomographic images of involved anatomy. This enables clinical decisions to be made based on the visualization of dose distributions with respect to patient anatomy instead of dose tables. It has almost become routine in the last several years that the target volume and normal organs be localized from CT, magnetic resonance imaging (MRI), or ultrasound (US) images. Pre-implant planning can now be carried out by inspecting the instantaneous dose distribution changes resulting from different source placements. In addition, some modern analytic tools including dose volume histograms, normal tissue complication probabilities, and tumor control probabilities can be used to estimate the quality of an implant more quantitatively.
Different Types of Brachytherapy Based on Dose Rate and Duration of Brachytherapy Different Durations of Brachytherapy Typically Used Permanent Implants In a permanent implant, the radioactive sources are permanently implanted into the tumor, the patient is released from the hospital with radioactive material in him or her, and radioactivity is allowed to decay within the patient. Hence, the dose, dose rate, or the dose distribution cannot be changed after the initial insertion. Many of these procedures are relatively simple, and can be performed in an outpatient clinic and one-day surgery units. Other clinical advantages of the permanent implant are that, in deep-seated tumors, it may be safer because of the lower risk of infection and that a second operation for its removal is not required. Permanent implants are performed with relatively short half-life radioisotopes like 125I, 103 Pd, or 198Au so that the radioactivity decays to a safe level within a few weeks or months and does not present the risk of radiation-induced carcinogenesis due to long-term chronic radiation exposure. Temporary Implants In a removable implant, the radioactive material is temporarily implanted into or close to the tumor and is removed once the desired radiation dose has been delivered. When the treatment time is short (less than 20 minutes), there is better control of the total dose and the dose distribution. Removable implants with
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short treatment times require more precision compared to permanent implants. On the other hand, lower precision and accuracy can be tolerated in LDR temporary implants, because in such implants dwell locations and times may be adjusted to correct for poor needle placement. In intracavitary implants, if the dosimetry is inadequate, the dose for that fraction may be reduced and positioning or packing adjusted for subsequent fractions. If radiobiologically equivalent doses for each modality have been adopted, the decision to use either temporary or permanent implantation depends upon the judgment of the radiation oncologist and his skill and experience with a given system.
Dose Rates Used in Brachytherapy Brachytherapy can be delivered at enormously different dose rates over a very wide range of treatment times varying from minutes to months. Brachytherapy dose rates have been divided into low, medium, and high dose rates by International Commission for Radiation Units and Measurements (ICRU) Report No. 38 as follows: 1. Low Dose Rate (LDR): 0.4 to 2.0 Gy per Hour This is the traditional dose rate for permanent and manually afterloaded brachytherapy. Temporary implant LDR techniques result in typical treatment times of 3 to 5 days. This requires hospitalization of the patient. These implants are generally manually afterloaded, although LDR remote afterloaders are also available, but not commonly used. Most of the long-term clinical experience with brachytherapy is with LDR. Many of the advantages of brachytherapy are attributed to the radiobiology of continuous LDR irradiation in LDR brachytherapy. 2. Medium Dose Rate (MDR): 2 to 12 Gy per Hour This dose rate, also called “intermediate dose rate,” is rarely used since it gives excessive exposure if such an implant is manually loaded, and this dose rate does not have the advantages of outpatient brachytherapy afforded by the HDR technique. Pulsed dose rate (PDR) brachytherapy afterloaders were developed in this dose rate realm to replicate the LDR experience in terms of total treatment duration but with the source exposed in pulses for only 5 to 10 minutes per hour. 3. High-Dose-Rate (HDR): More than 12 Gy per Hour HDR brachytherapy utilizes very high activity sources, typically a 10 Ci 192Ir source, which produces a very intense radiation field around the source. Since HDR brachytherapy is associated with high radiation exposure rates, it is only used in well-shielded bunkers. Depending on distance and usage factors, 1 to 2 feet of concrete shielding or its equivalent in other materials is required. Treatment is delivered by remote-control techniques rather than manual loading. The usual dose rate in the commercially available HDR brachytherapy systems is about 100 to 300 Gy per hour, allowing the treatments to be given in only a few minutes on an outpatient basis. The introduction of HDR remote brachytherapy, with its advantages of thorough radiation protection and outpatient treatments, has led to a resurgence of interest in brachytherapy.
Chapter 2
Radiobiology: A Briefing for the Brachytherapist Marco Zaider, Ph.D. Department of Medical Physics Memorial Sloan-Kettering Cancer Center New York, New York Who Needs It? . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7 The Nuts and Bolts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8 Kinds of Effects . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8 Dose-Response Curves and Their Meaning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8 Hits: What Are They? . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 10 Temporal Aspects (I) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 10 Enhancement Ratios: RBE and OER . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 10 Cellular Proliferation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 12 Mechanisms of Radiation Action . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 13 Microdosimetry in a Nutshell . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 13 A Microdosimetric Account of the Linear-Quadratic Model . . . . . . . . . . . . . . . . . . . . . . . . . . . . 15 Temporal Aspects (II) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 17 Applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 18 Equivalent Treatments (Painless Version) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 18 Equivalent Treatments (II) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 20 Radiation Quality . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 21 The Real World . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 22 Further Reading . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 22
Who Needs It? The present function of radiation biology in treatment planning is threefold: (1) to provide information on biologically equivalent temporal patterns of dose delivery; (2) to quantify the effect of radiation quality; and (3) to guide the implementation of biologically based optimization algorithms. Of these three functions the first one is most often invoked, particularly in brachytherapy where low and high dose rate (LDR and HDR) regimens are commonly employed. The issue of radiation quality (to be defined more precisely below) is mostly relevant to brachytherapy where low-energy electrons (“low” compared to standard megavoltage radiation) are known to have enhanced biological effectiveness. The topic of biological optimization, although relatively new, is expected to dominate treatment planning in the years to come and, in fact, to replace the current dose-prescription modus operandi. The goal of this chapter is to familiarize the medical physicist with the basic language and some of the fundamental theories of quantitative radiobiology and to offer the brachytherapy practitioner the tools necessary for its implementation in daily activities. The first section explains the basic dose-response relationships, relative biological effectiveness (RBE), and enhancement factors, in particular oxygen enhancement ratio (OER). Factors that impinge on the dose-survival response, e.g., the effects of cell proliferation or radiation quality are briefly introduced. The key tools for understanding the mechanisms of radiation action and their connection to the physical description of the radiation field are provided by the discipline of microdosimetry. Elements of microdosimetry are introduced with the purpose of explaining the well-accepted linear-quadratic (LQ) formalism. An account of elementary [the effect of dose rate, biological equivalent dose (BED)] and advanced [tumor control probability (TCP), normal-tissue complication probability (NTCP)] applications concludes this chapter.
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In line with the adage “teach a man to fish and he will never go hungry,” the author could think of no better introduction to the application of these topics than letting the reader deduce a number of familiar results as applied to brachytherapy. Many important topics, as well as specific references, are absent. Because of limitations of space and time a choice was necessary and I hope that references given under “Further Reading” will stimulate the reader to complete his knowledge.
The Nuts and Bolts Kinds of Effects It is useful to recognize that radiation effects may be classified into two broad classes: stochastic and non-stochastic (or deterministic). Stochastic effects are of the all-or-none variety; for instance, cell killing. Quite obviously, radiation— or any other agent—has no effect on the nature of this sort of effect. What radiation does instead is to change the probability of producing the effect. For instance, a dose-response relationship connects probability of effect to absorbed dose. Tumor control probability (TCP), which measures the likelihood that a particular treatment dose leaves all cancer cells sterilized, illustrates this state of affairs. In contrast, deterministic effects are such that their severity (or magnitude) can be changed as a function of dosage. Cataract-induction by ionizing radiation is one such example. Effects of radiation on normal tissues are mostly of the deterministic kind—as the ability of that particular organ to function is progressively altered with increasing dose. However, one can always transform a non-stochastic effect into a stochastic one by setting up a threshold that divides the deterministic effect in question into, say, acceptable and unacceptable domains. Normal tissue complication probability (NTCP) is the result of dichotomizing the effect as described. In the following we shall denote probability of effect by E(D). The absence of effect (colloquially known as survival probability) is S(D) = 1-E(D).
Dose-Response Curves and Their Meaning Progress in molecular biology notwithstanding, the dose-response relationship remains the principal tool available to the radiation biologist for testing assumptions concerning the mechanisms responsible for radiation action. Examples of questions commonly asked are: (a) Is the effect a result of single-lesion action or a consequence of the accumulation of multiple sublesions? (b) Does the temporal pattern of dose delivery matter? (c) Does the response depend on the position of the cell in the cell cycle? (d) Is radiation-induced damage repairable? Consider the two dose-response curves shown in Figure 1 (note the semilogarithmic representation). The curve in Figure 1a is a purely exponential function, exp(−αD) where α > 0; it describes an outcome where equal increments in dose result in equal fractions of the exposed population acquiring the effect. The interpretation of this kind of response is that surviving cells1 have no memory of previous exposure
1
For convenience we shall refer throughout this chapter to “cells” as the biological units under observation. However, the statement applies equally well to any kind of biological object, from DNA to a human being.
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Figure 1a. An exponential dose response curve indicates a single-hit mechanism of radiation action.
Figure 1b. A concave downward dose effect curve shows an accumulation of sublesions and represents a multihit mode of radiation action.
to radiation. Alternatively, one speaks of single-hit mechanism in the sense that, when exposed to radiation, the cell will either show the effect or remain unaffected. The curve in Figure 1b is concave downward and depicts the LQ equation: S ( D ) = e− α D − β D , α , β > 0 . 2
(1)
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In this case sequential exposures to equal increments in dose result in an increasing fraction of affected cells. The obvious interpretation of this type of response is that cells exposed to radiation carry (and accumulate) damage not sufficient to produce the effect on its own but enough to sensitize the cell to further radiation exposure. The term sublesion will be used for this kind of damage, and lesion will denote the biological alteration responsible for the end-point under investigation. The radiation response illustrated in Figure 1b is known as multi-hit action. Finally, if—as a result of exposure to radiation—cells become less sensitive, one expects a doseresponse curve, which is concave upwards. Mathematically, the function that determines the behavior of the curves shown in Figure 1 is2:
h( D) = −
d log ( S ) dD
constant ( curve A ) = monotonicallly increasing ( curve B) monotonically decreasing .
(2)
Hits: What are They? We have spoken of single-hit and multi-hit action and at this time we need a precise definition of the term “hit.” Loosely speaking, a hit (microdosimetrists prefer the name event) indicates the traversal of the radiation-sensitive region of the system under observation by an ionizing particle (Figure 2). More precisely, a hit (or event) is the production of statistically correlated alterations in the sensitive site. Essentially, alterations result from local energy deposition via ionizations or excitations. The crucial feature here is that events are statistically independent.
Temporal Aspects (I) If cells respond to radiation according to the single-hit mechanism, the temporal distribution of hits is clearly a matter of indifference. However, in multi-hit mode and in the presence of repair mechanisms sublesions may be eliminated before the next hit arrives and thus dose rate (proportional to the number of hits per unit time) becomes relevant. As the dose rate is progressively lowered, the term responsible for multi-hit behavior [βD2 in equation (1)] must grow increasingly smaller; at the limit of very low dose rates the dose response curve becomes exponential (Figure 3). (A more precise definition of “low” dose rate will be given below; however, it should be apparent that this quantity depends on the characteristic sublesion repair time, which is of the order of one hour.)
Enhancement Ratios: RBE and OER The shape of the dose-response curve depends on many factors (radiation quality, oxygen concentration, position in the cell cycle, etc.). The enhancement ratio (ER) quantifies such differences as follows: let DA and DB be isoeffective doses corresponding, respectively, to dose-response curves A and B. Then: ER =
2
In epidemiology h is known as the hazard rate.
DA DB
.
(3)
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Figure 2. The dots indicate the locations of energy deposition (ionizations or excitations) by a hypothetical charged-particle track in tissue. Also shown are two cells—one traversed by the main track and the other by a secondary electron (delta ray). The ensemble of energy deposition sites within the radiosensitive volume, which are associated with one track, defines a (microdosimetric) event.
Figure 3. As the dose rate decreases the quadratic term (βD2) becomes increasingly smaller. At very low dose rates only the linear term, αD, remains (dashed curve). Lower solid curve: acute exposure; upper dashed curve: protracted exposure.
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The following are two important examples: Relative biological effectiveness (RBE) is a quantity used to compare the biological effects of two different radiations. If X denotes the reference radiation, the RBE of A is given by: RBE =
DX DA
.
(4)
Typically, high-energy electrons (or, equivalently, high-energy photons) are taken as reference radiation. Anoxic cells (poor in oxygen) are more resistant to radiation than aerobic (oxic) cells. This phenomenon is known as oxygen effect. To quantify the oxygen effect, one makes use of the oxygen enhancement ratio (OER) defined as: OER =
Danoxic Doxic
.
(5)
In general, the ER depends on the dose level [e.g.. DA in equation (4)]. An agent for which the ER does not depend on dose is known as a dose-modifying factor. (The oxygen is believed to be a dosemodifying factor.)
Exercises 1. Find an expression for RBE for two radiations characterized by (αA, β) and (αX, β): see equation (1). 2. Plot RBE as a function of DA for the result obtained above. 3. Find the limiting value of RBE when DA → ∞, and when DA → 0. 4. Show that when β = 0, RBE does not depend on dose. 5. Find a linear-quadratic expression for an agent, which is a dose-modifying factor.
Cellular Proliferation Stem cells (responsible for recovery of self-renewing tissues from radiation injury) and malignant cells (which account for tumor growth) have the capability to proliferate. Knowledge of the growth curve of these cell populations is of obvious importance in treatment planning. Additionally, one must be aware that radiation response depends drastically on the cell position in the cell cycle. We shall start by a brief review of this latter aspect. The cell cycle of actively dividing cells is conventionally divided into four stages: G1, S, G2, and M. G1 and G2 (known as first and second gaps) bracket the S phase (where DNA synthesis takes place) and the M phase—where the actual cell division occurs. Empirical evidence shows that (1) cells are most radiosensitive in M phase and most radioresistant in the latter part of S; (2) exposure to radiation results in a lengthening of the G2 phase (this is known as the G2 block); and (3) exposure to radiation decreases the division probability.
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The kinetics of cell growth can be described in terms of three quantities: (1) the average duration of the cell cycle, Tc; (2) the growth fraction, GF, which is the fraction of cells actively proliferating; and (3) the cell loss factor, φ, which represents the rate of cell loss divided by the rate of cell production. With this, the cell number doubling time, Td, is given by: Td = Tc
1
ln 2
1 − ϕ ln (1 + GF )
.
(6)
In particular, if no cell loss occurs (φ=0), the doubling time is known as Tpot, which stands for potential doubling time. Tpot = Tc
ln 2 ln (1 + GF )
(7)
Another useful relation is:
ϕ = 1−
Tpot Td
.
(8)
For a typical human tumor Tc ≈ 1−2 d, φ ≈ 0.8−0.9, and GF ≈ 0.4−0.5. The expressions, equations (6 to 8), will be useful for calculating TCP.
Mechanisms of Radiation Action Microdosimetry in a Nutshell The energy deposited by ionizing radiation in any given site is a stochastic quantity and absorbed dose— the quantity commonly used to measure the “amount of radiation”—is in fact only the average value of its distribution. The response of an organism (cell, virus, etc.) to radiation depends on the energy locally deposited and, understandably, bears no relation to its expected value over an ensemble of such individuals. The belief that the average effect will depend on dose (average energy per unit mass) is not borne out by experiment either; it is well established that the same dose of radiation may result in vastly different effects, when delivered by different types of radiation. The reason for this state of affairs can be found in the observation that generally the biological effects of radiation depend on the (highly non-uniform) macroscopic pattern of energy deposition. Microdosimetry is the systematic study and quantification of the spatial and temporal distribution of energy in irradiated matter. The principal microdosimetric quantity is specific energy, z, defined as the quotient of ε by m, where ε is the energy imparted by ionizing radiation to matter of mass m: z=
ε
.
(9)
m
The unit of specific energy is the gray (Gy), equal to 1 J/kg. Specific energy z can be imparted to matter in single events or in multiple events; the respective distributions are labeled f1(z) and f(z). Note that f(z)dz—i.e., the probability that specific energy is in the interval dz about z—depends on the geometry and physical composition of the site in which z is determined. In contrast, absorbed dose, D, is defined
14
Marco Zaider
at a point in matter through the following limiting process. D = lim z
(10)
m→ 0 ρ = const
where is the first moment of f(z) and the limit is taken at constant mass density, ρ. When the site is sufficiently small (almost invariantly the sizes of interest are of the order of micrometer) D = z . Two averages of the single-event distribution, zF (the frequency average) and zD (the dose average) are defined as follows: ∞
zF = ∫ zf1dz ;
1
zD =
zF
0
∞
∫z
2
f1dz .
(11)
0
For a given dose, D, the average number of events traversing the site is n=
D
.
(12)
zF
Alternatively: ∞
n = ∑ k P ( k,; n )
(13)
k=0
where P(k;n)—the probability of exactly k events—is the Poisson distribution: P ( k; n ) = e n
nk . k!
(14)
Let fk(z) denote the distribution of z in exactly k events. It follows that: f0 ( z ) = δ ( z )
f1 ( z ) ≡ f1 ( z ) f2 ( z ) =
∞
∫ f ( z ') f ( z − z ') dz ' 1
1
0
(15)
fk ( z ) =
∞
∫ f ( z ') f ( z − z ') dz ' 1
k −1
0
and thus: ∞
f ( z ) = ∑ p ( k; n ) k=0
fk ( z ) .
(16)
2–Radiobiology: A Briefing for the Brachytherapist
15
The expression, equation (16), shows that—given the single-event distribution f1(z)—one can calculate f(z) for any dose, D (or, equivalently, n).
Exercises 1. Let Φ(ω,n) and Φ1(ω) be the Fourier transforms of f(z) and f1(z), respectively. Prove, using the convolution theorem that Φ (ω , n ) = exp ( − n ) exp [ nΦ1 (ω )] . 2. Let mk and Mk represent, respectively, the k-th moments of f1(z) and f(y). Show that: k −1
mk + 1 = M k + 1 / n − ∑ C jk M k − j m j + 1 . j=0
From the result obtained at Exercise 2, it follows that: M 1 = z = nzF M 2 = z 2 = z ( z + zD ) ,
(17)
where [from equation (11)] zF = m1 and zD = m2/m1. If D = , equation (17) becomes: z 2 = zD D + D 2 .
(18)
A Microdosimetric Account of the Linear-Quadratic Model The LQ expression for the survival probability, equation (1), is understood to stand for the probability that following exposure to dose D, no lesions were produced in the organism. If the actual number of lesions (ε, a stochastic quantity) is Poisson-distributed3, it follows that
ε ( D) = α D + β D2
(19)
is the average number of lesions. Moreover, the LQ dependency of the yield of lesions on dose is compatible with the assumption that each lesion is the product of two sublesions. To make this clear, consider the “prototype” lesion responsible for preventing cell proliferation4, namely the dicentric aberration (Figure 4). A dicentric is the result of two (unrepaired) chromosome breaks misjoining their split ends. A cell that has a dicentric in it at the time of mitosis will not be able to go through the normal process of division. The two single-chromosome breaks are the sublesions and the dicentric is the lesion. Each sublesion is the result of one particle (event, or hit) traversing a chromosome. The pair of sublesions needed to produce a dicentric can be generated in two different ways: (1) by a single particle traversing both chromosomes, or (2) by two (independent) particles. The yield of the former mechanism is proportional
3
The distribution of lesions is not Poissonian; however, for low-LET (linear energy transfer) radiation (energetic electrons or photons) the Poisson distribution is a fair approximation. 4 Radiation biologists refer to this end point as “cell killing” although the cell is not “killed” in any sense of the word.
16
Marco Zaider
Figure 4. A dicentric chromosomal aberration results when two adjacent single-chromosome breaks (sublesions) rejoin to produce a chromosome with two centromeres (a lethal lesion). In what is termed “intra-track” event, a single particle produces both breaks and the yield of lesions thus induced is proportional to the number of particles, or equivalently to dose (upper panel). In the other possible mechanism (inter-track event) two independent tracks produce each one chromosome break. The yield of dicentrics thus produced is proportional to D2 (lower panel).
to the number of particle traversals, that is to dose, D; while the latter will depend—for similar reasons— quadratically on dose. The result is equation (19). Microdosimetry offers a physical explanation of the constants α and β in equation (19). In line with the interpretation given above, assume that the yield of lesions (alterations responsible for the end point observed) depends quadratically on specific energy: E ( z) = β z2.
(20)
On average [see equation (18)]:
(
)
E ( D ) = β z 2 = β zD D + D 2 .
(21)
Thus: zD =
α β
.
(22)
To summarize: β is proportional to the yield of sublesions (e.g., single-chromosome breaks) per unit specific energy; and the ratio α/β, a purely physical quantity depends on radiation quality and on the distribution of radiosensitive matter in the cell.
2–Radiobiology: A Briefing for the Brachytherapist
17
Temporal Aspects (II) To account for sublesion repair—a factor that affects the quadratic term, βD2—the LQ expression is modified as follows:
ε ( D ) = α D + q ( t ) β D 2.
(23)
The dose-rate function, q(t), quantifies sublesion damage repair that occurs in-between events. It is calculated as follows: ∞
∫ τ (t ) h (t ) dt ,
q=
(24)
0
where τ(t) = exp(−t/t0) gives the rate of sublesion damage elimination as a function of time, t0 is the sublethal repair constant, and h(t)dt is the distribution of time intervals t between consecutive events (hits). Thus: h (t ) =
2 D
2
∞
∫ I ( s ) h ( s + t ) ds .
(25)
0
In this expression I(s) is the dose rate as a function of time. As expected, for acute (HDR) exposures q → 1; for protracted exposures q → 0. For instance: (a) For irradiations at a constant dose rate: q(r ) =
2 r
−
2 f2
(1 − e ) −r
(26)
where r = t/t0. In particular, when t >> t0 : q≅
2t 0
.
(27)
t
(b) For f well-separated fractions (complete sublesion repair in between fractions): q=
1 f
.
(28)
In a typical LDR treatment in brachytherapy the total irradiation time is of the order of several days and therefore q ª 0. It follows that the probability of cell survival, S(D), is quasi-exponential and the RBE is determined by the linear (α) coefficient—or in microdosimetric terms, by zD. In HDR brachytherapy the dose rate is of the order of 1.5 Gy/min; for a prescription dose of, say, 6 Gy and taking t0 = 60 min one obtains q = 0.978 and thus both terms, linear (αD) and quadratic (βD2), contribute significantly.
18
Marco Zaider
Exercises 1. Demonstrate equation (26). 2. Demonstrate equation (27). 3. Demonstrate equation (28).
Applications In this section we discuss some simple applications of the items described above. For numerical estimations the following may come in handy: Typical initial dose rates in brachytherapy are 0.07 Gy/h (125I) and 0.24 Gy/h (103Pd) for permanent implants; 0.5 Gy/h for interstitial implants; 50 to 100 Gy/h for socalled “HDR” treatments. It is common to take α/β = 10 Gy for early responding tissues (tumors and tissues with fast turnover) and α/β = 3 Gy for late responding tissues. For the sublethal repair constant t0 of equation (24), a value of 0.5–1 h would be representative for most cell lines. The simplest (but not necessarily accurate) way to account for proliferation during time t is through a term exp(t/Tpot).5 Since only the quadratic term in dose (qβD2) is affected by temporal effects, the terminology “low” or “high” dose rate must be understood in terms of the relative contribution of this term to the total yield of lesions. Thus, if the dose rate is such that αD >> qβD2, one is in low dose rate regime. Note that the lower the dose, the larger the range of dose rates that classify as “low.”
Exercise Using α = 0.1 Gy−1, β = 0.01 Gy−2, and t0 = 1 h, find—as a function of dose—the range of dose rates that qualify as “low” if one defines low dose rate by the condition 0.1 αD = qβD2.
Equivalent Treatments (Painless Version) Two treatments [different with respect to either one or any combination of the following: radiation quality and type of tissue (α, β), dose rate, Tpot, sublesion repair constant (t0)] can be made equivalent by changing the total dose, the dose rate, and/or the treatment time. It is understood that the temporal pattern of dose delivery is given by prescription. However, one must ask first “equivalent in terms of what?” The two choices are equivalent TCP or equivalent NTCP. It is generally assumed (incorrectly, in fact) that for both TCP and NTCP treatments that result in equal survival probabilities for the cells that make up the respective tissues cause equivalent biological outcome. For instance, in a tumor containing n malignant cells treated uniformly to dose D, the average number of surviving cells is nS(D). If the number of surviving cells is Poisson-distributed (this is true only under certain conditions), the probability that no cell will survive the treatment is: TCP = e− nS ( D ) ,
(29)
and in this case iso-survival probability is indeed equivalent to iso-TCP. The case for NTCP is less clear and the assumption mentioned above (equal survival means equal normal tissue damage) is by and large a matter of faith. 5
See, however, the section herein Equivalent Treatments (II).
2–Radiobiology: A Briefing for the Brachytherapist
19
Under these conditions to find biologically equivalent treatments, one would use the following expression:
α D + β qD 2 −
t Tpot
= α 1 D1 + β1q1 D12 −
t1 Tpot ,1
.
(30)
This equation must be applied separately to early- (e.g., tumor) and late-responding tissues; and if the dose is designed to match, say, the effect on the tumor, one must evaluate if the effects on normal tissues will increase or decrease. A worked example [In the following we shall use α/βearly = 10 Gy, α/βlate = 3 Gy, Tpot = 2 d (early) or 60 d (late), β = 0.01 Gy−2, t0 = 1 h). 1. A tissue is known to tolerate N = 30 daily fractions of 1.8 Gy each (total dose, D). What is the equivalent total dose, D1, required if the treatment is delivered in one single fraction over 7 days? (ignore cell proliferation) Solution:
2t 1 2 D = α D1 + β 0 D12 t N
αD + β
α 2t 0 2 α 1 2 D1 + D1 − D + D = 0 N β t β 2
2t α α α 1 − + + 4 0 D + D2 N β t β β D1 = 4t0 t For early responding tissues D1 = 59.5 Gy. For late-responding tissues D1 = 68.0 Gy. 2. If the dose is designed to match the tumor tissue, will the effect on late-responding tissues increase or decrease? 3. Repeat the calculation above by including cell proliferation. Now (with obvious notations and ∆t = 1 day):
20
Marco Zaider
S( D) = e
2
− α D − β qD + t / T pot
2t t 1 2 t D − = α D1 + β 0 D12 − 1 N t2 T T1
αD + β
2t 0 2 α α 1 2 N ∆t t1 t D1 + β D1 − β D + N D − T β + T β = 0 1 1 2
− D1 =
N ∆t t 2t α α α 1 + 1 + + 4 0 D + D2 − N β T β T1 B t1 β β
4t0 t 1
Answer again question 2, but for this case.
Equivalent Treatments (II) For any temporal protocol of dose delivery, a general formulation that describes the probability of tumor control—defined here as the probability that at time t there are no clonogens alive—is:
S ( t ) e(b − d ) t TCP ( t n, S ) = 1 − t dt ' (b − d )t 1 + bS t e ( ) ∫ S ( t ' ) e(b − d ) t ' 0
n
.
(31)
In equation (31) S(t) is the survival probability (proliferation processes excluded) at time t of the n clonogenic tumor cells present at the time treatment started (t = 0), and b and d are, respectively, the birth and (radiation-independent) death rates of these cells. Equivalently, b = 0.693/Tpot and d/b is the cell loss factor, φ, of the tumor. In this expression t refers to any time during or after the treatment; typically, one would take for t the end of the treatment period or (better) the expected remaining life span of the patient. [Unlike equation (29) the expression, equation (31), is valid for both temporary and permanent implants.] The quantity: Seffective =
S ( t ) e(b − d ) t t
1 + bS ( t ) e(b − d ) t ∫ 0
dt '
(32)
S ( t ') e
(b − d )t '
in equation (31) represents the effective survival probability (proliferation included) of a cell present in the system at time t. In the same way, one may define a biologically equivalent dose (BED) by observing that
2–Radiobiology: A Briefing for the Brachytherapist
BED = −
1
α
21
S ( t ) e(b − d ) t
log
1 + bS ( t ) e
(b − d )t
(33)
t
dt '
∫ S (t ') e(
b − d )t '
0
has a dimension of dose. In the absence of cell proliferation (b = d = 0) BED takes the more familiar form:
β BED = D 1 + q ( t ) D . α
(34)
For a fractionated treatment (f fractions separated by time ∆1, S0 = survival probability after a single fraction), the probability of cure—defined as the probability that as time post-treatment tends to infinity there will be no clonogenic tumor cells left—is given by the following expression:
f λ∆ S0 e ) ( TCP = 1 − λ∆ λ ∆ ( f − 1) − 1) b eλ ∆ 1 + 1 − e− λ ∆ S0 e ( S0 e ) ( λ1 S0 e λ ∆ − 1 1
1
1
1
1
1
n
1
1
(35)
1
1
1
1
(λ1 = b-d) or, within a very good approximation:
f λ∆ S0 e ) ( n TCP ≈ exp − . λ∆ λ ∆ ( f − 1) − 1) b eλ1 ∆1 1 + 1 − e− λ ∆ S0 e ( S0 e ) ( λ1 S0 e λ ∆ − 1 1
1
1
1
1
1
(36)
1
1
1
1
Radiation Quality The photon or electron energies used in brachytherapy are generally low. At low doses or low dose rates the RBE of 100 keV photons relative to, say, 1-MeV photons may exceed 2. For instance, the RBE for 125 I has been extensively studied and results of 1.2 to 2 have been reported for the dose-rate range of 0.03 to 9 Gy/h. Similarly, for 103Pd a study performed at 0.07 to 0.8 Gy/h reported an RBE value of 1.9. While considerably lower RBE values apply to the much higher doses or dose rates usually employed in external radiotherapy, differences in biological effectiveness of the order of 10% to 15% remain. In a typical LDR treatment in brachytherapy the total irradiation time is of the order of several days and therefore q ª 0. It follows that the probability of cell survival, S(D), is quasi-exponential and the RBE is determined by the linear (α) coefficient—or in microdosimetric terms, by zD. The expression, equation (22), links the ratio α/β to the microdosimetric quantity zD and because β changes very little with electron energy one can assume that—on a relative scale—the coefficient α changes in the same manner. The table below gives numerical examples of ratios RBE = zD/zD(60Co) for radiations of interest in brachytherapy.
22
Marco Zaider
zD/zD(60Co)
Radionuclide 103
Pd
2.3
125
I
2.1
Am
2.1
241
192
Ir
1.3
60
Co
1.0
The Real World The treatment presented here disregards many important factors present in real-life clinical situations and the unwary should not be carried away by mathematical enthusiasm. Here are four examples: 1. It is rarely the case, and it never happens in brachytherapy, that any given treatment area (tumor or healthy organ) is uniformly exposed to the same dose. To the extent that individual cells (or functionally linked groups of cells) respond independently to radiation the device known as dosevolume histogram (DVH) can be used to take this into account. For instance, the TCP expression, equation (31) becomes: N
TCP ( t ) = ∏ TCP ( t ni , Si ) ,
(37)
i =1
where ni tumor cells have the same (dose-dependent) survival probability, Si. 2. Tissues (malignant or not) consist of cells that have a wide spectrum of α, β, Tpot, and φ values. Theoretical considerations indicate that—unless one particular cell type numerically dominates the tissue in question—tumor response and normal-tissue toxicity is primarily determined by a (possibly) small number of cells—radioresistant, fast-proliferating in the case of TCP, or radiosensitive for NTCP. Empirical evidence for this conjecture comes from the fact that, quite invariably, analyses of TCP dose-response data indicate that the number of relevant cells [n in equation (31)] is remarkably small. 3. As actively proliferating cells progress through the mitotic cycle, they change radiosensitivity parameters by as many as one to two orders of magnitude. 4. Cells do not generally exist in isolation and their response to radiation may depend on many extracellular factors and fluctuating environments. Nevertheless, one should remember that crude though such mathematical models may be in terms of biological realism, their limitations fade into insignificance when compared with the clinical rewards that can result from increased knowledge of the mechanisms responsible for radiation action.
Further Reading Brahme, A., J. Nilsson, and D. Belkic. (2001). “Biologically optimized radiation therapy.” Acta Oncol 40:725–734. Dale, R. G. (1985). “The application of the linear-quadratic dose-effect equation to fractionated and protracted radiotherapy.” Br J Radiol 58:515–528. Fowler, J. F. (1989). “The linear-quadratic formula and progress in fractionated radiotherapy.” Br J Radiol 62:679–694.
2–Radiobiology: A Briefing for the Brachytherapist
23
Hall, E. J. Radiobiology for the Radiologist. New York: Harper and Row Publishers, 1994. Kellerer, A. M., and H. H. Rossi. (1978). “Generalized formulation of dual radiation action.” Radiat Res 75:471–488. Kiefer, J. Biological Radiation Effects. New York: Springer-Verlag, 1990. Rossi, H. H., and M. Zaider. Microdosimetry and Its Applications. Springer-Verlag Telos, 1996. Steel, G. Gordon. Growth Kinetics of Tumors. Gloucestershire, UK: Clarendon Press, 1977. Steel, G. Gordon (ed.). Basic Clinical Radiobiology. Oxford: Oxford University Press, 1997. Thames, H. D., and J. H. Hendry. Fractionation in Radiotherapy. Bristol, PA: Taylor and Francis, 1987. Zaider, M., and G. N. Minerbo. (2000). “Tumor control probability: A formulation applicable to any temporal protocol of dose delivery.” Phys Med Biol 45:279–293. Zaider, M., and H. H. Rossi. Radiation Science for Physicians and Public Health Workers. Norwell, MA: Kluwer Academic Publishers, 2001.
Chapter 3
Sources and Delivery Systems I: Radionuclides Ravinder Nath, Ph.D. Department of Therapeutic Radiology Yale University School of Medicine New Haven, Connecticut Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 25 Photon-Emitting Radionuclides Used in Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 25 226 Ra . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 25 137 Cs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 26 192 Ir . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 26 198 Au . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 27 125 I . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 27 103 Pd . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 28 Beta-Emitting Radionuclides Used in Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 28 90 Sr . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 28 32 P . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 28 90 Y . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 28 Neutron-Emitting Radionuclides Used in Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 29 252 Cf . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 29 Other Radionuclides with Potential Applications in Brachytherapy . . . . . . . . . . . . . . . . . . . . . . 29 Acknowledgments . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 30 Further Reading . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 30
Introduction In the early applications of brachytherapy, naturally occurring radionuclides of radium and radon were used extensively. As the artificially produced radionuclides became available from nuclear reactors and particle accelerators, many new radionuclides entered the brachytherapy field. Today in medicine, there is an extensive menu of radionuclides emitting gammas, betas, or neutrons with wide spectra of energies and half-lives. For brachytherapy, these radionuclides are encapsulated in sealed, biocompatible capsules made of materials such as titanium, stainless steel, etc. The physical characteristics of some of the commonly used radionuclides in brachytherapy are described in this chapter. Details of the fabrication and composition of actual brachytherapy sources in use are described in the next chapter.
Photon-Emitting Radionuclides Used in Brachytherapy 226
Ra
226
Ra is the sixth member of the uranium decay series that starts with 238U and ends with stable 206Pb. Radium decays into radon, with a half-life of about 1622 years; radon is a heavy, inert gas that in turn disintegrates into its daughter products. As a result of the decaying processes from 226Ra to 206Pb, at least 49 photons are emitted, with energies ranging from 0.184 to 2.45 MeV. The average energy of the gamma rays from radium in equilibrium with its daughter products and filtered by 0.5 mm of platinum is 0.83 MeV. There are also some high-energy beta particles and alpha particles emitted by this radionuclide that are absorbed by the encapsulation material. The half-value layer (HVL) of the gamma rays from an encapsulated source containing this radionuclide is about 14 mm of lead. 226 Ra and 222Rn are virtually unused today, primarily because of the hazards of chemical and radioactive toxicity of radium and its by-products and because of the high energy of photons emitted by these
25
26
Ravinder Nath
sources. The high energy of the photons makes it difficult to shield health professionals and others from unwanted radiation exposure. For these reasons, several other radionuclides (as described below), which emit lower-energy photons, have been introduced. In the early radium sources, the radioactive material was supplied mostly in the chemical form of radium sulfate or radium chloride mixed with an inert filler and loaded into cells about 1 cm long and 1 mm in diameter. A typical radium source might contain 1 to 3 cells, depending on the source length. Radium sources were manufactured as needles or tubes in a variety of lengths and activities. 137
Cs
137
Cs is a by-product of nuclear fission and is generated in nuclear reactors. It has a half-life of about 30 years and decays through beta emission (93.5%) to the metastable state of 137Ba (half-life 2.5 minutes) and then through gamma-ray emissions to its ground state (Figure 1). The gamma rays emitted have a photon energy of 662 keV. The HVL for the gamma rays from this radionuclide is about 5.5 mm of lead. For 137Cs sources, the radioactive material is supplied in the form of insoluble powders or ceramic microspheres labeled with cesium and doubly encapsulated in stainless steel for both needles and tubes. The beta particles and low-energy characteristic x-rays are totally absorbed by the stainless steel encapsulation, making the clinical source a pure gamma emitter. The active length of the tube sources typically used is 10 to 50 mm and diameters are approximately 1.5 to 2.0 mm. 137 Cs sources are less hazardous and require less shielding compared to that required for 226Ra sources. Although the half-life of 137Cs is much less than that of 226Ra, some of the cesium sources may have to be replaced after about 7 years. In clinical use, the activity of the sources must be adjusted for the decay of 137 Cs over time. 137Cs sources have been used in both interstitial and intracavitary brachytherapy. 192
Ir
192
Ir is produced in a nuclear reactor via neutron capture by stable 191Ir. 192Ir has a half-life of 73.83 days and decays primarily by beta emission and electron capture to excited states of 192Pt and 192Os. Subsequently, the daughters decay to the ground states by gamma-ray emission. 192Ir has a very complicated gamma-ray spectrum. The average energy of gamma rays is about 0.38 MeV. Because the gamma-ray energy is lower than 226Ra or 137Cs7, 192Ir sources require less shielding. The HVL thickness for this radionuclide is about 2.5 mm. 192 Ir sources are available in the form of small sources (about 3 mm long with a diameter of about 0.5 mm) placed in nylon ribbons for safety purposes. Two different source designs are commercially available in the United States. One of them has an inner core alloy composed of 30% iridium and 70% platinum, encapsulated in stainless steel. The other has an inner core alloy composed of 10% iridium and 90% platinum surrounded by a 0.1-mm thick cladding of platinum. Typically, these sources are press-fitted,
Figure 1. Decay scheme for 137Cs.
3–Sources and Delivery Systems I: Radionuclides
27
normally at 1-cm intervals, into nylon ribbons with an outer diameter of about 0.8 mm. 192Ir sources are also available as wires composed of an alloy of 25% iridium and 75% platinum enveloped in pure platinum. The wires are available in two sizes with outer diameters of 0.3 mm and 0.5 mm. These are more popular in Europe than in the United States. 192Ir wires and sources in ribbons are particularly suitable for afterloading technique. These sources are used for temporary brachytherapy implants only, because their high average photon energy makes them unsuitable for permanent implants due to excessive public radiation exposure. High-intensity 192Ir sources are also available for the high dose rate (HDR) remote afterloader units. Very high activity sources (typically 10 Ci) of 192Ir can be fabricated in a very small volume source because of the high specific activity of this radionuclide. 192 Ir is the only gamma isotope that has been clinically used for intravascular brachytherapy (IVB) thus far. In one of the IVB systems, 192Ir sources were used as trains of sources in nylon ribbons containing 6, 10, 14, 18, or 22 seeds with 1 mm seed spacing (i.e., 4 mm center to center seed spacing). Seed activities were typically 33 mCi per seed. Treatment times for IVB were typically 15 to 40 minutes to deliver 15 Gy at a distance of 2 mm from the source center. A second 192Ir system for IVB consisted of a wire source (192Ir source inside a 0.4 mm diameter nitinol wire) rather than seeds. 198
Au
198
Au is produced in a reactor by bombarding a gold target with neutrons. 198Au has a half-life of 2.7 days and 99.9% of its gamma rays are emitted with an energy of 0.412 MeV; its primary beta emission has a maximum energy of 0.96 MeV. 198Au has three gammas (412, 676, and 1088 keV) and three betas. The nuclide may also undergo an isomeric transition (five prominent gammas from 100 to 334 keV) prior to beta decay. A typical gold seed, also known as a gold “grain,” is encapsulated in 0.1 mm of platinum, which is sufficient to absorb the beta rays emitted by 198Au. The outside dimensions of a 198Au source are 2.5 mm long and 0.8 mm in diameter. Because of their short half-life, 198Au seeds are used in permanent implants only. Although commonly used in Canada, the 198Au sources have been used only sparsely in the United States. 125
I
125
I has a half-life of 59.6 days decaying exclusively by electron capture process to an excited state of 125Te followed by spontaneous decay to the ground state with the emission of 35.5 keV gamma rays. Characteristic x-rays in the range of 27 to 32 keV are also emitted as a consequence of the electron capture and internal conversion processes. One of the advantages of this radionuclide over radon and gold is the lower energy of the photons that it emits, which makes reducing radiation exposures around the patient much easier. The HVL thickness for the photons emitted by encapsulated sources containing this radionuclide is about 0.025 mm of lead. Two models of the 125I seeds were originally introduced in the 1960s. In one of the models, the 125I was adsorbed onto a 3-mm long silver rod that is the central core of the source. The rod was encapsulated in 0.5 mm of titanium; the resultant seed is 4.5 mm long and 0.8 mm in diameter. In the other original model, the seeds consisted of 2 to 4 ion exchange resin spheres containing 125I inside a titanium tube. Titanium encapsulation served to absorb low-energy electrons and x-rays. The ion exchange resin beads were impregnated with 125I in the form of iodide ion. This technique offered an efficient absorption of 125I, which allows fabrication of higher activity sources than is possible with the other model. Except for their use in 125 I sources in eye plaques for temporary brachytherapy, 125I seeds are used principally for permanent implants. Seeds are available in strengths up to 50 U. In the last decade, numerous other models of 125I sources have become commercially available, as described further in the next chapter.
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Ravinder Nath
103
Pd
103
Pd can be produced in a nuclear reactor when stable 102Pd captures a thermal neutron. It can also be produced in a cyclotron by bombarding protons into a rhodium target. With a half-life of 17 days, it decays via electron capture process with the emission of characteristic x-rays in the energy range of 20 to 23 keV and Auger electrons. The weighted mean photon energy is 20.7 keV, and the HVL of photons emitted from encapsulated sources of 103Pd is about 0.004 mm of lead. Original 103Pd sources consisted of a cylindrical titanium tube sealed at both ends with laser welded titanium end cups. Enclosed in the tube were two 103Pd graphite cylinders and a lead rod x-ray marker for radiographic identification. The cylindrical tube was about 5 mm long and 0.8 mm in diameter. In the last decade, numerous other models of 103Pd sources have become commercially available, as described further in the next chapter. 103 Pd seeds are primarily used for permanent implants. However, high-activity 103Pd sources have been fabricated which make it possible to use 103Pd for temporary brachytherapy applications.
Beta-Emitting Radionuclides Used in Brachytherapy 90
Sr
90
Sr decays with a half-life of 28.9 years to 90Y and yields only beta rays with a maximum energy of 0.5 MeV. The daughter isotope 90Y is a nearly pure beta emitter (99.99% of decay events) with a 64-hour halflife, and emits betas with a maximum energy of 2.27 MeV. Like 90Sr, 90Y decays solely by beta emission, but a tiny fraction of decays (0.01%) result in one of two excited levels of the daughter isotope, which then relaxes by emission of low-intensity gamma rays. It is primarily the 90Y betas that are useful for therapy. This radionuclide is a beta-ray emitter suitable for treatment of superficial lesions, e.g., a 90Sr ophthalmic applicator is used for treatment of lesions in the eye where the depth of penetration needed is a few millimeters. The second most common application for 90Sr is IVB for prevention of restenosis in coronary or peripheral arteries. One of the IVB systems consists of a string (“train”) of 90Sr seeds. Each seed is encapsulated in stainless steel, and the activity is contained in a ceramic matrix. The seeds are stored in a hand-held delivery device and are advanced by a closed loop hydraulic system, which uses sterile saline to advance (and then retract) the seeds into and out of a non-centered catheter. 32
P
32
P is a pure beta emitter with 14-day half-life, and betas with a maximum energy of 1.71 MeV. Recently, it has been used in a commercial IVB system, which consisted of a 32P wire source. In one version, the delivery system consisted of a sophisticated computer-controlled stepping motor, similar to a conventional HDR unit, and a spiral-shaped perfusion-centering catheter. The source itself contained 32P hermetically sealed inside a 0.5 mm diameter nitinol wire. Another IVB system utilizing 32P consisted of an angioplasty balloon impregnated with 32P. 90
Y
90
Y is the daughter product of the decay of 90Sr and is a beta emitter with a 64-hour half-life, and 2.27 MeV maximum beta energy. 90Y has been tested as a stand-alone source (i.e., not in equilibrium with 90Sr) in clinical trial for IVB. In this trial the 90Y source consisted of a 0.1 mm diameter 90Y wire coated with
3–Sources and Delivery Systems I: Radionuclides
29
titanium with a computer-controlled HDR-like delivery device and a non-perfusion, segmented balloon centering system.
Neutron-Emitting Radionuclides Used in Brachytherapy 252
Cf
252
Cf has a half-life of 2.65 years decaying via alpha emission (97%) and nuclear fission (3%). The fission decay path emits neutrons with a spectrum of neutron energies similar to that of a fission reactor and with an average energy of 2.15 MeV. A significant number of gamma rays with an average energy of 0.7 to 0.9 MeV are also emitted from the fission events and from the decay of fission products. One of the clinical 252Cf sources has a central core consisting of a ceramic metal mixture of 252Cf oxide and palladium. This core is encapsulated with platinum-iridium alloy (10% iridium) to limit the combined gamma-ray and beta-ray dose to one-third the effective neutron dose. The seed is about 6 mm long and 0.8 mm in diameter. 252Cf sources are intended for use in temporary implants only because of the high risk of radiation-induced carcinogenesis from chronic low-dose irradiation by neutrons.
Other Radionuclides with Potential Applications in Brachytherapy Other radionuclides that have been used in brachytherapy include 60Co, 182Ta, 106Ru, 106Rh, 241Am, 145Sm, and 169Yb. The long half-life of 241Am and its intermediate photon energy of about 60 keV indicate that this radionuclide may be well suited for intracavitary use in which custom-designed shielding may be very effective. 169Yb has similar average photon energy but a much shorter half-life, which would make it commercially attractive as an HDR source. 145Sm has a half-life of 340 days and photon energies in the range of 38 to 45 keV. With its low energy and relatively long half-life, 145Sm is being considered as a substitute isotope for removable interstitial implants. Several other radionuclides have also seen limited clinical testing, but do not appear to be likely candidates for immediate commercial development. One of these radionuclides is 188W (69.4-day half-life, 349 keV maximum beta energy) in equilibrium with its short-lived daughter 188Re (17-hour half-life, 2.12 MeV maximum beta energy). The 188Re beta decay is accompanied by 155 keV gammas (10%) plus another 2% of gammas with energies up to 932 keV. In a version of the IVB system, 50 to 100 mCi of liquid 188Re were “milked” from the generator for use in the radioactive liquid balloon delivery system for IVB. Another version of this system, 188Re in a sealed, catheter-based 188W/188Re source has been investigated for IVB. In this case, the source comprised a neutron-activated tungsten coil mounted on a flexible nitinol wire. One can only surmise that many new brachytherapy applications will emerge in future as a vast number of radionuclides are either present in nature or can be artificially produced and only a small fraction of these radionuclides have been clinically exploited so far. Choice of a radionuclide for a particular type of brachytherapy is a complex decision. The selected radionuclide must have dosimetric characteristics (i.e., energy and half-life) suitable for the specific clinical need. In addition, other critical factors are easy availability, reasonable cost, high-activation cross sections, low levels of undesirable contaminants, and high specific activity. The search for the ideal radionuclide for specific clinical applications continues to engage research physicists. In summary, the basic physical characteristics of commonly used radionuclides in brachytherapy are presented in Table 1.
30
Ravinder Nath Table 1. Physical Characteristics of Various Radionuclides Used for Brachytherapy
Radionuclide
Half-life
Principal or mean energies from encapsulated sources, MeV* Photon
Radium-226 Cesium-137 Iridium-192 Gold-190 Iodine-125 Palladium-103 Strontium-90 Ytterbium-90 Phosphorus-32
1622 y 30 y 74 d 2.7 d 60 d 17 d 29 y 64 h 14 d
Californium-252
2.65 y
Beta
Neutron
0.830 0.662 0.380 0.412 0.028 0.021 0.50 2.27 1.71 2.15
*These are nominal values assuming typical encapsulation for sources.
Acknowledgments Some of the information presented here has been taken from an article that was coauthored by the present author. More details and many other items are presented in the original article entitled “Basic Physics of Brachytherapy” by Ali Meigooni, Cheng B. Saw, and Ravinder Nath in the textbook entitled Principles and Practice of Brachytherapy edited by Subir Nag, Futura Publishing Company, Inc., Armonk, NY, 1997.
Further Reading Meigooni, A., C. B. Saw, and R. Nath. “Basic Physics of Brachytherapy” in Principles and Practice of Brachytherapy. Subir Nag (ed.). Armonk, NY: Futura Publishing Company, Inc., 1997. Nath, R. “Physical Characteristics and Clinical Uses of Brachytherapy Radionuclides” in Brachytherapy Physics. J. F. Williamson, B. R. Thomadsen, and R. Nath. Madison, WI: Medical Physics Publishing, pp. 7–37, 1995.
Chapter 4
LDR Sources: Design and Delivery Systems Robert E. Wallace, Ph.D. UCLA Clinical Professor at Radiation Oncology Cedars-Sinai Outpatient Cancer Center at The Samuel Oschin Comprehensive Cancer Institute Los Angeles, California Low Dose Rate Brachytherapy: Definition and Rationale . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 31 History of LDR Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 32 Overview of Manchester, Quimby, Paris Systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 32 Patterson-Parker . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 32 Fletcher and Other Radium and Cesium Methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 33 Requirements on Source Design Deriving from Implanting Systems and Methods . . . . . . . . . . . 33 Encapsulation, Embodiments, and Desirable Features of Sources . . . . . . . . . . . . . . . . . . . . . . . 33 Quick History of Dosimetry and Strength Standards . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 34 Source Archetypes, 125I: Amersham and 103Pd Theragenics Seeds . . . . . . . . . . . . . . . . . . . . . . . . . . . 35 History of Source Designs for 125I and 103Pd . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 36 Conclusions and Remarks . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 40 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 42
Low Dose Rate Brachytherapy: Definition and Rationale The use of isotopes in treating tumors was first suggested soon after the discovery and separation of radium (Bell 1903). Bell observed one of the tenets of brachytherapy in that, by inserting (sealed) radioactive material directly into tumors, tumor dosage can be improved while the effects on healthy tissues can be reduced when compared to external x-rays (i.e., Crookes tube roentgen rays). That is, dose is not deposited in healthy tissues due to the traversal of x-rays from an external source to an internal target. Dose to healthy surrounding tissues in brachytherapy is a function of the placement of the radioactive sources and the type and energy of the radiation, the ability of the radiation to penetrate distances greater than the dimensions of the tumor, and the physical design of the source. There are radiobiological advantages to the dose distribution, the (usually short) time of treatment, and the use of low dose rate (Brenner 1997). Brenner observes that an array of implanted sources offers an “optimal conformal” dose distribution relative to that achievable using external beam systems. An appropriately designed implant can deliver high doses to the tumor while “minimizing radiobiological damage to normal adjacent tissues” (Brenner 1997). Brenner further observes that “good dose distributions spare: (1) early responding normal tissues, which, in external beam radiotherapy, typically produce the complications that force treatments to be prolonged over more than one month; and (2) late-responding normal tissues, which, in external beam radiotherapy, often represent the dose-limiting endpoint.” Shorter treatment time in brachytherapy generally overcomes accelerated tumor repopulation (Withers, Taylor, and Maciejewski 1988; Brenner 1993). Since in an implant, the dose to adjacent normal tissue is typically reduced, accelerated treatment has less effect on normal tissue. Low dose rate (LDR) irradiation generally reduces late effects in normal tissue more than it reduces tumor control (Coutard 1932; Lea and Catcheside 1942). From these observations, general aspects of brachytherapy source design are indicated: materials sealed in capsules, energy and radiation type selected for local penetration, activity and half-life selected for low dose rates, encapsulation, and internal isotope distribution designed to support conformal tumor dose distribution when an ensemble of sources are used.
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History of LDR Brachytherapy In the early twentieth century, rigid and large brachytherapy sources contained radium and were sealed to contain radon gas and other toxic decay products and to filter α and β ray emission (Dominici 1907; Regaud 1924, 1930). By midcentury, man-made isotopes, tantalum-182 (182Ta) and iridium-192 (192Ir), in the form of wires made implantation more feasible (Van Miert and Fowler 1956; Hall, Oliver, and Shepstone 1966; Pierquin et al., 1978). Other such isotopes including cesium-137 (137Cs), gold-198 (198Au), cobalt-60 (60Co), iodine-125 (125I), ytterbium-169 (169Yb), phosphorus-32 (32P), and palladium-103 (103Pd) have been used in discrete, encapsulated sources of different sizes and applications (Hendee and Ibbott 1996). Fission spectrum neutron sources encapsulating californium-252 (252Cf) have been used to exploit the differences in tissue/tumor repair capacity to damage from high LET (linear energy transfer) radiations (Oliver and Wright 1969; Krishnaswamy 1971; Anderson 1974; Smith, Almond, and Delclos 1974; Rivard 2000). The development of LDR sources in the past 10 years has largely focused on sources containing 125I and 103Pd owing to the increase in their use in treating prostate cancer (Russell and Blasko 1993). Development in the field of brachytherapy sources includes designs encapsulating cesium-131 (131Cs) (Murphy et al., 2004), first suggested earlier (Henschke and Lawrence 1965). The use of sources in treating cancer quickly led to the development of systems, tables, and nomograms to facilitate and expedite clinical applications. The assumptions and the dosimetry of these systems indicate some of the desired features and some requirements for source design. Overview of Manchester, Quimby, Paris Systems Prior to the advent of robust computer treatment-planning systems for brachytherapy, several different systems of distribution (i.e., placement) rules, dose tables, and dose specification criteria were developed to optimize particular implant geometries. Each system had assumptions about what determined a “good dose distribution” and each was constrained by the available sources. The Manchester system (Paterson and Parker 1934, 1938, 1995; Meredith 1967) was designed to maximize dose homogeneity within a treated plane or volume using planar or volume arrays of radium sources. The loading rules dictate a variable loading in the implant, some regions having more or stronger sources than others. As such, implementation of the system requires a large variety of radium needles of different active lengths and activities. Using wires or seeds having uniform activity, one can achieve the differential loading of the Manchester system by altering the geometric loading rules and source-spacing requirements of the system, originally fixed to 1-cm spacing. In one approach, the differential loading is accomplished by using sources of differing lengths and strengths in a uniform geometry. In another, identical sources are placed in a variable geometry to achieve differential loading. The Quimby system (Quimby 1944; Quimby and Castro 1953) for implants is distinguished by its use of uniform spatial distribution of uniform source activity within a plane or a volume. Spacing is uniform but varies in dimension between 1 to 2 cm between sources. The Paris system (Pierquin, et al. 1978; Dutreix and Marinello 1987) was designed for uniformly spaced, uniform linear density 192I wire sources in one or two planes for planar and volume implants. Spacing between wires is 1/2 to 2 cm. In two-plane implants covering a volume, wires are distributed in equilateral triangle or square geometry. The system applies also to strands of 192Ir seeds, spaced uniformly (Pierquin and Marinello 1997). Paterson-Parker In the development of the Manchester system for radium implants, Parker considered the non-uniform distribution of activity along linear sources that are placed on a uniform grid (Meredith 1967). The calculations of this model indicate a ratio of activity of 3:1 between the rind (periphery) and the core (contained
4–LDR Sources: Design and Delivery Systems
33
volume). That is, approximately 75% of the activity is placed on the periphery, the balance uniformly distributed in the enclosed volume. Modified Loading for Modern Prostate Implants. In planning a permanent implant of the prostate using low-energy sources of 125I or 103Pd, a modified Paterson-Parker-like loading can limit the urethra dose to just above the prescribed prostate dose. In studies of possible loading schemes for prostate therapy, a modified uniform loading placing sources toward the periphery while removing sources proximal to the urethra can improve dose coverage while reducing urethral dose (Butler et al., 2000). The recommendations of the AAPM Task Group 56 (TG-56) are that “the treatment plan is designed to place seeds peripherally to improve dose homogeneity and to avoid unnecessary radiation damage to the urethra” (Nath et al., 1997). Whether the ratio of peripheral to core activity is 2:1 or 3:1, modern prostate implants reflect the earlier observations of Parker (Meredith 1967) regarding the achievement of uniform volume dose by using non-uniform distribution of activity, or a non-uniform distribution of a number of sources of equal activity. Fletcher and Other Radium and Cesium Methods Intracavitary applicators for the treatment of cervical cancers at Manchester were first described for use with radium in 1938 (Tod and Meredith 1953; Meredith 1967a). The applicators were similar to later devices in geometry, having an intrauterine linear array of sources with paired vaginal sources possibly with some shielding for the bladder and rectum (Fletcher 1953; Haas, Dean, and Mansfield 1985). Sealed sources of radium and later 137Cs are used with no specific requirement on source design other than consistency with the apertures in the rigid applicator bodies. Modern cesium sources from any manufacturer have utilized one of the few designs that existed when the applicators themselves were designed. This reduces cost because patents no longer protect those designs and using an established design speeds regulatory approval. Requirements on Source Design Deriving from Implanting Systems and Methods The requirements on modern sources are fairly obvious: sealed and durable encapsulation; availability (in quantity) in different strengths; existence of a standard for source strength; the existence of standardized single-source dosimetric parameters in a form usable in treatment planning; visibility using radiographic, ultrasound, or magnetic resonance imaging; and compatibility with existing source designs and source delivery systems (e.g., applicators, needles, catheters). The use of discrete sources in ultrasound-guided prostate implants leads to some desirable features in x-ray source dosimetry: high dose per contained activity (may reduce cost), relatively isotropic distribution around the source (since source orientation is possibly indeterminate), and limited penetration (i.e., low energy) to spare adjacent tissues. Sources in permanent implants for prostate brachytherapy should be relatively small, given that the prostate is not a large target.
Encapsulation, Embodiments, and Desirable Features of Sources Bell had suggested encapsulating radium salts in glass tubes (Bell 1903), while the benefits of filtering and sealing radium sources in a platinum-iridium needle were observed by Dominici (1907). Radium needles were manufactured in different active lengths with a variety of internal distributions of activity: uniform, dumbbell (higher activity at needle ends), and club (higher activity at one end). Cesium-137 is a substitute for radium and is available in needles and in tubes. The advent of afterloading allowed the replacement of needle-based brachytherapy with tube sources that would be placed within implanted catheters, hollow needles, or intracavitary applicators. Typical dimensions for stainless steel, encapsulated, ceramic 137Cs sources in current production are approximately 3 mm in diameter by 10 to 20 mm physical length with the tube ends plugged and welded (Devine 2004). Recent evaluation
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Robert E. Wallace
of dosimetry for a variety of 137Cs source designs allows for treatment planning using the current AAPM TG-43 formalism (Liu, Prasad, and Bassano 2004; Rivard et al., 2004) Outdated in use, 182Ta sources had been available for temporary implants in the form of flexible wires, 0.2 mm diameter, encased in a 0.1 mm thick platinum sheath (Cohen 1955). Wires of 192Ir are constructed with a core of 25% iridium, 75% platinum, 0.1 to 0.3 mm in diameter, encased in a 0.1 mm thick platinum sheath (Hendee and Ibbott 1996). Iridium seed sources are available in two designs, the first has a 0.3 mm diameter core 10% 192Ir/90% platinum clad in a 0.1 mm thick platinum sheath (Best Industries, Springfield, VA). The second design has a 0.1 mm diameter core of 30% iridium/70% platinum clad in a 0.2 mm thick stainless steel sheath (Alpha-Omega Services, Bellflower, CA). Iridium seeds are typically delivered in nylon strands at uniform 1 cm center-to-center spacing. As a substitute for gold-encapsulated pellets containing radon gas, “seed” type small sources containing 198Au are used in some applications (Slanina et al. 1982; Hendee and Ibbott 1996). Present manufacture provides these sources with the active gold in a 2.2 mm by 0.5 mm diameter pellet encapsulated in a 0.15mm thick platinum sheath, resulting in a seed that is 2.5 mm long by 0.8 mm in diameter (Best Industries, Springfield, VA). Seed sources containing 125I or 103Pd are available in a number of designs, differing principally in internal configuration. At one point the market in the United States had nearly 20 different manufacturers of 125 I seeds and 5 manufacturers of 103Pd seeds. The large number of manufacturers and differences in source design derive from the increased demand due to increased popularity in the late 1990s of their use in permanent-implant prostate brachytherapy, and the existence of active, enforceable patents on existing source designs. The market competition has led to novel designs that variously improve single-source dosimetry, improve visibility, improve availability, and reduce manufacturing costs.
Quick History of Dosimetry and Strength Standards The dosimetry of brachytherapy sources has developed through a series of reports over the past 18 years. The reports covered source strength standards, dosimetry, and use of brachytherapy sources. Reports of exposure standards were published by the National Bureau of Standards [NBS, which is now the National Institute of Science and Technology (NIST)] for 137Cs (Loftus 1970), 192Ir (Loftus 1980), and 125I (Loftus 1984). Kubo observed that 125I exposure measured in the NBS 1984 standard using the Ritz chamber (Ritz 1960) included the effects of fluorescent x-rays produced in the titanium encapsulation of the 125I sources. These titanium soft (approx. 5 keV) x-rays have little practical significance in phantom dose measurements (Kubo 1985; Williamson 1988). As long as a consistent standard and individual source dose-rate constant data were used, under the 1984 standard, patient administered dose was not significantly affected. In 1987, the AAPM TG-32 published, in AAPM Report No. 21 (AAPM 1987), the recommendation that brachytherapy source strength should be specified in terms of air kerma strength and that such should be adopted by NBS/NIST and by the medical physics community for clinical use. NBS responded with a standard (Loftus 1988), a subsequent paper provided direction for clinical implementation (Williamson and Nath 1991), and the use of air kerma strength was endorsed by the American Endocurietherapy Society (Williamson et al., 1993). Also, in approximately 1988, a manufacturer implemented an ad hoc 103Pd source activity standard that was based on intercomparison with a 109Cd source having NIST-traceable activity calibration (Williamson et al., 2000). In 1999, NIST reported an air kerma strength standard (Seltzer et al., 1999; Seltzer and Lamperti 1999; Seltzer et al., 2003) for both 125I and 103 Pd using a new device (Loevinger 1993) that addressed the effect observed by Kubo. The 1999 NIST standard for 125I and 103Pd air kerma strength had direct bearing on revision of dosimetric parameters. Because air kerma strength under the NIST 1999 standard for 103Pd was significantly different from that derived under the existing ad hoc standard, patient dose prescription when using 103Pd was also affected. Suggestions for revision of the prescribed dose were enunciated by the AAPM (Williamson et al., 2000).
4–LDR Sources: Design and Delivery Systems
35
The 1995 AAPM TG-43 report unified dosimetry for 192Ir, 125I, and 103Pd brachytherapy sources (Nath et al., 1995). Subsequent papers described clinical implementation of the recommendations (Luse, Blasko, and Grimm 1997; Bice et al., 1998). In anticipation of the 1999 NIST standard for 125I, the report of the AAPM ad hoc committee on 125I sealed-source dosimetry addressed the implementation of TG-43 dosimetry in the clinic, dose prescription changes for 125I using the published TG-43 parameters, and clinical incorporation of the proposed air kerma strength standard from NIST (Kubo et al., 1998). That report addressed the 125I sources available in September 1997, models 6702 and 6711 (Nycomed-Amersham, Arlington Heights, IL), and provided a formulary for implementing the 1999 NIST standard. The arrival of a new source design in the market in early 1998 (Wallace and Fan 1998) and the prospect of further designs (Wallace and Fan 1999) prompted the delineation (Williamson et al., 1998) of requisite procedures to characterize any new source not considered in TG-43. These recommendations by the AAPM ad hoc subcommittee on low-energy brachytherapy source dosimetry were followed by specific instructions to users (Williamson et al., 1999) regarding changes in the NIST standard for 125I sources implemented early in 1999. The arrival in January 1999 of new design 103Pd sources (Wallace and Fan 1999, 1999a; Li, Palta, and Fan 2000; Nath et al., 2000; Williamson 2000) and the contemporaneous enunciation of the 1999 NIST standard for 103Pd air kerma strength raised two issues. The first issue concerned the appropriate therapeutic dose prescription for 103Pd under the new 1999 NIST standard and how to convert a decade of clinical experience using a source having an activity intercomparison standard. This was addressed in an exhaustive report of the ad hoc subcommittee (Williamson et al. 2000). This report also provides instructions for clinical implementation of the new 103Pd source strength standard and new dosimetric evaluations. The second issue was the continued use of apparent activity both by manufacturers and in clinical application. In their communication, the ad hoc subcommittee reaffirmed earlier recommendations and provided direction in the interim use of apparent activity during the transition to TG-43 based dosimetry (Williamson et al., 1999a). NIST and the manufacturing community recognized anomalous results in air kerma strength standardization measurements made under the primary standard for late 1998 through 1999. Discrepancies of 2% to 7% were identified, but users of the affected 125I and 103Pd source designs were advised to make no changes to any dosimetry parameters until a detailed formal analysis was completed (Williamson et al., 2000a). An update of the TG-43 report was subsequently published that reviewed dosimetric formalism, methods for measurement and analysis of data for new sources, and that compiled consensus data for selected sources (Rivard et al. 2004, 2004a). All measured dose-rate constant data in the updated report reflect corrections for the 1999 anomaly and should be applied to the affected sources in the clinic.
Source Archetypes, 125I: Amersham and 103Pd Theragenics Seeds In the United States, low-energy 125I sources were first manufactured four decades ago by the Lawrence Soft X-ray Company. The 3M Company took over the manufacturing and then sold that part of their business to what are now the Amersham Health/MediPhysics (Arlington Heights, IL) and Oncura (Plymouth Meeting, PA) companies. These sources, the models 6701, 6702, 6711, and the more recent 6733 demonstrate most of the design issues in 125I sources. In 1988–1989, Theragenics Corporation (Norcross, GA) released one design of a 103Pd source, the model 200, that has undergone small, incremental manufacturing variations to the present day. A 103Pd source, model MED 3933, was released in 1999 by North American Scientific, Inc. (NAS, Chatsworth, CA) and a new source from Mills Biopharmaceuticals, Inc. (MBI, Oklahoma City, OK) recently was reported (Rivard, Melhus, and Kirk 2004). These three designs provide archetypes for 103Pd sources. A unique 103Pd source design, RadioCoil™, (RadioMed Corporation, Tyngsboro, MA) has been reported where the source is a coiled wire of activated rhodium (Meigooni et al., 2004).
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History of Source Designs for 125I and 103Pd The model 67xx source designs are depicted in Figure 1. The first source introduced in this series, the 6701, was a titanium cylinder, 4.5 mm nominal length, and 0.8 mm outer diameter, end plugged and welded, depicted in Figure 1a. This design contained two active resin spheres (0.5 mm diameter) that flank a central radiopaque marker. 125I is incorporated into the resin via an ion-exchange mechanism when the resin spheres are soaked in a concentrated solution of an ionic salt containing 125I. This is one of the more common mechanisms of incorporating radioactive iodine in sources. The use of only three spheres appears to have been dictated by the process of welding plugs at the source container ends where heat dissipation may have been an issue. Modern laser welding techniques are designed to minimize this problem and typically remain as industrial secrets. The 6701 is visible in a radiograph while the 6702 design is not. In the 6702, the central marker is replaced with a third active sphere, apparently to increase the activity range of the design. Both designs are presently unavailable for routine use. In Figure 1b, the model 6711 design is shown. In this design, the 125I is adsorbed (i.e., plated) onto a 3 mm long, 0.5 mm diameter silver rod using proprietary techniques. This design provides improved radiographic visibility with the possibility of determining both location and orientation. The incorporation of the silver substrate adds characteristic x-rays from silver to the 125I spectrum. This had a significant effect on the calibration of 125I source strengths at NIST, as outlined above. The use of ultrasound localization in prostate brachytherapy generated a desire to improve acoustic visibility of sources. The model 6733 source (Figure 1c) represents an approach to improved visibility of individual sources under ultrasound. The principal difference between this design and the 6711 is the encapsulation. The corrugation of the cylindrical titanium tube portion of the capsule provides a wide-angle dispersion of reflected sound back to the ultrasound transducer. Since reflection is specular, a flat surface will reflect sound at an angle that mirrors the incident angle, usually away from the transducer. Corrugation provides several different reflection angles, thereby improving the likelihood that some reflected sound will be detected at the transducer. Also, in the 6733, the end-welds are less massive than in the 670x and 6711 designs. The design of the model 200 103Pd source is shown in Figure 2. In the model 200 design, the active elements have a graphite substrate and the marker is lead. The inverted end-cups sealing the ends of the titanium cylinder allow welding to be distant from the internal components while presenting lower attenuation than a 6701 type end plug. The concavity and extent of the end-cups serves to limit the motion of internal elements of the source. There are many other designs for sources that are available for use. The AAPM updated TG-43 report (Rivard et al., 2004) provides comprehensive analysis of the dosimetry of sources, while a review of how source design affects dosimetry of 125I sources is also available (Heintz, Wallace, and Hevezi 2001). In terms of dosimetry, the model 670x sources differ from model 6711/33 sources. In these latter source designs, the inclusion of characteristic x-ray from the interaction of the 125I x-rays in the silver substrate lowers the average energy of the emitted spectrum, decreases the penetration in water, and decreases the reference dose per unit source strength (i.e., the dose-rate constant). Compared to the 670x sources, the dose-rate constants for the 6711/6733 sources are approximately 6% to 7% lower than those of the 670x sources (Heintz, Wallace, and Hevezi 2001; Rivard et al. 2004, 2004a). The Mills 125I and 103Pd sources share the same design, Figure 3. These sources have a standard size, 4.5 mm length by 0.8 mm diameter, titanium capsule with laser-welded ends. The end welds of these sources are somewhat less massive than those found in the 670x or 6711 sources. The isotope bearing elements are five silver spheres, plated with either 125I or 103Pd. These active and radiopaque elements move within the source, with random orientations shown in Figure 3a that compare to an idealized configuration, Figure 3b, used for dosimetry modeling studies (Li 2002). The net effect of internal motion is to introduce a small (+0.3/–0.2) uncertainty in the dose-rate constant for the 103Pd source that is about 3% lower than for other available sources (Rivard, Melhus, and Kirk 2004). This is not clinically untenable considering that the spread in the dose-rate constant among the various 125I source designs is on the order
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Figure 1a. 6701, 6702 sources. Model 6702 is shown. In the 6701, the central resin sphere is replaced with a radiopaque marker. [Reprinted from Heintz, B. H., R. E. Wallace, and J. M. Hevezi, “Comparison of I-125 sources used for permanent interstitial implants,” Med Phys 28:671–682. © 2001, with permission from AAPM.]
Figure 1b. Model 6711. [Reprinted from Heintz, B. H., R. E. Wallace, and J. M. Hevezi, “Comparison of I-125 sources used for permanent interstitial implants,” Med Phys 28:671–682. © 2001, with permission from AAPM.]
Figure 1c. Model 6733. [Reprinted from Meigooni, A. S., S. A. Dini, K. Sowards, J. L. Hayes, and A. Al-Otoom, “Experimental determination of the TG-43 dosimetric characteristics of EchoSeed™ model 6733 125I brachytherapy source,” Med Phys 29:939–942. © 2002, with permission from AAPM.]
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Figure 2. Model 200. [Reprinted from Rivard, M. J., B. M. Coursey, L. A. DeWerd, W. F. Hanson, M. S. Huq, G. S. Ibbott, M. G. Mitch, R. Nath, and J. F. Williamson, “Update of AAPM Task Group No. 43 Report: A revised AAPM protocol for brachytherapy dose calculations,” Med Phys 31:633–674. © 2004, with permission from AAPM.]
Figure 3a. MBI 125I and 103Pd sources, radiographs. [Reprinted from Rivard, M. J., C. S. Melhus, and B. L. Kirk, “Brachytherapy dosimetry parameters calculated for a new 103Pd source,” Med Phys 31:2466–2470. © 2004, with permission from AAPM.]
Figure 3b. MBI 125I and 103Pd sources containing five 0.5-mm diameter silver spheres plated with 125I or 103Pd. [Reprinted from Li, Z., “Monte Carlo calculations dosimetry parameters of the Urocor Prostaseed 125I source,” Med Phys 29:1029–1034. © 2002, with permission from AAPM.]
of 5% (Heintz, Wallace, and Hevezi 2001; Rivard et al., 2004, 2004a). These designs are similar to the 6711 and 6733 in that the isotope is distributed on the surface of a silver substrate. The model IS-125-01 125 I source (Imagyn Medical Technologies, Denton, TX) is similar in design to the Mills sources where the five silver spheres are 0.64 mm in diameter. The model 1A1-125A (IsoAid, Port Richey, FL) resembles the 6711 design, yet with thinner end-welds. The model 130.002 125I (Nucletron Corporation, Columbia, MD) also resembles the 6711 source with a longer marker, 3.4 mm versus 3 mm (Karaiskos et al., 2001). Consequently, the Imagyn, IsoAid, Mills, and Nucletron 125I sources have an emitted spectrum that includes characteristic x-rays from silver. The palladium substrate of the model BT-125I 125I source (no longer available; Syncor Corporation, Golden, CO) provides characteristic x-rays comparable to those from silver. The use of silver or palladium substrate results in a lowered dose-rate constant that is similar to that for the models 6711/6733 (Heintz, Wallace, and Hevezi 2001; Karaiskos et al., 2001;
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Rivard et al. 2004, 2004a). The use of a tungsten substrate (and marker rod) in the model 2301 (Best Medical Industries, Springfield, VA) adds characteristic x-rays that do not significantly affect the otherwise pure 125I emitted spectrum (Heintz, Wallace, and Hevezi 2001). The double-walled container of the model 2301 presents very thin end walls (about 0.04 mm) and 0.08 mm thick body walls. (Most other sources have body wall thickness on the 0.04 to 0.05 mm.) The model 3500 125I source (Implant Sciences Corporation, Wakefield, MA) contains a silver marker of smaller diameter than used in the 6711 but with pointed ends to enhance radiographic visibility. Due to the reduced mass of the silver marker as compared to the model 6711, the dose-rate constant for this source is between that of the 6702 and 6711 sources (Rivard 2002). In the model 3500, 125I is incorporated into a quartz tube via the ion implantation of 124Xe, subsequently neutron-activated to 125Xe (17 hour half-life) and 125I. After decay of the 125Xe, the active quartz tubes are processed into brachytherapy sources (Munro 2004). The model I125.S06 (BeBig/Theragenics, Norcross, GA), and the model STM125I (Source Tech Medical, Carol Stream, IL) each contain linear gold markers in novel configurations that place the 125I in either a ceramic matrix, in the I125.S06, or plated on a copper sheath covering an aluminum tube that itself contains the gold marker rod. As the case for the tungsten marker in the model 2301, the characteristic x-rays from the gold, copper, and aluminum components do not significantly affect the dosimetry of these sources (Heintz, Wallace, and Hevezi 2001). Of the sources discussed thus far, all but the models 6701 125I and 200 103Pd sources are essentially linear sources (or an uninterrupted linear array of active elements—spheres). In radionuclide distribution, these sources resemble the uniform radium needles. The model 200 103Pd source, the NAS model MED-3631-A/M 125I , MED3633 103Pd sources, the Best Industries model 2335 103Pd source, and the model LS-1 125I source (Draximage, Inc., Kirkland, Quebec, Canada) each place the active elements on organic materials separated by a marker. This echoes the 6701 design. These designs also reflect the “dumbbell” designs of radium needles, and for the same reasons. The dose distributions around these sources should be more isotropic than a line source (Heintz, Wallace, and Hevezi 2001; Rivard et al., 2004, 2004a), particularly when the end-walls/-welds are less massive than those of the 670x designs. The NAS sources are depicted in Figure 4, and contain four active elements (resin beads suffused with the isotope) separated by two gold/copper alloy markers. In the Best model 2335 source, six polymer beads containing the 103Pd are separated by a short tungsten marker within a double-wall container (Meigooni et al., 2001). The Draximage model LS-1 source is likely the most complicated design that is currently available, shown in Figure 5 (Heintz, Wallace and Hevezi 2001). In the source, two glass beads carrying the 125I are separated by a Pt/Ir(10%) marker with centralized extra wall thickness using internal titanium spacers that also fix the locations of the glass beads. The weld of this source is also centralized to allow minimal end-wall thickness. In this source, the activity contained
Figure 4. NAS Model 3631-A/M 125I and Model 3633 103Pd sources. [Reprinted from Heintz, B. H., R. E. Wallace, and J. M. Hevezi, “Comparison of I-125 sources used for permanent interstitial implants,” Med Phys 28:671–682. © 2001, with permission from AAPM.]
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Figure 5. Draximage model LS-1 125I source. [Reprinted from Heintz, B. H., R. E. Wallace, and J. M. Hevezi, “Comparison of I-125 sources used for permanent interstitial implants,” Med Phys 28:671–682. © 2001, with permission from AAPM.]
in/on the glass beads should be identical in order to obtain an ideally isotropic dose distribution. The same is true for the model 200 source, Figure 2. The use of multiple active elements in the model MED3631A/M, MED 3633, and 2335 sources allows some variation in the activity carried by each bead so long as the net activity at each source end is nearly identical. The manufacturing issue reduces to efficient sorting in these sources rather than exact activation in the models 200 and LS-1. The model 1251L (International Brachytherapy, Norcross, GA) 125I source, Figure 6, is a concentric tubular design that allows a suture to be run through the source to form strands of sources. Dumbbellstyle end-loading and relatively thin encapsulation provide for a relatively high dose-rate constant and better isotropy than a 6711 type source (Heintz, Wallace and Hevezi 2001). A novel design 103Pd source, the RadioCoil, is outlined in Figure 7. This source, available in various lengths, combines the flexibility of the Paris system wires with the low-energy of 103Pd. The source is fabricated of a coiled flat-wire of high-purity rhodium that is neutron-activated to 103Pd. The 5 mm length source, a size comparable to other 103Pd seeds, also demonstrates equivalent dosimetry: dose-rate constant, penetration in tissue, and dose isotropy (Meigooni et al., 2004). A new 131Cs source, model CS-1 (IsoRay, Inc., Richland, WA), has been reported that resembles the model I125.S06, 3500, and STM-125I sources. In Figure 8, the source is shown with an axial gold marker rod, a glass/ceramic tube bearing the 131Cs, with ring welded end-caps. Relative dosimetry appears to be comparable to 125I sources in penetration (due to similar average energy, 29 to 34 keV) and dose anisotropy. The initial dose rate is higher than 103Pd due to the 9.7-day half-life and the dose-rate constant is nearer that for 125I sources than 103Pd sources (Murphy et al., 2004). As of this writing, no formal air kerma strength standard exists for 131Cs brachytherapy sources such as are available for either 103Pd or 125I sources.
Conclusions and Remarks There has been considerable effort in improving the design and manufacture of brachytherapy sources in the past decade, focusing primarily on seeds containing 103Pd and 125I. Legacy designs for 198Au, 192Ir, and 137Cs remain available for use. A new seed for 131Cs has been reported for clinical use, possibly in place of 125I, but an air kerma strength standard needs to be enunciated. The principles and desirable features of source design were laid out in the early twentieth century, notably by Paterson and Parker in 1934, and endure in current designs. Not specifically mentioned in this chapter are significant advances in packaging: seed cartridges, preloaded needles and catheters, uniform and customized non-uniform spaced strands of sources, or automated needle loading systems.
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Figure 6. IBT model 1251l. [Reprinted from Heintz, B. H., R. E. Wallace, and J. M. Hevezi, “Comparison of I-125 sources used for permanent interstitial implants,” Med Phys 28:671–682. © 2001, with permission from AAPM.]
Figure 7. RadioCoil™ 103Pd activated rhodium source 6733. [Reprinted from Meigooni, A. S., H. Zhang, J. R. Clark, V. Rachabatthula, and R. A. Koona, “Dosimetric characteristics of the new RadioCoil™ 103Pd wire line source for use in permanent brachytherapy implants,” Med Phys 31:3095–3105. © 2004, with permission from AAPM.]
Figure 8. IsoRay model CS-1 131Cs brachytherapy source. [Reprinted from Murphy, M. K., R. K. Piper, L. R. Greenwood, M. G. Mitch, P. J. Lamperti, S. M. Seltzer, M. J. Bales, and M. H. Philips, “Evaluation of the new cesium-131 seed for use in low-energy x-ray brachytherapy,” Med Phys 6:1529–1538. © 2004, with permission from AAPM.]
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Loftus, T. P. (1970). “Standardization of 137Cs gamma-ray sources in terms of exposure units (roentgens).” J Res Nat Bur Stand 74A:1–6. Loftus, T. P. (1980). “Standardization of 192Ir gamma-ray sources in terms of exposure.” J Res Nat Bur Stand 85:19–25. Loftus, T. P. (1984). “Exposure standardization of 125I seeds used for brachytherapy.” J Res Nat Bur Stand 89:295–303. Loftus, T. P. (1988). “Calibration of gamma-ray emitting brachytherapy sources.” Nat Bur Stand Spec Pub 89:2250. Luse, R. W., J. Blasko, and P. Grimm. (1997). “A method for implementing the American Association of Physicists in Medicine Task Group-43 dosimetry recommendations for 125I transperineal prostate seed implants.” Int J Radiat Oncol Biol Phys 37:737–741. Meigooni, A. S., Z. Bharucha, M. Yoe-Sein, and K. Sowards. (2001). “Dosimetric characteristics of the Best® doublewall 103Pd brachytherapy source.” Med Phys 28:2568–2575. Meigooni, A. S., S. A. Dini, K. Sowards, J. L. Hayes, and A. Al-Otoom. (2002). “Experimental determination of the TG-43 dosimetric characteristics of EchoSeed™ model 6733 125I brachytherapy source.” Med Phys 29:939–942. Meigooni, A. S., H. Zhang, J. R. Clark, V. Rachabatthula, and R. A. Koona. (2004). “Dosimetric characteristics of the new RadioCoil™ 103Pd wire line source for use in permanent brachytherapy implants.” Med Phys 31:3095–3105. Meredith, W. J. “Physical Aspects of the Interstitial Treatment System.” Chapter 10 in Radium Dosage: The Manchester System, ed. 2. Edinburgh: E. & S. Livingstone, pp. 83–107, 1967. Meredith, W. J. “Dosage for Cancer of the Cervix Uteri.” Chapter 6 in Radium Dosage: The Manchester System, ed. 2. Edinburgh: E. & S. Livingstone, pp. 42–50, 1967a. Munro, J. J. (2004). Private communication, Implant Sciences Corporation, Wakefield, MA. Murphy, M. K., R. K. Piper, L. R. Greenwood, M. G. Mitch, P. J. Lamperti, S. M. Seltzer, M. J. Bales, and M. H. Philips. (2004). “Evaluation of the new cesium-131 seed for use in low-energy x-ray brachytherapy.” Med Phys 6:1529–1538. Nath, R., L. L. Anderson, G. Luxton, K. A. Weaver, J. F. Williamson, and A. S. Meigooni, (1995). “Dosimetry of interstitial brachytherapy sources: Recommendations of the AAPM Radiation Therapy Committee Task Group No. 43.” Med Phys 22:209–234. Also available as AAPM Report No. 51. Nath, R., L. L. Anderson, J. A. Meli, A. J. Olch, J. A. Stitt, and J. F. Williamson. (1997). “Code of practice for brachytherapy physics: Report of the AAPM Radiation Therapy Committee Task Group No. 56. American Association of Physicists in Medicine.” Med Phys 24:1557–1598. Also available as AAPM Report No. 59. Nath, R., N. Yue, K. Shahnazi, and P. J. Bongiorni. (2000). “Measurement of dose-rate constant for 103Pd seeds with air-kerma strength calibration based upon a primary national standard.” Med Phys 27:655–658. Oliver, G., and C. Wright. (1969). “Dosimetry of an implantable 252Cf source.” Radiology 92:143. Paterson, R., and H. M. Parker. (1934). “A dosage system for gamma-ray therapy.” Br J Radiol 7:592–632. Paterson, R., and H. M. Parker. (1938). “A dosage system for interstitial radium therapy.” Br J Radiol 11:252–266. Paterson, R., and H. M. Parker. (1995). “A dosage system for gamma ray therapy. 1934.” Br J Radiol 68:H60–100. Pierquin, B., and G. Marinello. A Practical Manual of Brachytherapy. Translated by F. Wilson, B. Erickson, and J. Cunningham. Madison, WI: Medical Physics Publishing, 1997. Pierquin, B., A. Dutreix, C. H. Paine, D. Chassagne, G. Marinello, and D. Ash. (1978). “The Paris system in interstitial radiation therapy.” Acta Radiol Oncol 17:33–48. Quimby, E. H. (1944). “Dosage table for linear radium sources.” Radiology 43:572. Quimby, E. H., and V. Castro. (1953). “The calculation of dose in interstitial radium therapy.” Am J Roentgenol 70:739–749. Regaud, C. (1924). “Some biological aspects of radiation therapy of cancer.” Am J Roentgenol 10:97. Regaud, C. (1930). “Sur les principes radiophysiologiques de la radiothérapie des cancers.” Acta Radiol 11:455. Ritz, V. H. (1960). “Standard free-air chamber for the measurement of low-energy x-rays (20 to 100k – constant potential).” J Res Nat Bur Stand 64:49. Rivard, M. J. (2000). “Neutron dosimetry for a general 252Cf brachytherapy source.” Med Phys 27:2803–2815. Rivard, M. J. (2002). “Comprehensive Monte Carlo calculations of AAPM Task Group Report No. 43 dosimetry parameters for the Model 3500 I-Plant 125I brachytherapy source.” Appl Radiat Isotop 57:381–389.
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Rivard, M. J., C. S. Melhus, and B. L. Kirk. (2004). “Brachytherapy dosimetry parameters calculated for a new 103Pd source.” Med Phys 31:2466–2470. Rivard, M. J., B. M. Coursey, L. A. DeWerd, W. F. Hanson, M. S. Huq, G. S. Ibbott, M. G. Mitch, R. Nath, and J. F. Williamson. (2004). “Update of AAPM Task Group No. 43 Report: A revised AAPM protocol for brachytherapy dose calculations.” Med Phys 31:633–674. Also available as AAPM Report No. 84. Rivard, M. J., W. M. Butler, L. A. DeWerd, M. S. Huq, G. S. Ibbott, Z. Li, M. G. Mitch, R. Nath, and J. F. Williamson. (2004a). “Erratum: ‘Update of AAPM Task Group No. 43 Report: A revised AAPM protocol for brachytherapy dose calculations.’” [Med Phys 31:633–674 (2004)].” Med Phys 31:3532–3533. Russell, K. J., and J. C. Blasko. (1993). “Recent advances in interstitial brachytherapy for localized prostate cancer.” Probl Urol 7:260–278. Seltzer, S. M., and P. J. Lamperti. (1999). Status of NIST Primary Standards for 125I and 103Pd Therapy Seeds Based on the Wide-Angle Free-Air Chamber (WAFAC): Proceedings of a CIRMS Workshop, National Institute of Standards and Technology. Seltzer, S. M., P. J. Lamperti, R. Loevinger, C. G. Soares, and J. T. Weaver. (1999). “New NIST air-kerma strength standards for 125I and 103Pd brachytherapy seeds (abstr).” Med Phys 25:170. Seltzer, S. M., P. J. Lamperti, R. Loevinger, M. G. Mitch, J. T. Weaver, and B. M. Coursey. (2003). “New national air-kerma-strength standards for 125I and 103Pd brachytherapy seeds.” J Res Nat Bur Stand 108:337–358. Slanina, J., M. Wannenmacher, K. Kuphal, H. Knufermann, C. Beck, and W. Schilli. (1982). “Interstitial radiotherapy with 198Au seeds in the primary management of carcinoma of the oral tongue: Results in Freiburg/Breisgau from January 1964 to July 1980.” Int J Radiat Oncol Biol Phys 8:1683–1689. Smith, A., P. Almond, and L. Delclos. (1974). “Evaluation of 252Cf neutron emitter for interstitial and intracavitary radiation therapy.” Eur J Cancer 10:369. Tod, M., and W. J. Meredith. (1953). “Treatment of cancer of the cervix uteri, a revised Manchester method.” Br J Radiol 26:252–257. Van Miert, P. J., and J. F. Fowler. (1956). “The use of tantalum-182 in the treatment of early bladder cancer.” Br J Radiol 29:508. Wallace, R. E., and J. J. Fan. (1998). “Evaluation of a new brachytherapy iodine-125 source by AAPM TG43 formalism.” Med Phys 25:2190–2196. Wallace, R. E., and J. J. Fan. (1999). “Report on the dosimetry of a new design 125iodine brachytherapy source.” Med Phys 26:1925–1931. Wallace, R. E., and J. J. Fan. (1999a). “Dosimetric characterization of a new design 103palladium brachytherapy source.” Med Phys 26:2465–2470. Williamson, J. F. (1988). “Monte Carlo evaluation of specific dose constants in water for 125I seeds.” Med Phys 15:686–694. Williamson, J. F. (2000). “Monte Carlo modeling of the transverse-axis dose distribution of the model 200 103Pd interstitial brachytherapy source.” Med Phys 27:643–654. Williamson, J. F., and R. Nath. (1991). “Clinical implementation of AAPM Task Group 32 recommendations on brachytherapy source strength specification.” Med Phys 18:439–448. Williamson, J. F., L. L. Anderson, D. W. Grigsby, A. Martinez, R. Nath, D. Neblett, A. Olch, and K. Weaver. (1993). “American Endocurietherapy Society recommendations for specification of brachytherapy source strength.” Endocuriether/Hypertherm Oncol 9:1–7. Williamson, J. F., B. M. Coursey, L. A. DeWerd, W. F. Hanson, and R. Nath. (1998). “Dosimetric prerequisites for clinical use of new low energy photon interstitial brachytherapy sources.” Med Phys 25:2269–2270. Williamson, J. F., B. M. Coursey, L. A. DeWerd, W. F. Hanson, R. Nath, and G. Ibbott. (1999). “Guidance to users of Nycomed Amersham and North American Scientific, Inc. I-125 interstitial sources: Recommendations of the AAPM Radiation Therapy Committee ad hoc Subcommittee on low-energy seed dosimetry.” Med Phys 26:570–573. Williamson, J. F., B. M. Coursey, L. A. DeWerd, W. F. Hanson, R. Nath, M. J. Rivard, and G. Ibbott. (1999a). “On the use of apparent activity (Aapp) for treatment planning of 125I and 103Pd interstitial brachytherapy sources: Recommendations of the AAPM Radiation Therapy Committee Subcommittee on low-energy brachytherapy source dosimetry.” Med Phys 26:2529–2530.
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Williamson, J. F., B. M. Coursey, L. A. DeWerd, W. F. Hanson, R. Nath, M. J. Rivard, and G. Ibbott. (2000). “Recommendations of the American Association of Physicists in Medicine on 103Pd interstitial source calibration and dosimetry: Implications for dose specification and prescription.” Med Phys 27:634–642. Also available as AAPM Report No. 69. Williamson, J. F., B. M. Coursey, L. A. DeWerd, W. F. Hanson, G. Ibbott, R. Nath, and M. J. Rivard. (2000a). “Important notice for radiation therapy physicists using 125I or 103Pd brachytherapy sources (Letter).” Online: http://rpc.mdanderson.org/rpc/htm/Home_htm/Low-energy%20documents/ImpNotice_EndUsers_v12.pdf. Withers, H. R., J. M. G. Taylor, and B. Maciejewski. (1988). “The hazard of accelerated tumor clonogen repopulation during radiotherapy.” Acta Oncol 27:131–146.
Chapter 5
Quality Management of Low Dose Rate Brachytherapy Sources Gary A. Ezzell, Ph.D. Mayo Clinic Scottsdale Scottsdale, Arizona Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 47 Acceptance Testing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 47 Preparation for Calibration . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 47 Long-lived Sources . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 48 Short-lived Sources . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 48 Periodic Testing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 50 Per-patient Testing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 51 Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 51 Appendix . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 52 Form 1: Radioactive Materials Requisition . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 53 Form 2: Cesium In-Use Inventory . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 55 Form 3: Iridium In-Use Inventory . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 56 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 57
Introduction The purpose of this chapter is to review one particular part of the quality management (QM) program for low dose rate brachytherapy: that associated with the sources themselves. It is the first link in the quality chain, since a problem with the sources will certainly affect the quality of the brachytherapy treatment. The QM program has three main components: acceptance testing of new sources, periodic testing, and per-patient testing. This discussion is organized in that fashion. The requirements and procedures for each component depend somewhat on the characteristics of the sources being used. Useful divisions are between sources used for temporary or permanent implants (sources which emit primarily photons with an average energy >50 kV and other sources, respectively), long- or short-lived isotopes, and small or extended sources (e.g., 125I seeds or 137Cs tubes.) The discussion for each section reflects these major divisions. Major references guiding this discussion are, of course, the AAPM TG-40 report on comprehensive quality assurance in radiation oncology (Kutcher et al. 1994) and the AAPM TG-56 report on the brachytherapy code of practice (Nath et al. 1997).
Acceptance Testing Preparation for Calibration The local physicist has the responsibility of assuring the accuracy of the calibration of the sources used for brachytherapy. TG-56 states “Every institution practicing brachytherapy shall have a system for measuring source strength with secondary traceability for all source types used in its practice.” Secondary traceability means that the source strength is measured either with an instrument that has been itself calibrated for that source type either at the National Institute of Standards and Technology (NIST) or an Accredited Dosimetry Calibration Laboratory (ADCL) or by comparison to a source of the same type that has been calibrated at NIST or an ADCL. In practice, such measurements are best done with a well-type chamber. So the first step in the process is to obtain such a chamber/electrometer system and arrange with
47
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an ADCL to calibrate either that system or an appropriate source. (The details of the calibration process and its associated QM program are covered in another chapter.) Long-lived Sources For long-lived sources such as 137Cs tubes, TG-56 recommends the following steps: 1. Document the physical/chemical form and encapsulation based on the manufacturer’s specification in order to support dosimetry calculations. 2. Document the initial leak test based on the manufacturer’s certification, and, optionally, repeat with local equipment (see below). 3. Verify the uniformity of the activity distribution within each source (most likely with an autoradiograph; no tolerance limit is given). 4. Verify the location of the activity distribution within each source relative to the exterior dimensions to a tolerance of 1 mm (most likely by combining an autoradiograph with a transmission radiograph). 5. Verify the identification of each source (serial numbers, color coding, etc.). 6. Calibrate each source. If the measured source strength agrees with the manufacturer’s specification to within 3%, then either may be used for dose calculations. Differences larger than 3% should be investigated, and differences larger than 5% should be reported to the manufacturer. TG-56 asserts “It is always advisable to ask the manufacturer to review its calibration of the sources to help resolve these discrepancies.” When sources are batched for dosimetry purposes, then the 3% tolerance applies to the mean of the batch, and the range of source strengths within the batch should not exceed 5%. Long-lived sources need to be leak tested at intervals specified by the radioactive materials license, typically 6 months or 3 years, depending on the source type. Such sources should have certified leak tests before being shipped, and that documentation can show initial compliance. Users of sources such as 137Cs tubes may decide to repeat the leak test locally during acceptance; certainly that should be done if there is any suggestion of damage. Properly leak testing sealed sources requires sensitive instrumentation, such as a NaI well counter, a means of determining the counting efficiency of the system for the energies being analyzed, and careful consideration of counting times in order to determine the minimum detectable activity (MDA). In the United States, the MDA must be less than 0.005 µCi. A comprehensive discussion of leak test considerations is given in Thomadsen (2000, pp 21–28). Figure 1 demonstrates localizing the activity distribution within 137Cs tubes. This was accomplished using Gafchromic RTQA film (International Specialty Products, Wayne, NJ). First, the position of each source was drawn on the film surface with lines showing the axis and endpoints of each tube assembly, along with the serial numbers. Then, the sources were laid on the film. After approximately 10 minutes of exposure for these nominal 20 mgRaeq sources, the self-developed autoradiograph shows the location of the distribution relative to the physical source. Short-lived Sources For short-lived sources such as 125I seeds, TG-56 recommends the following steps: 1. Document the physical/chemical form and encapsulation based on the manufacturer’s specification in order to support dosimetry calculations.
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2. Verify the uniformity of the activity distribution within each source, where applicable, or the distribution of seeds along an extended ribbon. 3. Calibrate the sources, either individually or as a batch. For large batches of loose seeds, TG-56 allows for “secondary traceability by statistical inference” in which a “suitable random sample” is calibrated with secondary traceability. TG-40 calls for that sample to comprise at least 10% of the total batch. Some institutions have developed procedures for measuring all the seeds in a batch together, applying a correction factor that depends on the batch size (Lee et al. 1999; Mellenberg and Pennington 1999). A disadvantage of batch measurements is that one or more significant outlier sources may be lost in the background. For large numbers of seeds in ribbons, TG-40 calls for measuring a minimum of 10% of the seeds or two ribbons, whichever is larger. It is possible to obtain inserts for well chambers that allow ribbons to be coiled into reproducible positions, but it may be necessary to develop correction factors for ribbons of different lengths (Thomadsen et al. 1999) The same as with long-lived sources, the expectation is that the local calibration should agree with that provided by the manufacturer within 3%. The range of source strengths should not exceed 5%, but TG-56 acknowledges that verifying this for a large batch of seeds may not be practical. For sources purchased in a sterile assembly, TG-40 recommends, “purchasing and calibrating a single (non-sterile) seed for each designated-strength grouping.” This problem of sterile assemblies deserves additional discussion. Calibrating a loose seed as a surrogate for the actual seeds implanted assumes that there is some meaningful connection between the manufacturing process that produced the loose seed and that for the assembly. For example, if the loose seed comes from the same manufacturing lot as those used to create the assembly, then one might consider it a random sample of the seeds used to treat the patient. If the loose seed comes from a different lot, but the manufacturer declares that it has passed through the same
Figure 1. Autoradiograph of 137Cs tube sources using Gafchromic RTQA film. The exposure time was approximately 10 minutes for these nominal 20 mgRaeq sources. Left: lines drawn to show the axis and endpoints of the sources, along with the serial numbers. Middle: sources placed on the prepared lines. Right: self-developed densities indicating the radioactive material distribution.
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calibration process as those used to create the assembly, then the connection is weakened, but may still be sufficient. There are several ways to deal with the problem of assuring the calibration of seeds that come in sterile assemblies. 1. Purchase and calibrate some loose seeds of the same type and activity. 2. Purchase additional assemblies, which are then calibrated and discarded. 3. Break sterility on some assemblies, which are then calibrated and re-sterilized. 4. In the operating room (OR), calibrate some of the assemblies under sterile conditions (Feygelman et al. 1996; Butler et al. 1998). 5. In the OR, remove some of the seeds from the assemblies under sterile conditions, and then go to a non-sterile area to calibrate and subsequently discard them. 6. Purchase the seeds from the manufacturer and have them sent to a third party that can provide an independent calibration and sterilize the assemblies. The third-party supplier of pre-loaded needles or stranded seeds may be asked to hold aside a few loose seeds from the manufacturer’s shipment for end-user calibration, sterilization, and possible implantation. None of these options is clearly best. The loose seeds in option (1) above may not be representative of the seeds implanted. Purchasing material for calibration only as in (2) or (5) adds expense, and (5) treats less of the patient because the assemblies are shortened. Resterilization (3) may not be possible or permitted by the package insert. Sterile calibration (4) is probably the best solution but has some logistical challenges. The last option listed, (6), provides an independent calibration that is not under the control of the local institution and with no assurance of the third party’s qualifications of traceability. Whether or not that is acceptable is somewhat debatable and is presently under review by an AAPM working group. Note that leak testing short-lived sources is not usually necessary, since they are typically not kept in active inventory for longer than 6 months. TG-56 does recommend leak-testing 125I seeds before using them for a second patient or after extensive handling, since the titanium encapsulation is relatively fragile and the radioactive material highly volatile.
Periodic Testing For sources that are maintained in inventory, the quality management program should include some periodic activities. TG-56 recommends the following steps: 1. A formal inventory check needs to be completed quarterly; this is a Nuclear Regulatory Commission (NRC) requirement that applies to federal and state licensees. Note that this could apply to 125 I seeds if an inventory is maintained. 2. The calibration of each source should be checked annually and the identification system (e.g., color-coding) renewed, if applicable. Deviations from the apparent half-life may indicate contamination with a different isotope. Should any such deviation be found and confirmed, the dose distribution around the source may differ from standard tables. Assessing that dose distribution is beyond the capabilities of most clinics, so any such sources would typically be returned to the manufacturer. 3. Leak testing should be performed at the intervals required by the radioactive materials license. Note that this could also apply to 125I seeds kept in service for longer than 6 months (an unlikely but not impossible situation.)
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Rechecking the uniformity of the activity within the source, while not specifically recommended by the AAPM, is another possible element of the QM program.
Per-Patient Testing For long-lived sources maintained in an inventory, such as 137Cs tubes, each time the sources are used there should be a mechanism for checking the proper selection and arrangement of the sources. Typically, this is accomplished by using a color-coding system. Where staff resources permit, one can have one person build the source assembly and another check it. It is a regulatory requirement that an inventory system be established that records the use of the sources. The system should clearly show which sources are out of storage and where they are. The importance of inventory control cannot be overstated. Dealing with a lost source is best read about, not experienced firsthand. The appendix shows worksheets developed at Wheeling Hospital, Wheeling, WV, to aid in specifying the sources to be used in an implant and in documenting the inventory. For short-lived sources that are obtained for a specific patient’s treatment, the tests listed under “Acceptance Testing” are both necessary and sufficient.
Conclusion The topic for this chapter has been the quality management system for the sources themselves. The table below summarizes the recommendations. It should be apparent to any physicist with some clinical experience that this element of the overall quality management program is necessary, since an error here can affect a large number of patients. Most errors that do occur, however, relate to how the sources are used. Thus, quality management of the sources is where vigilance in brachytherapy begins.
52
Gary A. Ezzell Table 1. Quality Control Tests for Low Dose Rate Brachytherapy Sources Long-lived sources, maintained in an inventory Test
Physical/chemical form; encapsulation Check integrity with leak test
Activity location and uniformity of distribution Verify identification system Calibration Inventory
Method
Frequency
Shipment documentation
Acceptance
Leak test
Acceptance; then every 6 months unless otherwise specified by license Acceptance; optionally at annual calibration Acceptance; annually; each use Acceptance; annually Acceptance; quarterly; each use
Radiograph/autoradiograph Check color-coding or equivalent Well-chamber Inspection
Short-lived sources, ordered for specific patients Physical/chemical form; encapsulation Activity location and uniformity of distribution (where applicable, such as wires or source trains) Calibration
Shipment documentation
Acceptance
Inspection or autoradiograph
Acceptance
Well-chamber
Acceptance
(Adapted from Kutcher et al. 1994; Nath et al. 1997; and Thomadsen 2000).
Appendix Worksheets for specifying the sources to be used in an implant and in documenting the inventory. (Courtesy of Dr. Wayne Butler, Wheeling Hospital, Wheeling, WV.)
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53
Form 1: Radioactive Materials Requistion
Radioactive Materials Requisition Patient: ____________________________________
Room: ____________________
( ) Cesium - 137 Tandem Loading (mg Ra eq.)
Upper, inner end ________
________
________
________
________
________
________
________
Outer end ________
________
Color
————— Spacers —————
10 = green
—————
15 = yellow
—————
20 = orange
————— Color
Ovoid Loading Right _________ ________
(mg Ra eq.)
Left __________ ________
( ) Iridium - 192 Number of Ribbons
No. of Seeds per Ribbon
Seed Spacing (mm)
Activity per seed (mg Ra eq) (mCi)
Ribbon Color
Date and Time needed: ________________________________ Radiation Oncologist: ___________________________________ Date: _______________________ Requisition filled by: ___________________________________ Date: _______________________ Checked by: __________________________________________ Date: _______________________
54
Gary A. Ezzell Form 1: Radioactive Materials Requistion (continued)
Radioactive Materials Requisition Iridium - 192 (>4 ribbon types) Patient: ____________________________________ Number of
No. of Seeds
Seed Spacing
Ribbons
per Ribbon
mm
Room: ______________________ Activity per seed
mg Ra eq
mCi
Ribbon Color
Date and Time needed: ________________________________ Radiation Oncologist: ___________________________________ Date: _______________________ Requisition filled by: ___________________________________ Date: _______________________ Checked by: __________________________________________ Date: _______________________
5–Quality Management of Low Dose Rate Brachytherapy Sources Form 2: Cesium In-Use Inventory
Cesium In-Use Inventory
Patient: ________________________________________________
Room: _____________
Sources removed: ________________________________________
Date: _____________
________________________________________
Time: ____________
Cesium in Hot-Cell Storage Safe Quantity After Removal
Quantity After Return
Color Code
Quantity Stocked
Identity
Green
6
10 mg Radium equivalent
Yellow
6
15 mg Radium equivalent
Orange
4
20 mg Radium equivalent
←
Authorized Handler Initials
→
Sources returned: _________________________________________
Date: _____________
_________________________________________
Time: _____________
*********************** Authorized Handlers:
(List here)
55
56
Gary A. Ezzell Form 3: Iridium In-Use Inventory Iridium-192 Inventory Log Receipt Manufacturer: _____________________ Received by: _________________ Survey instrument check
Package Damaged?
Package Wipe Test 0.99 ?
is tdwell(120)/120 1.000 ± 0.010 ?
Source Strength Calibration 1.
Setup control console with L=995 for 5F/6F caths in HDR 1000+ and 10 seconds each for positions 14-22. Use 200 nA scale on the Keithley 6517A and the AVG(10) filter function.
Position 22 21 20 19 18 17 16 15 14 Distance [mm] 942.5 945.0 947.5 950.0 952.5 955.0 957.5 960.0 962.5 Rdg Determine maximum reading (MAXRdg) ____________________ at time: ________________ 2. Calculate in-house measure of air kerma strength (NEMCSK); 192Ir t½ = 73.83 days with Ȝ = 0.009388 days–1: Decay factor (NEMCF) to 0:00 for the current day _____________> 1, typically ~ 1.0071. Measure temperature _________ [ºC] and pressure ___________ [mbar] and calculate CTP ___________ Standard Imaging HDR-1000 well chamber S#A940704 calibrated 2/27/2003 CWELL = 0.5035 mGy m2 h–1 nA–1 Keithley 6517A electrometer S#0857926 calibrated 9/15/2003 CEL = 0.999 A Rdg–1 on 200 nA scale or Keithley MK 602 (CNMC) electrometer S#34635A calibrated 1/25/2002 CEL = 0.0981 nA Rdg–1 10–8 A scale
SK = MAXRdg x NEMCF x CTP x CWELL x CEL = _____________ [mGy m2 h–1] For air kerma rate @ 1 meter (as required by license) divide above number by 10 _____________ [cGy/h] NEMC
Using a conversion factor of 2.45 Ci-h/cGy, enter Aapp [Ci] on the yellow placard on the RAL and console ________ if not already done by the Nucletron Service Engineer.
Figure 2. New Source QA.doc. page 1
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Tufts-New England Medical Center, Department of Radiation Oncology New HDR 192Ir Source QA 3. Comparison of in-house determination of air kerma strength with source strength reported by Mallinckrodt. The microSelectron remote afterloader (RAL), model # 080.000 serial # 9511, is made by Nucletron BV. The source is manufactured by Mallinckrodt Medical BV, model CI L BV. Source ID # ____________________ Calibration date __________________________ Mallinckrodt air kerma strength (VENDORSK) calibration [mGy m2 h–1] ÷ 10 ________________ [cGy m2 h–1] Decay factor (VENDORF) to Mallinckrodt calibration (X days +5 hours due to timezone differences) ____________< 1
Calibration ratio
=
NEMC VENDOR
SK
F ×VENDOR S K
=
_____________ Is this ratio 1.00 ± 0.05? ___________
If yes, 1. update source strength (NEMCSK) in the microSelectron treatment console (submode 11) compare predicted treatment time for a standardized VagCyl treatment plan with catalog 2. update source strength (NEMCSK) in the Plato v.14.1.2 planning system (passwd=hdr) 3. update source strength (NEMCSK) in the Pinnacle planning system (passwd=physics) 4. update source strength (NEMCSK) in the MammoSite Excel spreadsheet 5. make 2 new decay tables and post at the planning workstation and in HDR console logbook
___________ ___________ ___________ ___________ ___________
If ratio is not within 5% of unity, contact another Physicist or the RSO and consider cross-calibration using the IVB-1000 well chamber.
Safety Checks
(Y or N)
1. HDR 192Ir source retraction following AC power failure simulation?
___________
2. Door light and Primalert function with AC power off using Cs micrad check source? (Unplug “Connector In” from PRIMAPAK)
___________
3. Emergency STOP (on wall near console) functions with printout of irradiation time?
___________
4. Door interlock functions with printout of irradiation time?
___________
137
5. Interrupt button on the console functions with printout of irradiation time?
___________
6. Survey meter battery check and constancy check OK?
___________
7. Was survey meter calibrated within 365 days ? (If not, notify Health Physics)
___________
8. Are Emergency Procedures in HDR Log Book?
___________
9. Is error 129 detected when indexer ring is not latched?
___________
10. Is error 129 detected when applicator is not inserted?
___________
11. Is error 129 detected when interstitial interconnect tube is installed without catheter present?
___________
12. Review of daily checks since last monthly/new source QA?
___________
13. Daily checks performed? (attach and complete QA-Daily.doc)
___________
Additional Notes (if needed): Authorized Physicist Name _____________________________ Signature/Date _____________________________
Figure 2. New Source QA.doc. page 2
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Mark J. Rivard and Christopher S. Melhus
printed copies of these forms, and hardcopies of the irradiated films, will soon be replaced with an electronic spreadsheet-based approach that is paperless (see Figure 3). Based on past inspections by the Massachusetts Radiation Control Program, sole use of electronic records is an acceptable practice. While somewhat quite painful, it is recommended by AAPM TG-56 that QA of device-related equipment (catheters, applicators, tubes, dummy wires, etc.) be performed once every 3 months. In addition to confirming device integrity and inventory, as these items occasionally get moved or are lost just when they are needed for that rare and complex procedure, one should physically determine that the device performs according to expectations. For instances, the metal nibs on a dummy marker can become uncrimped and slide along the wire. When used in vivo, misleading distances may be obtained and the whole HDR 192Ir brachytherapy procedure could be aborted, or worse yet, could proceed incorrectly due to erroneous information. Similarly, the parts for an HDR vaginal cylinder kit may get mixed up with an HDR tandem-and-ovoid kit. In addition to being misplaced, possibly in the operating room (OR) storage room, parts could become damaged if subjected to incompatible sterilization procedures (e.g., flash versus gas). Monthly QA While no longer the accepted practice in the U.S. with the advent of the revised Code of Federal Regulations 10 CFR part 35, our license demands that patients not be treated if monthly or new source QA was performed more than 31 days ago. When patients are not scheduled for treatment, there are occasions when we will actively decide to let the date for QA creep into the next month so as to exert resources elsewhere. This approach is feasible at our institution since we can generally complete the QA tasks outlined below in the Monthly QA form (Figure 4) within a matter of 4 hours since the RAL is in a vault with a seldom-used linac. Toward resources management, we are working with the Massachusetts Radiation Control Program to resolve the monthly versus quarterly QA requirements, and integration of NRC regulations into our agreement state regulations. Daily QA Based upon the level of staffing at a given facility, either the medical physicist or the radiation therapy technologist (RTT) performs daily QA tests preceding patient treatments. At some facilities, a trained radiotherapist will perform the daily QA tests and the physicist will occasionally assist the RTT team by performing the daily QA toward hastening patient treatments and coordinating departmental activities. The TG-56 report provides guidance on core daily QA tests recommended for HDR brachytherapy facilities. An example of a Daily QA checklist used at Tufts-NEMC is provided in Figure 5. Treatment-related QA A form, Treatment QA (Figure 6), has been in use for many years at Tufts-NEMC to coordinate responsibilities of all parties preceding, during, and immediately after HDR brachytherapy. This checklist approach ensures that tasks are appropriately delegated and consistently performed. We keep this completed form in the patient chart, along with a printed record of the treatment that has been signed by the physician, treatment plan, and written directive. QA for Imaging-related Equipment Simple QA tasks of image quality, spatial accuracy and orientation, and accurate data depiction can and should be performed during the patient simulation phase. Sometimes a therapist will mistakenly label a film as right anterior oblique instead of left anterior oblique, or record the wrong source-to-film distance value. These mistakes can significantly delay the treatment-planning process for the physicist, and caution versus cynicism is recommended. However, these simple tasks must supplement a rigorous, regularly performed QA program for the imaging component of the brachytherapy simulation process.
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HDR brachytherapy is a complex endeavor that relies on a multitude of imaging equipment to properly deliver the therapy. While somewhat outside the scope of this chapter, it goes without saying that “garbage-in is garbage-out” regarding the quality of imaging data used to prepare an HDR brachytherapy treatment plan. QA of the simulator unit is typically performed monthly, sometimes only quarterly. However, more sophisticated plans rely on 3-D computed tomography (CT)- and magnetic resonance imaging (MRI)-based image datasets. With physicist specialization, it is often the case that the physicist primarily involved with brachytherapy planning and delivery will not be the one performing QA of the imaging equipment. Remember that brachytherapy is limited in that it does not have the in vivo delivery confirmation feature that teletherapy does through port-films, electronic portal imaging devices, or kVpbased image-guided radiation therapy. Use of a diode integrating output or surface dose at a point is not a substitute. Treatment-Planning System and Related Equipment QA After the patient has been simulated, the clock is often ticking and the physicist is under pressure to complete the radiation treatment plan (RTP). Everyone, including the sick patient, is waiting to start the treatment. This is not the time to feel uncertain about your HDR QA program and possible inadequacies of RTP QA. At a minimum, consistency of RTP performance regarding source decay and calculation of point doses results. At our institution and many others, an atlas of preplanned HDR treatment plans is available for the purposes of QA and for facilitating quick implementation of relatively simple multifractionated brachytherapy implants. For our vaginal cylinder atlas, plans are available for a wide range of doses, cylinder diameters, and prescription points. We use air kerma strength at Tufts-NEMC, and simply taking the ratios of the source strengths permits a physicist opportunity to check the corrected treatment times (both total time and individual dwell times) with the times calculated by the HDR control console (in computer memory) and HDR RTP system. Of course, it is senseless to perform all these tests, and not document your efforts. Strive for an electronic document (backed up occasionally) that permits easy data entry and record keeping. In addition to tests of the RTP system, QA tests must be performed of the associated peripherals such as the digitizer, printer, treatment card, or export function, etc. One must have confidence during the treatment-planning process that printed and exported results are accurate and as expected. An example of comprehensive QA tests of modern RAL equipment, albeit for LDR 125I, is provided by Rivard, Evans, and Kay (2005) where suitability of various AAPM task groups is considered.
Emergency Procedures The AAPM TG-59 report provides guidance toward the development of adequate HDR brachytherapy emergency procedures. Emergency procedures should be able to direct treatment personnel through the worst-case scenario accident, e.g., the HDR 192Ir source becoming detached from the source guide wire and remaining inside the catheter or applicator. In some extreme instances, source retrieval from the patient could include surgical intervention. However, it is more likely for minor emergencies to occur, which include, but are not limited to, loss of power and abnormal performance of the RAL. Because of the severe consequences that could occur after only a short period of time, the NRC directs the licensee to provide appropriate staff and equipment during HDR brachytherapy. Hence, the presence of an authorized medical physicist and authorized user is often required during HDR brachytherapy treatment.
2.399 2.376 2.354 2.332 2.310 2.289 2.267 2.246 2.225 2.204 2.184 2.163 2.143 2.123 2.103 2.084 2.064 2.045 2.026 2.007 1.988 1.969 1.951 1.933 1.915 1.897 1.879 1.862 1.844 1.827 1.810 1.793
2.387 2.365 2.343 2.321 2.299 2.278 2.257 2.236 2.215 2.194 2.173 2.153 2.133 2.113 2.093 2.074 2.054 2.035 2.016 1.997 1.979 1.960 1.942 1.924 1.906 1.888 1.870 1.853 1.835 1.818 1.801 1.785
7/5/05 7/6/05 7/7/05 7/8/05 7/9/05 7/10/05 7/11/05 7/12/05 7/13/05 7/14/05 7/15/05 7/16/05 7/17/05 7/18/05 7/19/05 7/20/05 7/21/05 7/22/05 7/23/05 7/24/05 7/25/05 7/26/05 7/27/05 7/28/05 7/29/05 7/30/05 7/31/05 8/1/05 8/2/05 8/3/05 8/4/05 8/5/05
1.472 1.458 1.445 1.431 1.418 1.405 1.391 1.378 1.366 1.353 1.340 1.328 1.315 1.303 1.291 1.279 1.267 1.255 1.243 1.232 1.220 1.209 1.197 1.186 1.175 1.164 1.153 1.142 1.132 1.121 1.111 1.100
1.465 1.452 1.438 1.425 1.411 1.398 1.385 1.372 1.359 1.346 1.334 1.321 1.309 1.297 1.285 1.273 1.261 1.249 1.237 1.226 1.214 1.203 1.192 1.181 1.170 1.159 1.148 1.137 1.126 1.116 1.106 1.095
Figure 3. Decay table for an HDR 192Ir brachytherapy source (1266.xls). continued on next page
5/14/05 5/15/05 5/16/05 5/17/05 5/18/05 5/19/05 5/20/05 5/21/05 5/22/05 5/23/05 5/24/05 5/25/05 5/26/05 5/27/05 5/28/05 5/29/05 5/30/05 5/31/05 6/1/05 6/2/05 6/3/05 6/4/05 6/5/05 6/6/05 6/7/05 6/8/05 6/9/05 6/10/05 6/11/05 6/12/05 6/13/05 6/14/05
3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05
3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082
3.890 3.854 3.818 3.782 3.747 3.712 3.677 3.642 3.608 3.575 3.541 3.508 3.475 3.443 3.411 3.379 3.347 3.316 3.285 3.254 3.224 3.194 3.164 3.134 3.105 3.076 3.047 3.019 2.991 2.963 2.935 2.908
3/23/05 3/24/05 3/25/05 3/26/05 3/27/05 3/28/05 3/29/05 3/30/05 3/31/05 4/1/05 4/2/05 4/3/05 4/4/05 4/5/05 4/6/05 4/7/05 4/8/05 4/9/05 4/10/05 4/11/05 4/12/05 4/13/05 4/14/05 4/15/05 4/16/05 4/17/05 4/18/05 4/19/05 4/20/05 4/21/05 4/22/05 4/23/05
3.908 3.872 3.836 3.800 3.764 3.729 3.694 3.660 3.625 3.592 3.558 3.525 3.492 3.459 3.427 3.395 3.363 3.332 3.301 3.270 3.239 3.209 3.179 3.149 3.120 3.091 3.062 3.033 3.005 2.977 2.949 2.921
DATA ENTRY Date of Strength Calibration
HDR 192Ir source #D35A-1266 = 3.908 cGy m2/h at 00:00:00 3/23/2005 by CSM length 995 mm 73.83 day HL Strength (cGy m2/h) Strength (cGy m2/h) Strength (cGy m2/h) date MIDNIGHT NOON date MIDNIGHT NOON date MIDNIGHT NOON
84 Mark J. Rivard and Christopher S. Melhus
1.776 1.760 1.743 1.727 1.711 1.695 1.679 1.663 1.648 1.632 1.617 1.602 1.587 1.572 1.557 1.543 1.528 1.514 1.500 1.486
1.768 1.751 1.735 1.719 1.703 1.687 1.671 1.655 1.640 1.625 1.609 1.594 1.579 1.565 1.550 1.536 1.521 1.507 1.493 1.479
8/6/05 8/7/05 8/8/05 8/9/05 8/10/05 8/11/05 8/12/05 8/13/05 8/14/05 8/15/05 8/16/05 8/17/05 8/18/05 8/19/05 8/20/05 8/21/05 8/22/05 8/23/05 8/24/05 8/25/05
1.090 1.080 1.070 1.060 1.050 1.040 1.030 1.021 1.011 1.002 0.992 0.983 0.974 0.965 0.956 0.947 0.938 0.929 0.921 0.912
1.085 1.075 1.065 1.055 1.045 1.035 1.026 1.016 1.006 0.997 0.988 0.979 0.969 0.960 0.951 0.942 0.934 0.925 0.916 0.908
Figure 3. Decay table for an HDR 192Ir brachytherapy source (1266.xls). continued from previous page
6/15/05 6/16/05 6/17/05 6/18/05 6/19/05 6/20/05 6/21/05 6/22/05 6/23/05 6/24/05 6/25/05 6/26/05 6/27/05 6/28/05 6/29/05 6/30/05 7/1/05 7/2/05 7/3/05 7/4/05
3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05 3/23/05
3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082 3.9082
2.880 2.854 2.827 2.800 2.774 2.748 2.723 2.697 2.672 2.647 2.622 2.598 2.574 2.550 2.526 2.502 2.479 2.456 2.433 2.410
4/24/05 4/25/05 4/26/05 4/27/05 4/28/05 4/29/05 4/30/05 5/1/05 5/2/05 5/3/05 5/4/05 5/5/05 5/6/05 5/7/05 5/8/05 5/9/05 5/10/05 5/11/05 5/12/05 5/13/05
2.894 2.867 2.840 2.814 2.787 2.761 2.736 2.710 2.685 2.660 2.635 2.610 2.586 2.562 2.538 2.514 2.490 2.467 2.444 2.421
DATA ENTRY Date of Strength Calibration
HDR 192Ir source #D35A-1266 = 3.908 cGy m2/h at 00:00:00 3/23/2005 by CSM length 995 mm 73.83 day HL Strength (cGy m2/h) Strength (cGy m2/h) Strength (cGy m2/h) date MIDNIGHT NOON date MIDNIGHT NOON date MIDNIGHT NOON
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Tufts-New England Medical Center, Department of Radiation Oncology Monthly HDR 192Ir Source QA Autoradiograph Tape clear 6F endobronchial catheters to readypack V-film with two GYN dummy #1 inserted. Place on simulator table at 100 cm SFD with 4 x 25 cm2 field size (no wires present) and take 3 exposures (small focus, 75 kVp, 320 mAs). Remove dummy markers and bring to HDR unit. Program HDR control console with step=5.0 mm; L=995 mm (separately); and t=0.2 second for positions below (tIRR = 3.0 s total). Autoradiograph film without bolus to minimize scatter. Process film, and analyze shifts within 0.5 mm.
Position1 1000 mm 6F Shift [mm] 1
3
5
1
7
9
3
11
5
9
11
13
15
17
13 15 17 19 21 23 25 27 29 31
19
23
25
27
29
31
33
33 35 37 39 41 43 45
ŶŶŷŶŷŶŷŶŷŶŷŶŶŷŶŷŶŷŶŷŶŷŶŷŶŷŶŷŶŷŶŷŶŶŷŶŷŶŷŶŷŶŷŶŷŶŷŶŷŷŷŷŷ Perform two autoradiographs using tdwell = 3 s and 30 s to evaluate source homogeneity.
Source homogeneity acceptable (visual assessment) ? ___________ Average measured offset _______________ mm (range 0.0 ± 1.0 mm) If offset is within range, write “NA” below. Otherwise enter the correct L (z 995 mm) and complete the following 3 tasks: Update daily check program with new length L Update program cards for any ongoing patients with new length L Post new length L in planning room
___________ ___________ ___________
Timer Offset, Linearity, and Accuracy Insert a clear, 6F catheter into HDR-1000. Setup control console with L=995, step 2.5 mm, position=18, and variable dwell time (tdwell). Use electrometer without filter, fit to straight-line, and determine toffset and time linearity (R2). Use stopwatch to independently measure tdwell(120) accuracy. If toffset, R2, or tdwell(120) are not within specifications, have second physicist review results. If they are still outside specification, contact Nucletron.
tdwell [seconds] 0.1 1.0 10.0 0.5 5.0 0.2 2.0 20.0
measured charge (2 PC scale)
what is the value of toffset ?
what is the value of R2 ?
what is the value of tdwell(120)/120 ?
is toffset –0.55 ± 0.03 s ?
is R2 > 0.99 ?
is tdwell(120)/120 1.000 ± 0.010 ?
Source Strength Calibration 1. Setup control console with L=995 for 5F/6F caths in HDR 1000+ and 10 seconds each for positions 14-22. Use 200 nA scale on the Keithley 6517A and the AVG(10) filter function. Position 22 21 20 19 18 17 16 15 14 1000 mm 6F Determine maximum reading (MAXRdg) ____________________ at time: ________________ 2. Calculate Monthly QA air kerma strength (SK); 192Ir t½ = 73.83 days with Ȝ = 0.009388 days–1: Decay factor (NEMCF) to 0:00 for the current day _____________> 1, typically ~ 1.0071. Measure temperature _________ [ºC] and pressure ___________ [mbar] and calculate CTP ___________ Std Imaging HDR-1000 well chamber S#A940704 good until 2/17/2006 CWELL = 0.5029 mGy m2 h–1 nA–1 Keithley 6517A electrometer S#0857926 good until 9/15/2005 CEL = 0.999 nA Rdg–1 on 200 nA scale or Keithley MK 602 (CNMC) electrometer S#34635A good until 2/9/2006 CEL = 0.0998 nA Rdg–1 10–8 A scale Monthly QA SK = MAXRdg x NEMCF x CTP x CWELL x CEL = _____________ [mGy m2 h–1] For air kerma rate @ 1 meter (as required by license) divide above number by 10 _____________ [cGy/h] 3. Comparison of Monthly QA determination of SK with tabulated value from decay chart. Source ID # ____________________ Calibration date __________________________ Tabulated (Midnight) SK value from decay table _________________ Ratio- divide Tabulated by air kerma rate @ 1 meter __________ Acceptable (Y/N)? _____ (1.000 r 0.020)
Figure 4. Monthly QA.doc. page 1
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Tufts-New England Medical Center, Department of Radiation Oncology Monthly HDR 192Ir Source QA Check of Guide Tubes GYN #1: length = __________ mm Acceptable [748.5 mm +/- 1.0 mm]? GYN #2: length = __________ mm Acceptable [748.5 mm +/- 1.0 mm]? GYN #3: length = __________ mm Acceptable [748.5 mm +/- 1.0 mm]? Is the mechanical integrity of all applicators, connectors, and guide tubes acceptable based on visual inspection? Safety Checks
___________ ___________ ___________ ___________ (Y or N)
1. HDR 192Ir source retraction following AC power failure simulation?
___________
2. Door light and Primalert function with AC power off using Cs micrad check source? (Unplug “Connector In” from PRIMAPAK)
___________
3. Emergency STOP (on wall near console) functions with printout of irradiation time?
___________
4. Door interlock functions with printout of irradiation time?
___________
5. Interrupt button on the console functions with printout of irradiation time?
___________
6. Survey meter battery check and constancy check OK?
___________
7. Was survey meter calibrated within 365 days ? (If not, notify Health Physics)
___________
8. Are Emergency Procedures in HDR Log Book?
___________
9. Is error 129 detected when indexer ring is not latched?
___________
10. Is error 129 detected when applicator is not inserted?
___________
11. Is error 129 detected when interstitial interconnect tube is installed without catheter present?
___________
12. Review of daily checks since last monthly/new source QA?
___________
13. Daily checks performed? (attach and complete QA-Daily.doc)
___________
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Additional Notes (if needed):
Authorized Physicist Name _____________________________ Signature/Date _____________________________
Figure 4. Monthly QA.doc. page 2
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Tufts-New England Medical Center Department of Radiation Oncology Daily HDR 192Ir Source QA The microSelectron remote afterloader, model # 080.000 serial # 9511, is made by Nucletron BV. The source is manufactured by Mallinckrodt Medical BV, model CI L BV. Name: ____________________________ Signature: _________________________
Date: _______
1. Are the TV monitors at the control console operational? 2. Is the intercom operational? (CL-4 must be powered) 3. Does the radiation area monitor function? 4. Does the radiation survey meter function? 5. Is the emergency basket stocked and the emergency storage container in place?
_______ _______ _______ _______ _______
[forceps/tongs, cutters/scissors, suture removal kit, and “CAUTION RADIATION AREA” tape]
6. Are the source guide tubes free of kinks, etc? 7. Did control console pass self-diagnostic checks when key switch is first turned? 8. Is the emergency STOP button in the treatment room operational?
_______ _______ _______
[verify illumination of “Activate” light on Emergency Stop box near console]
9. Is the emergency STOP button on the HDR unit operational?
_______
[verify illumination of “Activate” light on Emergency Stop box near console]
Attach source position check ruler to channel 2 (incorrect channel). Turn on ceiling light and adjust TV camera so that check ruler is clearly visible. Run program card #1. This will program: Positions 4, 8, 12, 16, 20 Length 995 Step size 2.5 mm Dwell times at least 3 sec for positions 8,12,16,20 at least 20 sec for position 4 10. Confirm program and dwell times, press START to initiate treatment, and verify that Error Code 0129 (or 014) appears on printout and that the source is not exposed.
_______
Attach the source position check ruler to channel 1 (correct channel). Press START to initiate treatment. 11. Do the ends of the live and dummy sources align with the marked distance indicators to within 1 mm of 950, 960, 970, 980, and 990?
_______
While the source is dwelling at position 4, check the door light, interrupt key, door interlock, and emergency stop. The source should return to position 4 (end at 990) after each restart. 12. Is the “HDR IN USE” door light on? 13. Is the door interlock operational? 14. Does the interrupt button on the console function properly? 15. Is the emergency STOP button near the HDR console operational?
_______ _______ _______ _______
On the printout: 16. Is the date correct? 17. Is the time correct? 18. Is the source strength on the printout within 0.02 cGy m2/h of the posted activity? 19. Is the paper supply in the console printer sufficient for the day's treatment?
_______ _______ _______ _______
Figure 5. Daily QA.doc.
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Figure 6. Treatment QA.doc.
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System Failures Physical Plant Emergencies For natural disasters such as fires, earthquakes, tornadoes, or floods, the brachytherapy treatment team should generally stop HDR treatment and follow hospital or clinic-specific emergency procedures. Even if electrical power is interrupted, RAL units are designed with sufficient redundancy (battery backup) to retract the HDR 192Ir source and to record the amount of treatment completed using a backup battery or power pack. The medical physicist should include regular tests of backup power systems through QA tests. AAPM TG-56 recommends quarterly testing of backup power systems. Minor Emergencies Minor emergencies arise through abnormal operations of the RAL unit. These may include kinked treatment catheters, catheter obstruction, loose guide tube connections, operation of treatment vault doors during treatment, patient motion or discomfort during treatment, or other similar events. Manufacturers of HDR brachytherapy devices incorporate displayed messages and error codes for the operator to interpret the root cause of a minor emergency. Training should prepare hospital brachytherapy personnel to recognize common error codes and/or enable them to interpret appropriate reference manuals to determine the meaning of a given error code. In addition, operators must be able to determine whether treatment should continue in the wake of a minor emergency. Typically, minor emergencies do not result in misadministrations to the patient, and treatment may resume after correcting the problem. Major and Medical Emergencies A major emergency is one that can only be resolved through physical intervention of a member of the treatment team. For example, source retraction failure, controller (computer) failure, or patient medical emergency would qualify as a major emergency. As this type of emergency requires entry into the treatment vault while the source is out of the safe, the potential exists for very high personnel and patient doses. The dose rate can be over 7 Gy per minute at a distance of 1 cm from a 10 Ci 192Ir source. Therefore, individuals should be trained to work quickly and efficiently during the resolution of a major emergency. RAL manufacturers often prepare a single-page instruction sheet on how to retract the source manually in the event of a major emergency. All treatment personnel should be trained to perform this task and retraining should be offered at least annually. Ideally, source retraction/retrieval would occur within 1 to 2 minutes of a major emergency to limit detrimental outcomes to the treatment personnel and to the patient. The medical physicist should maintain an emergency kit near the HDR treatment vault. At a minimum, this kit should include tools for transferring a loose HDR source to a backup shielded container, a suture removal kit, and appropriate signage to secure a room after a major emergency. The contents of an emergency kit should be checked before every treatment. An example of an Emergency Procedure form is included in Figure 7. When printed in color and posted in the clinic, the form facilitates rapid understanding of the proper actions to manually retract the HDR 192Ir source for personnel responding to an emergency.
Staffing, Credentialing, Assignments, Equipment, and Training Frequency Radiation Safety Officer (RSO) While the RSO generally does not directly operate the HDR brachytherapy device, the RSO is directly responsible for the safe use of radiation administration at an institution. In addition, the RSO communicates with regulatory agencies regarding the administration of a radioactive materials program. As a result, the RSO shall review and approve procedures and emergency plans. However, minor changes to forms for improvements to workflow, clarity, or practicality need not go through a laborious review process.
7–Quality Assurance for High Dose Rate Remote Afterloaders
Figure 7. Emergency Procedure.doc.
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Physician The radiation oncologist should be board-certified, have substantial expertise and training in brachytherapy, and be licensed (authorized) to operate an HDR 192Ir unit. While the physician is most knowledgeable about clinical aspects such as diagnoses, prescription doses, and short-/long-term effects, the radiation oncologist should have annual emergency procedure training and QMP review, and be shown the practical operation techniques of the RAL at least annually. Without a high frequency and running the risk of “crying wolf,” it is often humbling to call for surprise emergencies to glean how the physicians would address a medical emergency and perform their responsibilities of the emergency procedures. Radiation Therapy Technologist (RTT) The treatment unit operator is typically a radiation therapy technologist (RTT), and should have vendorsupplied training on the HDR unit or be trained by an in-house expert using a documented and RSO-approved training program. Like the physician and physicist, the RTT should have annual emergency procedure training and should review the QMP for a general sense of the program goals. Furthermore, the treatment-unit operator should have the opportunity to practice operating the machine before administering patient treatment. This is why it is not suggested that the medical physicist perform the daily QA as a favor since this gesture could inadvertently result in a misadministration. Medical Physicist AAPM TG-59 recommends that one full-time equivalent medical physicist be allocated for an average of 10 HDR fractions per week, including treatment planning, treatment supervision, and periodic QA and training. The medical physicist should be board-certified in therapeutic radiological physics and should attend vendor-supplied training for both the treatment unit and the treatment-planning system. If possible, the medical physicist should spend 1 to 3 weeks at an established HDR program to observe all aspects of HDR brachytherapy treatment and QA. Since physicists typically coordinate the training and technical aspects of the HDR QA program, sometimes the continuing training of the physicist is overlooked. In addition to the annual emergency procedure training and QMP review, the physicist should re-examine the credentialing program to assure that it is up-to-date, and that s/he would still pass the competency test. A sample test (Figure 8) for physicians, therapists, and physicists is included (with a key).
Case Reports from NRC Public Postings Towards learning from the mistakes of others, it is instructive to review the problems with HDR brachytherapy experienced at other institutions. Follows are six examples which document treatment mishaps that could have been prevented had high QA standards been enforced. Example 1: NRC Information Notice 00-05 In NRC Information Notice 00-05 titled “Recent Medical Misadministrations Resulting From Inattention To Detail,” a series of eight preventable misadventures occurring between January 1999 and March 2000 were listed, including the following two in HDR brachytherapy. A patient received an underdose to an intended site and a dose to an unintended site during a high dose rate afterloader (HDR) treatment. An error, made by a technologist when programming the parameters into the HDR, was not detected by the physicist required to verify the information keyed into the console. The technologist responsible for setting up the HDR was unfamiliar with the skip process and did not program the skip into the HDR, even though it was clearly marked on the written directive. The skip process, in this particular remote afterloader, involves the sending of the source to the end of the catheter as a starting point and then moving the source back to the designated first radiation point. Because the
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skip was not programmed, the source started the radiation treatment from the end of the catheter, resulting in a dose to an unintended site. Ineffective training and inattention to detail were considered the root causes of this misadministration. A patient received an underdose to the intended site and a dose to an unintended site during an HDR treatment. After several unsuccessful attempts to electronically transfer a patient’s treatment plan to the HDR treatment system, the licensee manually entered the treatment plan directly into the treatment system’s control station. While manually entering changes to the source dwell times, an unintended change to the source step size occurred. The licensee did not notice this change and the patient was treated using the incorrect step size resulting in the misadministration. The licensee attributed this unintended and unnoticed change in step size to a software problem. However, had the licensee properly reviewed the complete final treatment plan, which clearly showed the changed step size, this misadministration could have been prevented.
More information regarding the latter incident is available in NCR Information Notice 99-09. Example 2: NRC Significant Enforcement Action EA-04-093 In the instance described in EA-04-093, the NRC determined that a Severity Level III violation occurred when the licensee failed to ensure that the HDR source transfer tube of the correct length was attached to the treatment applicator. In summary, [the medical center] did not develop written procedures to assure the source transfer tube remained secured to the vaginal cylinder during an HDR brachytherapy treatment and thus ensure that NRC-licensed material was administered in accordance with the written directive from an authorized user physician, as required by 10 CFR 35.41, “Procedures for Administrations Requiring a Written Directive.” An authorized user physician prepared a written directive prescribing three HDR remote afterloading brachytherapy treatments to a patient after the patient received treatment from external beam radiation. Each HDR treatment was to consist of 700 rads from an Iridium-192 source of approximately 3.5 curies by vaginal cylinder. Erythema was noticed on the inner portion of each of the patient’s thighs, subsequent to the third treatment on March 11, 2004, indicating the radiation source did not transverse the entire length of the transfer tube during the HDR treatment. The authorized user physician did not expect any adverse medical effects to the patient as a result of this medical event. Corrective actions included: (1) developing written procedures to determine if the transfer tube is in place before and after treatment; (2) marking the transfer tube as part of a visual check to determine whether the transfer tube moves during treatment; and (3) training the medical physics and nursing staffs on the procedures.
Example 3: NRC Significant Enforcement Action EA-04-087 The misadventure in this example occurred when one medical physicist did not correctly implement the prescription depth requested by the authorized user in the written directive, and a second medical physicist did not review the work until two of five fractions were completed. Thus, there was a dose error of +60% for the first two fractions. This event reinforces the value of communication between the treatment planner (medical physicist or dosimetrist) and the authorized user, and the importance of treatment plan checking prior to radionuclide administration. An authorized user physician prepared a written directive prescribing five planned HDR remote afterloading brachytherapy treatments to a patient. Each treatment was to consist of 500 rads at the surface of a 25 millimeter (mm) vaginal cylinder for a total dose of 2500 rads, according to the directive prepared by the physician. A medical physicist, not realizing the physician’s written directive specified the intended dose of 500 rads per fraction was to be delivered at the surface of the cylinder, used
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KEY HDR 192Ir BRACHYTHERAPY EXAMINATION FOR PHYSICIAN AUTHORIZED USERS AND DEVICE OPERATORS 1. For HDR 192Ir, what requirements need to be satisfied to be a physician authorized user? 5 treatments, emergency procedure training, board certification or RSC & State Credentialing, 75% > this exam 2. For HDR 192Ir, what requirements need to be satisfied to be an authorized medical physicist? 5 treatments, emergency procedure training, board certification or RSC & State Credentialing, 75% > this exam 3. For HDR 192Ir, what requirements need to be satisfied to be an authorized radiation therapist? 5 treatments, emergency procedure training, registered RTT, 75% > this exam 4. Who may have access to keys to the HDR unit? Only authorized physician, physicist, and therapist 5. What is the required action should the door interlock malfunction? Guard the door entry during treatment, and contact Health Physics (x6-6168) 6. How may treatment planning data be entered into the HDR 192Ir control unit? plastic data card, standard programs in console memory, manual entry 7. According to our radioactive materials license, who must be physically present or within the audible range of normal human speech at the console area during all HDR treatments? Authorized physician, physicist, and therapist 8. What are the radiation survey requirements for HDR 192Ir procedures? Before and after treatment, survey the room and patient
Figure 8. HDR Exam3_key.doc. page 1
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KEY 9. According to our license, should the following checks be performed daily, monthly, or quarterly? a. b. c. d. e. f. g.
Check of the permanent radiation monitor Check of the TV monitor and intercom Check of treatment room door interlock Timer accuracy and linearity Check of backup battery Calibration of source Check of source homogeneity
D D D M M Q M
10. If the permanent radiation monitor is found to be inoperable, under what condition may a patient be treated? Guard the door during treatment (preferably Health Physics) 11. What is the maximum exposure level 1 meter from the surface of the HDR unit when the source is in the shielded position (in the tungsten safe)? (a) 0.25 mR/h
(b) 2.5 mR/h
(c) 25 mR/h
(d) 250 mR/h
12. What is the typical activity range of the HDR 192Ir source? (a) 4-10mCi
(b) 40-100 mCi
(c) 4-10 Ci
(d) 40-100 Ci
13. What is the typical source strength (U = 1 cGy/h/cm2) of the HDR 192Ir source? (a) 10-40 U
(b) 100-400 U
(c) 1,000-4,000 U
(d) 10,000-40,000 U
(c) 192Ir
(d) 252Cf
14. What radionuclide is used in the HDR unit? (a) 137Cs
(b) 60Co
15. What is meant by the term dwell time? Time HDR 192Ir source is in treatment positions 16. According to Nucletron’s definition, what is meant by the term secondary time? Time HDR 192Ir source is out of RAL safe 17. In the event of an emergency, what actions must be undertaken, in what order, and who is responsible for carrying them out? Therapist Physicist Physician Anyone
try all emergency off buttons enter room, crank source, survey patient remove applicator, place source in bailout pig contact Health Physics division (x6-6168)
18. Why should you not use a connecting tube for endobronchial or esophageal implants? Adding the connecting tube will not allow the source to reach the intended treatment site
Figure 8. HDR Exam3_key.doc. page 2
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Mark J. Rivard and Christopher S. Melhus the standard treatment parameter of 5 mm beyond the surface of the cylinder in developing the treatment plan. The patient was then treated on March 15, 22, and 29, 2004, according to the treatment plan prepared by the medical physicist. Prior to the fourth treatment, a second medical physicist independently reviewed the previously approved treatment plan and discovered the 5 mm error in that plan. As a result of the 5 mm error, an additional dose of 318 rads was delivered at the cylinder surface during each fractionated treatment for a total of 818 rads administered to the patient per treatment instead of the prescribed 500 rads. Corrective actions included: (1) requiring your staff to independently check treatment plans prior to each treatment; (2) requiring a second medical physicist to review the treatment plan and verify the treatment parameters are correct; (3) requiring the medical physicist to manually calculate the treatment dose and compare it to the dose developed in the treatment plan and written directive prior to administering the treatment to the patient*; and (4) disseminating the results of the investigation, including root cause and corrective actions, to all members of the medical physics staff and authorized user physicians involved in HDR treatments.”
Example 4: NRC Significant Enforcement Action EA-99-257 Although there were no adverse effects as a result of the incident described, in the following is an example of failure to verify the vendor-supplied source strength after HDR source exchange prior to patient treatment. [This action] involved the treatment of patients with your High Dose Rate (HDR) afterloader after the source was replaced, without first checking the source to ensure that the source strength was consistent with the value provided by the manufacturer. Although some quality assurance checks were performed on the HDR unit prior to use, your medical physicist was unable to verify the manufacturer’s source strength certification because the dosimetry system used to verify (calibrate) the source unit was not available. This was a violation of your facility’s procedures that require a check of the source strength after each installation. Although there were no actual consequences in this case, since the source strength was subsequently checked and was consistent with the manufacturer’s values, the failure to check the source prior to use on patients could have resulted in administration of an incorrect dose to the patient, possibly much higher than the prescribed dose for treatment. Credit for corrective actions is warranted because your corrective actions were considered prompt and comprehensive, once the violation was identified by the NRC. These actions include, but are not limited to: (1) a complete review of the HDR program by the Medical Physicist and the Radiation Oncology Manager; (2) a mandatory inservice for all staff associated with the HDR program on the requirements of the license; (3) the development of a “job specifics checklist” for newly hired medical physicists that includes instruction on all requirements of the NRC license; and (4) revision of HDR quality assurance forms highlighting items that must be completed prior to treating patients with the HDR unit.
Example 5: NRC Significant Enforcement Action EA-97-284 In this Enforcement Notice, the licensee was cited for administering radiotherapy treatments that were not explicitly allowed by their RML.
*
Note that item (3) was corrective action adopted by the particular medical center as a result of the misadventure. Manual calculation of HDR treatment doses is recommended by TG-59. However, there is significantly more reliance on computer-based treatment planning systems now than in 1998, when TG-59 was published. Thus, it is more likely for a physicist to carefully review the input parameters to the treatment plan, as opposed to manually calculating the dose at a select number of reference points.
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The violation involves the use of iridium-192 in a HDR afterloader for surface treatment of skin cancer without your NRC license authorizing such use. Your license states that the HDR unit may be used for the treatment of humans only for interstitial, intracavitary, or bronchial therapy. The license does not authorize the use of the HDR for topical skin treatments.
While this example is not useful for most day-to-day application of HDR brachytherapy, it is important for the medical physicist to be aware of the stipulations of their radioactive materials license. This applies to both LDR and HDR brachytherapy and the range of radionuclides and treatment sites/methods that may be employed at a given facility. Example 6: NRC Significant Enforcement Action EA-97-245 As in Example 5 above, this licensee did not ensure that patient treatments were conducted in accordance with their RML. In this instance, the medical physicist performing tasks associated with HDR brachytherapy treatment was not listed on the license, and the medical physicist did not follow protocols outlined in the RML. The violation involves the failure to follow your Quality Management Program (QMP) for your High Dose Rate Afterloading (HDR) treatment program in that certain activities required to be performed by your authorized Medical Physicist listed on the license were, at times, being performed by an individual other than the Medical Physicist listed on the NRC license. Specifically, from January 1997 to May 1997, for several HDR patients’ treatments, the calculations of the required treatment times and verification of the accuracy of the treatment input parameters were performed by an individual other than the Medical Physicist listed on the license. In addition to the violation, one of the requirements of NRC Bulletin 93-01, issued in 1993 following a serious HDR event at a facility in Pennsylvania, is that an authorized user and the RSO or the Medical Physicist must be physically present during each HDR treatment. There is only one individual who is named both the RSO and also the Medical Physicist on your NRC license. During the inspection, you stated that since January 1995, you used physicists other than the Medical Physicist named on the license, to be present during HDR treatments to meet the NRC Bulletin. Also, you indicated that since January 1997, some of the treatment plans were performed by an individual who is not authorized by the NRC license as a medical physicist. You further indicated that at times, the individuals, while in the general vicinity of the treatment areas, were not physically present for the treatment. At the conference, you indicated that since the inspection, the Medical Physicist named on the license has been present for all HDR treatments.
References Glasgow, G. P., J. D. Bourland, P. W. Grisby, J. A. Meli, and K. A. Weaver. “Remote Afterloading Technology.” AAPM Report No. 41. New York: American Institute of Physics, 1993. International Electrotechnical Commission (IEC), Geneva, Switzerland, http://www.iecee.org/cbscheme/ Standard/med.htm last accessed May 1, 2005. IEC 60601-2-17 http://www.astro.org/publications/astronews/2005/Jan/IECReport.htm last accessed May 1, 2005. Kubo, H. D., G. P. Glasgow, T. D. Pethel, B. R. Thomadsen, and J. F. Williamson. (1998). “High dose-rate brachytherapy treatment delivery: Report of the AAPM Radiation Therapy Committee Task Group No. 59.” Med Phys 25(4):375–403. Also available as AAPM Report No. 61. Lerma, F. A., and J. F. Williamson. (2002). “Accurate localization of intracavitary brachytherapy applicators from 3D CT imaging studies.” Med Phys 28:325–333.
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Nath, R., L. L. Anderson, J. A. Meli, A. J. Olch, J. A. Stitt, and J. F. Williamson. (1997). “Code of practice for brachytherapy physics: Report of the AAPM Radiation Therapy Committee Task Group No. 56.” Med Phys 24(10):1557–1598. Also available as AAPM Report No. 59. Rivard, M. J., D-A. R. Evans, and I. A. Kay. (2005). “A technical evaluation of the Nucletron FIRST® system: Conformance of a remote afterloading brachytherapy seed implantation system to manufacturer specifications and AAPM Task Group report recommendations.” J Appl Clin Med Phys 6:22–50. UL 60601 http://www.devicelink.com/mddi/archive/04/02/001.html last accessed May 1, 2005. U.S. Nuclear Regulatory Commission Enforcement Action 99-257: “Clara Maass Medical Center – Notice of Violation.” Washington, D.C., 2000. http://www.nrc.gov/reading-rm/doc-collections/enforcement/actions/materials/ ea99257.html. U.S. Nuclear Regulatory Commission Enforcement Action 04-093: “Missouri Baptist Medical Center – Notice of Violation.” Washington, D.C., 2004. http://www.nrc.gov/reading-rm/doc-collections/enforcement/actions/materials/ea04093.html. U.S. Nuclear Regulatory Commission Informational Notice 00-05: “Recent medical misadministrations resulting from inattention to detail.” Office of Nuclear Material Safety and Safeguards, Washington, D.C., 2000. http://www.nrc.gov/reading-rm/doc-collections/enforcement/actions/materials/ea99257.html. U.S. Nuclear Regulatory Commission Enforcement Action 97-284: “Centre Community Hospital – Notice of Violation.” Washington, D.C., 1997. http://www.nrc.gov/reading-rm/doc-collections/enforcement/actions/materials/ ea97284.html. U.S. Nuclear Regulatory Commission Enforcement Action 04-087: “St. Vincent Hospital & Health Care Center – Notice of Violation.” Washington, D.C., 2004. http://www.nrc.gov/reading-rm/doc-collections/enforcement/ actions/materials/ea04087.html. U.S. Nuclear Regulatory Commission Enforcement Action 97-245: “Mountainside Hospital – Notice of Violation.” Washington, D.C., 1997. http://www.nrc.gov/reading-rm/doc-collections/enforcement/actions/materials/ ea97245.html.
Chapter 8
Continuous Low Dose Rate and Pulsed Dose Rate Remote Afterloader Units John L. Horton, Ph.D., Ann Lawyer, M.S., and Firas Mourtada, Ph.D. University of Texas M. D. Anderson Cancer Center Radiation Physics Department Houston, Texas Continuous Low Dose Rate Remote Afterloader Units (LDR RAUs) . . . . . . . . . . . . . . . . . . . . . 99 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 99 Selectron LDR RAU . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 100 Advantages of LDR RAUs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 101 Disadvantages of LDR RAUs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 102 Facility Considerations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 103 Licensing Issues . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 103 License Application . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 103 Record Keeping . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 104 Training . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 104 Nucletron Pulsed Dose Rate Remote Afterloader Unit (PDR RAU) . . . . . . . . . . . . . . . . . . . . . 104 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 104 Advantages of PDR RAUs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 104 Disadvantages of the PDR RAUs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 105 Facility Considerations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 106 Licensing Issues . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 106 Training . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 107 Commissioning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 107 Quality Control . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 107 Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 108 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 108
Continuous Low Dose Rate Remote Afterloader Units (LDR RAUs) Introduction University of Texas M. D. Anderson Cancer Center (UTMDACC) has a long history of continuous low dose-rate (LDR) brachytherapy treatment for gynecological cancers. The Fletcher Suit Delclos (FSD) applicators and their variants are probably the most widely used applicators in the United States. Approximately 18 years ago UTMDACC adopted clinical use of the Selectron low dose-rate remote afterloader unit (LDR RAU), shown in Figure 1, as a replacement for manually afterloaded 226Ra treatments. The choice of the LDR RAU was predicated on the desire to maintain LDR treatments. Currently, approximately 90% of our gynecological tandem and ovoid and tandem and cylinder treatments are performed with the Selectron LDR RAU. The remaining 10% are performed with manually loaded 137Cs with an occasional patient treated with the Nucletron high dose-rate remote afterloader (HDR RAU). Treatments average about one patient per month with the HDR RAU, compared with about three per week with LDR treatments. The HDR RAU patients are typically those with medical problems that prevent them from spending 48 hours in a hospital bed. Our principal use of HDR RAU for gynecological patients is limited to those treated for vaginal cuff disease with dome and cylinders. Treatments average about three patients per week with this modality.
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Figure 1. Selectron Low Dose Rate Remote Afterloader Unit.
Selectron LDR RAU The Selectron LDR RAU was chosen as a replacement for 226Ra because an “equivalent” dose rate and dose distribution could be achieved, allowing us to build on several decades of prior clinical experience. The scattering and attenuation characteristics of 137Cs gamma rays in tissue are similar to 226Ra gamma rays, providing very nearly the same relative dose distribution. By proper choice of activity of the 137Cs pellets, the 226Ra dose rate could be approximated. The UTMDACC Selectron LDR RAUs have four channels, one for the tandem, one for each ovoid, and one for backup. Combinations of 48 active and inactive pellets are placed in each channel to simulate manually loaded intracavitary brachytherapy systems. The active pellets have a nominal activity of 5 mgRaeq (36.135 U) each. All pellets are 2.5 mm in diameter. The inactive pellets are ferromagnetic and the active pellets are nonmagnetic. The RAU is controlled by a microprocessor. The positions of the active pellets in each channel are programmed into the microprocessor. The pellets are transferred from the main safe to an intermediate safe. The pellets are sorted by a magnetic sorting technique. The pellets are stored in their programmed order in the intermediate safe. When treatment is initiated, the pellets are transferred from the intermediate safe in the afterloader to the brachytherapy applicators. When treatment is interrupted for nursing care or visitors, the pellets are transferred back to the intermediate safe. At the end of treatment, the pellets are transferred to the main safe. All pellet transfers are performed with compressed air. Prior to clinical implementation of the Selectron LDR RAU, extensive treatment planning comparisons were performed between historical 226Ra treatments and treatments to be performed with the Selectron LDR RAU. To ease the transition, a table was developed that translated 226Ra treatments to positions of active sources for the Selectron LDR RAU treatments.
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Advantages of LDR RAUs The principal advantage of all RAUs is the reduction of personnel exposure to all staff members, including the radiation oncologist, physicist, dosimetrist and nursing staff, because the sources remain in the shielded radiation safe when personnel are in the room caring for the patient. In the European Union nations with stricter limits on personnel exposure, this is more of an issue than in the United States. Usually, in the United States, the reduced personnel exposure is not a sufficient advantage to offset the increased costs in the minds of most hospital administrators. Manually afterloaded brachytherapy implants can be performed while adhering to personnel exposure limitations by following a well conceived and executed plan limiting the dose to radiation workers. A second advantage of the LDR RAU is the reduced dose to the rectum. This reduced dose results from the LDR RAU ovoids being 5 mm longer than the ovoids that were used for 226Ra treatments and from the source activity loaded within the ovoid being in a more anterior position. The standard LDR RAU ovoids in use at UTMDACC are 33 mm long rather than the 28 mm of the manually loaded FSD ovoid. Also, the activity within the 226Ra sources and within the 3M 137Cs sources is asymmetrically placed within the encapsulation, because of the eyelet in these sources. This arrangement places the activity to the posterior pole of the ovoid rather than being centered in the ovoid. The newer J sources from Amersham center the activity in the encapsulation, but the means of orienting the sources within the ovoid places the center of the activity posterior to the center of the ovoid. With the Selectron RAU, at UTMDACC the ovoid activity is centered within the ovoid. A third advantage is the higher degree of “optimization” of the dose distribution that can be achieved with the LDR RAU because of the flexibility of placing the sources within the applicator. During the time of 226Ra treatments, activity loadings were specified in terms of “sources” or “inches,” depending on the length of the tandem relative to its position with the external os and the ovoids. The 226Ra sources were 22 mm long. If the tandem sources were abutted, the loading was prescribed as a “sources” loading. However, if the physician desired to spread the sources out over a greater length, 3-mm spacers were used between the 226Ra sources; this prescription was considered an “inches” loading. As the change was made to manually afterloaded 137Cs sources, this schema continued. However, as the 137Cs sources were only 20 mm in length, rather than the 22 mm of the 226Ra sources, 2 mm-spacers were used for “sources” loading and 5-mm spacers for “inches” loading. “Short sources” loading was also used for 137Cs. For the short sources prescriptions for 137Cs, the sources were abutted. This nomenclature continues today at UTMDACC with the LDR RAU treatments, and is detailed in two sets of standard loading tables. Table 1 was developed for “sources” loading in the tandem, with the numbers referring to the position of the active pellets. Table 2 is similar except it is for “inches” loading in the tandem. Other tables are available for different mg Ra eq loadings in the ovoids. The “custom” loadings for tandem and ovoid patients usually affect only the more inferior sources in the tandem. This customization gives the radiation oncologist more flexibility in dealing with relative geometry between the tip of the tandem and the flange. In recent years the degree of custom loading of UTMDACC cases has increased over the examples seen in Tables 1 and 2. There are tables for “short sources” and tables for tandems with various length spacers at the tip. UTMDACC dosimetrists frequently customize the treatments for patients beyond these loadings, which the many tables address. The decisions on these custom loadings may be made on the physician’s experience, based on the extent of the disease, geometry of the implant, and external beam dose, or the physician may want to see a treatment plan before beginning treatment. Tables 1 and 2 are given here only for illustration and should not be used in other clinics without first verifying that they are appropriate for that practice. More importantly, the flexibility in source positioning provided by the LDR RAU is a great asset for tandem and cylinder treatments. The high degree of customization with the LDR RAU allows the
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John L. Horton et al. Table 1. “Sources” Loading Numbers refer to active source positions.
Superior source
Inferior source
5 mgRaeq
3
11
20
29
38
10 mgRaeq
1,3
10,12
19,21
29,31
37,39
15 mgRaeq
1,3,5
10,12,14
19,21,23
28,30,32
36,38,40
20 mgRaeq
1,2,3,4
10,11,12,13
19,20,21,22
28,29,30,31
37,38,39,40
Table 2. “Inches” Loading Numbers refer to active source positions.
Superior source
Inferior source
5 mgRaeq
3
13
23
33
44
10 mgRaeq
1,3
12,14
22,24
32,34
43,45
15 mgRaeq
1,3,5
11,13,15
21,23,25
31,33,35
42,44,46
20 mgRaeq
1,2,3,4
11,12,13,14
22,23,24,25
32,33,34,35
42,43,44,45
dosimetrist to design a plan that fulfills the prescription of the radiation oncologist much more easily than for the case with manually afterloaded patients. As can be seen from this discussion, intracavitary brachytherapy treatments as practiced at UTMDACC are almost as much art as science with the patient treatment based on semi-empirical rules articulated by Fletcher (Fletcher 1980) and Katz and Eifel (Katz and Eifel 2000).
Disadvantages of LDR RAUs The principal disadvantage of the LDR RAU is the possibility of failure during treatment. The great majority of these failures occur when the pellets are being transported from the radiation safe to the patient. Failures typically occur because of low air pressure, a power failure because of weather conditions, or a crimped transfer tube. The nursing staff can handle the majority of these failures, simply by restarting the transit. Occasionally a dosimetrist or medical physicist may be required to reprogram a unit. Maintenance may be indicated when the failures become too frequent. Maintenance issues typically arise when parts in the afterloader that are involved in the transfer of the pellets become worn and “out-of-round.” The pellets also can become out-of-round, leading to failures in their transport. To minimize the problems with pellets, Nucletron limits the number of pellet transfers to 7000. The second disadvantage of the LDR RAU is the desire of the patient to have visitors during treatment. Normally one would think that this would be an advantage of the LDR RAU. For patients being treated with manually loaded sources, visitors are limited to 30 minutes a day to prevent exceeding the visitor’s regulatory dose/exposure limits, while the patient treated with the LDR RAU can receive visitors for extended periods of time, since the visitor will not be irradiated. The sources will be within the radiation safe while a visitor is present. However, when the sources are within the safe, the patient is not being treated and the treatment time is extended. The physician counsels patients about the issue of extending the treatment time by having visitors. Given the choice, patients tend to limit their visitors in order to leave the hospital in a timely manner.
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Facility Considerations Planning for the facility to house an LDR RAU involves shielding, location, ancillary services, and audiovisual communication issues. To design appropriate shielding, one must decide what treatments will occur in the room and what will be the maximum air kerma strength per hour and maximum total reference air kerma (TRAK) per week for these treatments. The room shielding must limit the dose equivalent in uncontrolled areas to less than 1.0 mSv per year, with a maximum dose equivalent of 0.02 mSv in any one hour (NCRP 1987). Radiation workers may receive up to 0.05 Sv per year, but based on ALARA (as low as reasonably achievable), rooms should generally be designed to limit exposure of radiation workers to 0.1 mSv per week in controlled areas (NCRP 1987). Some states also limit the instantaneous dose equivalent rate received for both the general public and radiation workers. The instantaneous dose equivalent rate limit is usually not an issue for LDR RAU. As always, a facility and an individual must comply with the applicable regulations in their state. Although a bedside shield may be used to reduce the dose equivalent levels to comply with the regulations, adequately shielded walls are the preferable alternative. This situation does not rely on someone to place the shield in the appropriate position for each case and provides more usable floor space in the room. LDR RAU rooms are typically hospital rooms that should be located close to the nurses’ station where the nurse can keep the patient and the room under surveillance. If there are multiple shielded rooms, they should be adjacent to each other and remote from unshielded hospital rooms. The nurses need an intercom and a closed circuit television system to observe and to communicate with the patient during treatment. The LDR RAU treatment control console is placed immediately outside the patient room, with a remote at the nurses’ station. The power for the LDR RAU should be on the emergency power circuit. Compressed air is also required for the Selectron LDR RAU. The room door should be interlocked. An area radiation monitor should be in the room with a remote outside the room to indicate when the sources are exposed. As with any brachytherapy room, there should be emergency equipment and an emergency container (bail out pig) in the room. The door should have a plaque for mounting radiation signs and instructions when patients are being treated.
Licensing Issues RAUs and the sources they contain must be licensed by the U.S. Nuclear Regulatory Commission in nonAgreement States and federal facilities or the appropriate state or local agency in Agreement States. Before a source or device can be licensed, it must be on the Registry of Sealed Sources and Devices. Typically, the process of entering a source or device on this registry requires many months and significant paperwork from the vendors. The vendor should always supply a copy of its registration before you consider the purchase of the source or device. License Application Whatever you state you will do in your license, you must perform. You can always increase your tasks from what is stated in your license, but you can never decrease what you indicate in your license without a time-consuming license amendment and thorough justification. Always promise less than you plan to deliver in these situations. Place only the minimum required information in your license application. The license application requires a description of the source and its registry number, the manufacturer and model number of the afterloader, the authorized users and their qualifications, the planned use of the device, the location of this planned use, and suitability of the room shielding. You should also include the radiation detection devices to be used, the radiation warning devices, the audiovisual equipment to view and communicate with the patient, and a description of the security of the area and for the sources. Details of your quality control program must be discussed, including calibration techniques and frequency, routine
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quality control procedures and frequency, and leak test procedures and frequency. The qualifications of the individuals who will perform these tests and procedures, training of operators, emergency procedures and where they are posted, and disposal of decayed sources are required items. You also need to specify the maximum activity of the specified isotope that will be at your facility at any one time. Remember that when you have a source exchange, you have not only the new source but also the old source in your facility. It is usually acceptable to specify twice the activity of the new source to account for this. Of course, if you have other sources of the specified isotope in addition to what is in the afterloader, you also must ensure that these sources are accounted for in your specification of the maximum activity. Record Keeping Finally, the regulatory agency has requirements for record keeping of calibrations, quality control tests, room surveys, and treatment head surveys. These regulations include the length of time these records must be maintained. Be sure you adhere to all these requirements.
Training In addition to the standard radiation protection training, personnel require training specific for the RAU. Medical physicists and dosimetrists should be trained in the operation, programming, and emergency response procedures for the LDR RAU. The nursing staff needs to know how to start and stop the unit and be familiar with emergency response procedures. Physicians also need to be familiar with emergency response procedures.
Nucletron Pulsed Dose Rate Remote Afterloader Unit Introduction The Nucletron Pulsed Dose Rate Remote Afterloader Unit (PDR RAU), shown in Figure 2, is based on the Nucletron HDR afterloader described in a previous chapter. The physical construction of the PDR source is identical to the Nucletron HDR source (mHDR v2). The only difference is the lower activity of the PDR source (between 37 and 74 GBq at installation, depending on the institution’s needs and room shielding) compared to the HDR source (approximately 370 GBq at installation). The PDR RAU simulates continuous LDR treatments by delivering the same total dose in the same total time. With the higher activity source, this is accomplished by exposing the source from the afterloader for only a fraction of the time for each hour. Consider this example: the desired dose rate is 50 cGy/hr at the point of interest for a total treatment time of 48 hours. The PDR source delivers an instantaneous dose rate of 500 cGy/hr at that point. In this instance, the source should be exposed for a tenth of an hour, or 6 minutes, each hour for the same total treatment time of 48 hours. This schema delivers the same total dose in the same total time yielding the same average dose rate.
Advantages of PDR RAUs An essential component of the UTMDACC brachytherapy program is to continue the LDR program or a modality that provides a radiobiologically equivalent treatment. Radiobiological models and measurements indicate that PDR provides this capability (Brenner and Hall 1991; Fowler and Mount 1992; Millar, Hendry, and Canney 1996; Brenner et al. 1996; Armour et al. 1997; Sminia et al. 1998). The PDR RAU provides the same personnel protection as the LDR RAU, because no personnel are in the room when the sources are loaded into the applicators. However, the PDR provides the additional advantage in that nursing care can occur during the time between treatment pulses. For all patients, but
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Figure 2. Nucletron Pulsed Dose Rate Remote Afterloader Unit.
especially for patients who require extra nursing care, this means that the extra care can be given without extending the overall treatment time. This also permits the patient to have visitors without extending the treatment time. These two advantages should reduce the stress of the treatment for the patient. Another advantage of the PDR RAU is the ability to achieve greater conformation of the dose to the treatment volume than is possible with the LDR RAU. This advantage results from being able to perform computer optimization of individual dwell times. The PDR RAU provides the ability to vary the dwell times at each source position during each treatment pulse. As the LDR RAU sources have a nominal activity of 5 mgRaeq, two sources are required to simulate a 10 mgRaeq source and three sources for a 15-mgRaeq source, etc. The LDR RAU provides limited flexibility in plan optimization in that it may simulate a 12.5-mgRaeq source by programming the unit to deliver half the treatment with three sources and half with two sources. However, to accomplish this requires a dosimetrist or medical physicist to reprogram the unit halfway through the treatment. In reality, for true optimization, a critical aspect of the process is the ability to image the implant with computed tomography or magnetic resonance. Analogous to external beam treatment, the physician must be able to visualize and to contour the treatment volume and all critical structures to make meaningful decisions about the optimization process.
Disadvantages of PDR RAUs The most serious disadvantage of the PDR RAU is the potential for afterloader failures with the source in the exposed position. A very large dose could result if the high activity source were to stick at one dwell position during treatment. As a first order approximation, consider that a typical tandem and ovoid treatment may consist of 15 dwell positions, which are treated for a total time of 10 minutes each hour. This implies the average dwell time at each position is approximately 10 minutes divided by the 15 positions, or 40 seconds. Clearly, if the source remains at one dwell position, an excessive dose will be
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delivered in a very short time. If it takes 10 minutes to respond to the situation and retract the source, the overdose at that dwell position will be a factor of 15 too high. This situation requires an immediate response by a well-trained staff. Since the patient is in a hospital room for up to 72 hours, the medical physicist cannot be on site during the entire treatment. In a large hospital he/she may be remote from the treatment area and it may take 10 minutes to respond even if the medical physicist is on site. Preparation for the worst-case scenario requires that the nursing staff be able to respond to the emergency situation in a very timely manner. Fortunately, the probability of the source not being retracted is very low, based on data from the Academic Medical Center-Amsterdam (Blank 2004). They have reported their experience with PDR RAU for 417 treatments consisting of 5117 pulses. Of the 5117 planned pulses, 272 pulses (5.3%) encountered an obstruction during the check cable run. There were only five cases (0.1%) of an obstruction during an active sources run, all of which occurred when the active sources was exiting the PDR RAU. In these cases the machine retracted the active source without incident or the need for human intervention. Of the 417 treatments analyzed, five cases (1.2%) required an adjustment of the implant as a result of the catheters being kinked. Another five cases had their treatment discontinued for medical reasons. The Academic Medical Center-Amsterdam concluded based on their experience that 1. Errors can occur with the PDR RAU and on rare occasion require a treatment adjustment. 2. Errors do not always occur during the first pulse; hence an error-free start does not assure an errorfree treatment. 3. The check cable run has proven an effective system to prevent accidents with the active source. Academic Medical Center-Amsterdam has two physicists and four dosimetrists who are on call to address issues during the PDR treatments. They state that 95% of the issues can be solved by telephone conversation with the nurse.
Facility Considerations The facility requirements for a PDR RAU will be the same as for the Selectron LDR RAU, except for the compressed air. If the room has been shielded for LDR treatments, the shielding should be adequate for the equivalent PDR treatments. The instantaneous dose equivalent rate will be higher with the PDR RAU than for LDR treatments, but the dose equivalent rate averaged over an hour should be the same as for the LDR treatment. However, if local regulations include an instantaneous dose rate limit, it must be verified that these limits will not be violated. From a shielding standpoint, one advantage PDR has over LDR is that the half-value layer (HVL) in concrete for 192Ir is less than the concrete HVL for 137Cs (NCRP 1972), because of the lower average energy of the 192Ir. One caveat though is that if the new PDR treatments are planned that were not performed previously with LDR, these treatments may result in a higher average dose equivalent rate or higher total dose equivalent. For instance, large interstitial implants may result in higher average dose equivalent rate and higher total dose equivalent than the gynecological treatments. In this case, the adequacy of the shields must be verified.
Licensing Issues The issues involved with licensing this unit are essentially the same as discussed above for the LDR RAU and are common to all RAUs.
8 – Continuous Low Dose Rate and Pulsed Dose Rate Remote Afterloader Units
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Training As with the LDR RAU, medical physicists and dosimetrists should be trained in the operation, programming, and emergency response procedures for the PDR RAU. The nursing staff needs to know how to start and stop the unit and be very familiar with the emergency response procedures. Physicians also need to be familiar with emergency response procedures. Because of the much higher activity of the source, everyone’s knowledge of emergency response procedures is much more important for operating the PDR RAU than for the LDR RAU.
Commissioning Because the Nucletron PDR RAU is basically the same machine with different software as the Nucletron HDR RAU, commissioning of the PDR RAU is similar to the HDR RAU commissioning previously discussed in other chapters. However, for the PDR RAU the determination of the transit dose from the source is important because of the large number of source transits during a treatment. Consider a large interstitial implant of 18 catheters treated for 5 days. As the source will be exposed once each hour, the number of transits is 120 transits per catheter × 18 catheters = 2160 total transits. The treatment control station software accounts for much of the transit time between dwell positions by subtracting at least a 0.1 s from each dwell time. Most of the transit dose to critical structures near the implant results from the return of the source at the completion of a treatment session. Transit dose may also be delivered to normal tissue external to the implant as the source enters and exits the treatment volume. The transit dose may be estimated by simulating a large implant with a multicatheter applicator and measuring the transit dose with either high-speed radiographic film or another sensitive dosimeter, such as, a MOSFET (metal oxide semiconductor-field effect transistor) or diode. Another important aspect of a PDR unit is the ability to transfer the source to and retract the source from the applicator multiple times without failure. Nucletron warrants the PDR unit to perform 25,000 transfers without failure. Tests should also be performed with varying angles of flexion of the transfer tubes to determine the angle of flexion that will cause a failure. At UTMDACC the foremost consideration for gynecological treatments, after verifying machine reliability, establishing adequacy of the room shielding, developing procedures for emergency response, and training all personnel in emergency response and machine operation, is the difference in the gynecological applicator designs. Comparison of Monte Carlo calculations and radiochromic film measurements in phantom around the PDR RAU Fletcher Williamson applicators to similar calculations and measurements for the FSD applicators used with the LDR RAU is currently underway (Gifford et al. 2005). Also being compared is a selected group of patients for whom there are CT scans of their FSD treatments with the 137Cs LDR RAU available. Gifford (2004) performed a Monte Carlo analysis including the effects of the shields on the doses these patients received. These results will be compared to Monte Carlo calculations assuming the same source positions in the applicators, but for 192Ir and the PDR applicators. A research program to study other applicators and perhaps to develop new applicators that can provide an improved dose distribution with 192Ir has begun.
Quality Control Because the PDR RAU is basically the same as the HDR RAU discussed in an earlier chapter, the quality control should address the same issues of source position accuracy, room and head radiation surveys, interlock checks, audiovisual device checks, and source calibrations. The difference with the PDR RAU is the importance of the nursing staff training, as discussed above. Because the PDR RAU provides an order of magnitude lower dose rate than the HDR RAU, medical physicists are not required to be present
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throughout the treatment. However, because the PDR RAU has an order of magnitude higher dose rate than the LDR RAU, the nursing staff must be able to respond to an emergency situation in a very timely manner. The nursing staff must be the first line of defense. This means that the nursing staff must be well trained before beginning a PDR RAU program and frequent in-services to reinforce this training are a must. A well-planned and documented program of appropriate actions based on the errors codes from the computer is also essential.
Conclusions The PDR RAU offers several advantages over either manual afterloading or remote LDR afterloading. The PDR RAU provides the same average dose rate and radiobiologically equivalent treatment as continuous LDR treatment, allowing us to build on our many years of experience with continuous LDR treatments. PDR RAU provides personnel protection for both hospital staff and visitors. The nursing care and visiting time can occur between treatment pulses without extending the total treatment time. The PDR RAU provides the means to more tightly conform the radiation dose to the treatment volume than the manually afterloaded or remotely afterloaded LDR implants. However, a careful comparison of the applicators used and dose distributions delivered via the LDR RAU with the PDR RAU is essential before changing from a well-established technique. Finally, the nursing staff must be well trained and accept their new responsibilities before embarking on a PDR program.
References Armour, E. P., J. R. White, A. Armin, P. M. Corry, M. Coffey, C. DeWitt, and A. Martinez. (1997). “Pulsed low dose rate brachytherapy in a rat model: Dependence of late rectal injury on radiation pulse size.” Int J Radiat Oncol Biol Phys 38:825–834. Blank, L. Overview of LDR-PDR Differentiators. Presented at Nucletron Workshop: From Traditional to Modern Brachytherapy. Transferring from LDR to PDR. Amsterdam, The Netherlands, 2004. Brenner, D. J., and E. J. Hall. (1991). “Conditions for the equivalence of continuous to pulsed low dose rate brachytherapy.” Int J Radiat Oncol Biol Phys 20:181–190. Brenner, D. J., E. J. Hall, G. Randers-Pehrson, Y. Huang, G. W. Johnson, R. W. Miller, B. Wu, M. E. Vazquez, C. Medvedosky, and B. V. Worgul. (1996). “Quantitative comparisons of continuous and pulsed low dose rate regimens in a model late-effect system.” Int J Radiat Oncol Biol Phys 34:905–910. Fletcher, G. H. Textbook of Radiotherapy. Third Edition. Philadelphia, PA: Lea & Febiger, 1980. Fowler, J., and M. Mount. (1992). “Pulsed brachytherapy: The conditions for no significant loss of therapeutic ratio compared with traditional low dose rate brachytherapy.” Int J Radiat Oncol Biol Phys 23:661–669. Gifford, K. A 3-D CT Assisted Monte Carlo Evaluation of Intracavitary Brachytherapy Implants. Ph.D. dissertation. The University of Texas Graduate School of Biomedical Sciences. Houston, TX, 2004. Gifford, K., J. Horton, E. Jackson, T. Steger, M. Heard, F. Mourtada, A. Lawyer, and G. Ibbott. (2005).”Verification of Monte Carlo calculations around a Fletcher Suit Delclos ovoid with radiochromic film and normoxic polymer gel dosimetry.” Med Phys (Accepted for publication). Katz, A., and P. J. Eifel. (2000). “Quantification of intracavitary brachytherapy parameters and correlation with outcome in patients with carcinoma of the cervix.” Int J Radiat Oncol Biol Phys 48:1417–1425. Millar, W. T., J. H. Hendry, and P. A. Canney. (1996). “The influence of the number of fractions and biexponential repair kinetics on biological equivalence in pulsed brachytherapy.” Brit J Radiol 69:457–468. National Council on Radiation Protection and Measurements (NCRP). NCRP Report No. 40. Protection Against Radiation from Brachytherapy Sources. Bethesda, MD: NCRP, 1972. National Council on Radiation Protection and Measurements (NCRP). NCRP Report No. 91. Recommendations on Limits for Exposure to Ionizing Radiation. Bethesda, MD: NCRP, 1987. Sminia, P., C. J. Schneider, K. Koedooder, G. van Tienhoven, L. E. Blank, and D. G. Gonzalez. (1998). “Pulse frequency in pulsed brachytherapy based on tissue repair kinetics.” Int J Radiat Oncol Biol Phys 41:139–150.
Chapter 9
An Aperçu of Codes, Directives, Guidances, Notices, and Regulations in Brachytherapy Glenn P. Glasgow, M.S., Ph.D., F.A.A.P.M., F.A.C.R. Department of Radiation Oncology Loyola University Chicago Stritch School of Medicine Maywood, Illinois Brachytherapy Regulatory Overview (What’s New Since 1994 Summer School?) . . . . . . . . . 109 Is the regulatory climate hot (for more regulations) or cold (quiescent)? . . . . . . . . . . . . . . . . . . . . . 109 There is no excuse not to know the regulations: They are on the web! . . . . . . . . . . . . . . . . . . . . . . 110 What’s hot and what’s not in brachytherapy procedures? . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 110 Do you know where your sources are? Who else knows? . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 111 You still haven’t read the “new” (4-24-02) USNRC 10 CFR 20, 32, and 35 regulations? . . . . . . . . 111 I’m an agreement state licensee! Why should I be concerned about changes in federal codes? . . . 112 Brachytherapy Regulatory Overview (What’s Unchanged Since 1994 Summer School?) . . . 112 10 CFR 19 (Notices, Instructions, and Reports to Workers; Inspections) . . . . . . . . . . . . . . . . . . . . 112 10 CFR 20 (Standards for Protection Against Radiation) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 112 Regulatory Review: What’s Changed in 10 CFR 20 (Standards…Radiation)? . . . . . . . . . . . . . 112 10 CFR 20.1002/Scope; –.1003/Definitions, and –.1301/Dose Limits for Individual Members of the Public (USNRC 2002) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 112 Regulatory Review: What’s Changed in 10 CFR 32 (Specific…Material)? . . . . . . . . . . . . . . . 113 Regulatory Review: What’s “New” in 126 Sections of 10 CFR 35 (Medical Use…Material)? . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 113 Components of 10 CFR 35 Applicable to All Forms of Brachytherapy (Tables 4A,B) . . . . . . . . . . 114 Some Components of 10 CFR 35 (F) Applicable to Manual Brachytherapy (Table 4C) . . . . . . . . . 115 Some Components of 10 CFR 35 (H) for Photon-emitting Remote Afterloaders (Tables 4D,E) . . . 115 Some Components of 10 CFR 35 (L) (Record Retentions) (Table 5) . . . . . . . . . . . . . . . . . . . . . . . . 120 Some Components of 10 CFR 35 (M) (Reports…Medical Events…Sources) (Table 6) . . . . . . . . . 120 Bulletins, Directives, Guidances, Information Notices, Newsletters, and Regulatory Summaries, for Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 121 Bulletins . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 121 Directives . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 121 Guidances . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 122 Information Notices . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 122 Newsletters . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 122 Regulatory Summaries . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 122 Recent NRC Activities—Specialty Boards and Training Programs . . . . . . . . . . . . . . . . . . . . . . . . . 122 Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 124 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 125
Brachytherapy Regulatory Overview (What’s New Since 1994 Summer School?) Is the regulatory climate hot (for more regulations) or cold (quiescent)? Regulations are like taxes: Everyone thinks they should be reduced! However, it rarely occurs. And, like taxes, when the “Code” is revised, taxes or regulations that appear to have been eliminated in one section often appear in other sections, often at higher rates! Occasionally, a lonely voice cries out for reform, but usually the cry goes unanswered. Mossman proposes that regulatory effectiveness and efficiency would be improved by three changes: (1) adoption of a dose-based rather than the current risk-based system;
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(2) adoption of the International System of Units; and (3) establishing a single, independent office to coordinate nuclear regulations established by U.S. federal agencies and departments (Mossman 2003). However, the pace of regulatory change is glacial—usually advancing, not retracting! So, let us review the significant changes in the past 10 years in brachytherapy and the corresponding changes in regulations. The significant changes were: (1) Increased use of the World Wide Web Internet to disseminate regulations; (2) changes in types of brachytherapy procedures; (3) concerns about byproduct material security (or lack thereof) in medical facilities; (4) the adoption of new federal codes, and (5) implementation of these codes by Agreement States. The opinions expressed are exclusively those of the author, based on 30 years of experience working directly, and concurrently, with regulators under both an NRC license (Edward Hines VA Hospital) and Agreement State (Illinois) license (Loyola University Chicago Medical Center).
There is no excuse not to know the regulations: They are on the web! A comprehensive overview (Deye 1988), now dated, still serves as an excellent tutorial for those unfamiliar with organizations, such as the International Atomic Energy Agency (IAEA), International Commission on Radiological Protection (ICRP), and International Commission on Radiation Units and Measurements (ICRU), and codes and regulations. Table 1 presents useful, current web addresses for international organizations that often make recommendations that become the basis of subsequent federal regulations made by federal agencies. The National Council on Radiation Protection and Measurements (NCRP) makes, via statements and publications, recommendations on matters of radiation usage and, subsequently, safety issues. Through the Environmental Protection Agency (EPA), the federal agencies, Departments of Energy (DOE) and Transportation (DOT), Food and Drug Administration (FDA), currently coordinate their regulatory efforts. The U.S. Nuclear Regulatory Commission (NRC) is the lead agency for promulgating regulations related to brachytherapy sources and their use. Their regulations apply in 17 non-Agreement states and in federal facilities directly holding federal licenses. Agreement States, currently 33 in number, agree, under certain terms, to apply those regulations adopted by the NRC. The Conference of Radiation Control Program Directors (CRCPD) coordinates regulatory issues among its member states and proposes model state licensure and regulatory language. Their web site has electronic links to all state regulatory agencies. The Radiation Safety Officers Toolbox, via Idaho State University (http://www.physics.isu.edu/radinf/rsotool box.htm), provides direct links to the Code of Federal Regulations (CFR) discussed later in this chapter. Other organizations, such as the American Association of Physicists in Medicine (AAPM), Health Physics Society (HPS), Society of Nuclear Medicine (SNM), often issue position statements regarding proposed regulations. Higson (2001) offers a useful international view of the regulation of medical devices for public health and safety. Benedetto (1995) reviews how regulations arise in the United States.
What’s hot and what’s not in brachytherapy procedures? Brachytherapy prospers in the United States. Interstitial prostate implants with 125I and 103 Pd seeds are popular and show no signs of decaying! Remote high dose rate (HDR) afterloading procedures are increasing; mobile units serve multiple hospitals in heavily populated urban areas. Traditional single-session procedures (gynecology treatments with manually loaded 137Cs sources, interstitial sarcoma implants, etc.) are being replaced with multiple fractional HDR twice-a-day treatment regimens. Our original (7/88) HDR program at Hines VA Hospital was a Gamma Med IIi located in a 60Co teletherapy vault. Treating lung, esophagus, and the vagina, we performed about 49 procedures yearly. Since relocating (8/03) the program to a new facility at Loyola with a dedicated HDR vault, and opening a full gynecologic service and interstitial sarcoma and head and neck implant service, we perform 200+ cases yearly. Treatment of a new site (breast) with MammoSite® applicator additionally increased the total number of annual procedures.
9–An Aperçu of Codes, Directives, Guidances, Notices, and Regulations in Brachytherapy 111 Table 1. Useful Web Sites with Information about Radiation, Regulations, and Regulatory Issues Agency Conference Radiation Control Program Directors (CRCPD) Department of Energy (DOE) Department of Transportation (DOT) Environmental Protection Agency (EPA) Food and Drug Administration (FDA) Health Physics Society (HPS) International Atomic Energy Agency (IAEA) International Commission on Radiological Protection (ICRP) International Commission Radiation Units and Measurements (ICRU) National Council on Radiation Protection and Measurements (NCRP) Nuclear Regulatory Commission (NRC) Idaho State University
Internet Address (“http://www.) crcpd.org energy.gov dot.gov
Electronic Mail Address Not given on web page Not given on web page
[email protected] epa.gov
Not given on web page
fda.gov
Not given on web page
hps.org iaea.org
[email protected] [email protected] icrp.org
[email protected] icru.org
[email protected] ncrp.com
Not given on web page
nrc.gov
Not given on web page
physics.isu.edu/radinf/rsotoolbox
Not given on web page
Regulatory hurdles to treating with pulsed dose rate (PDR) remote afterloading devices have been lowered. Intravascular brachytherapy (IVB), popular for a short time, unfortunately exhibited a reasonably short half-life and was rapidly abandoned at many facilities.
Do you know where your sources are? Who else knows? A new international and national concern is the security of byproduct sources in medical facilities. Most medical licensees have small (multiple millicurie) quantities of long-lived byproduct materials (137Cs, 60Co, etc.), ideal components for a dispersal “dirty bomb.” The IAEA has developed an action plan to combat nuclear terrorism (Health Physics News and Notices 2002, 2003). These international efforts likely will lead to new national and state regulations requiring greater security for radioactive sources.
You still haven’t read the “new” (4-24-02) USNRC 10 CFR 20, 32, and 35 regulations? Shame on you! You will not be spared! This limited presentation (NB: think of it as “Regulations Lite”) reviews the cogent details of regulations that did not change as well as those that did.
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I’m an agreement state licensee! Why should I be concerned about changes in federal codes? Agreement States have certain periods (5 years or more) within which state regulations must become compliant, at certain levels of compliance, with NRC regulations. A licensee may be in “regulatory transition,” a dangerous period of change! For example, at this time (2005) Illinois is enforcing state regulations based on NRC regulations in force prior to the 2002 regulatory changes, as they prepare new state regulations compliant with the recent changes in federal codes. Other states are in similar regulatory transitions.
Brachytherapy Regulatory Overview (What’s Unchanged Since 1994 Summer School?) 10 CFR 19 (Notices, Instructions, and Reports to Workers; Inspections) This long-standing regulation (USNRC 1981), with 14 sections, issued 12/18/1981, un-revised, remains in force. Table 2 lists seven important sections with brief comments about their content. Staying compliant with “instructions” to a constantly changing workforce is a regulatory challenge!
10 CFR 20 (Standards for Protection Against Radiation) These standards, consisting of 69 sections, are (with one exception discussed later) mostly unchanged from the 5/21/1991 release (USNRC 1991). Tables 3A and 3B list 10 key headings with brief comments about their content. Note that security of radioactive materials (RAM) is addressed in §20.1801.
Regulatory Review: What’s Changed in 10 CFR 20 (Standards…Radiation)? 10 CFR 20.1002/Scope; –.1003/Definitions, and –.1301/Dose Limits for Individual Members of the Public (USNRC 2002) 20.1002/Scope now states that “…limits in this part do not apply…to exposures from individuals administered RAM and released under §35.75….” 20.1003/Definitions adds “Occupational dose does not include…dose…from individuals administered RAM and released under §35.75….” “Public dose does not include…dose…from individuals administered RAM and released under §35.75….” 20.1301/ Dose Limits for Individual Members of the Public now adds to the exclusion of dose from RAM in sanitary sewers, the following: “…does not exceed 0.1 rem (1mSv) in a year exclusive of the dose contributions from background radiation, from any medical administration to the individual, from individuals administered RAM and released under §35.75, from voluntary participation in medical research programs….” Also, added: “…a licensee may permit visitors to an individual…to receive a radiation dose greater than 0.1 rem (1 mSv) if (1) the radiation dose…does not exceed 0.5 rem (5 mSv) and (2) the authorized user has determined before the visit that it is appropriate.”
9–An Aperçu of Codes, Directives, Guidances, Notices, and Regulations in Brachytherapy 113 Table 2. 10 CFR 19 (Notices, Instructions, & Reports to Workers; Inspections) (Partial Contents) Section
Major contents of section
.3/Definitions
Workers, licenses, restricted areas defined.
.11/Postings notices to workers
a) Post regulations, (i) license & its conditions; (ii) operating procedures; (iii) violations; b) Documents, forms must be conspicuous.
.12/Instructions to workers
Inform about: a) storage, use RAM; b) health protection problems; c) procedures to reduce exposures; d) regulations; e) report conditions, violations; f) response to warnings; g) their exposures.
.13/Notification & reports to individuals
a) Written exposure reports; b) annual exposure reports per workers request; c) other provisions not stated here.
.14/Presence of licensee’s & workers representatives during inspections;
a) Licensee to allow inspections; b) inspectors may meet workers; c) representatives may accompany inspectors during inspections; d) other provisions not stated here.
.15/Consultations with workers during inspections
a) Inspectors may consult privately with workers; b) workers may consult privately with inspectors.
.16/Requests by workers for inspections
Workers may request, without retribution, inspections.
Regulatory Review: What’s Changed in 10 CFR 32 (Specific…Material)? The “new” (4-24-2002) 10 CFR 32 (Specific Domestic Licenses to Manufacture or Transfer Certain Items Containing Byproduct Material) changes are only notational bookkeeping, changing the paragraphs numbers and sections in Part 32 to correspond with the corresponding sections of the new 10 CFR 35.
Regulatory Review: What’s “New” in 126 Sections of 10 CFR 35 (Medical Use…Material)? With 126 sections, we focus only on those of direct interest or applicability to brachytherapy. Tables 4A through 4E summarize, using some shorthand notations, the major contents of the important sections. The bulk of regulatory changes relative to brachytherapy occur in these sections.
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.1101/Radiation Protection Program (RPP)
Major contents of section a) RPP must be developed, documented, implemented, commensurate with extent & scope of licensed activities; b) ALARA for occupational & public doses; c) Annually review RPP content & implementation.
.120/Occupational Dose Limits
a) Annual TEDE 0.05 Sv; sum of DDE & CDE of organs 0.5 Sv; eye DE 0.15 Sv; shallow skin or extremity DE 0.5 Sv; b) Excess DEs must be planned; c) Other provisions not stated here.
.1208/Dose to an Embryo/Fetus
a) 5 mSv dose to embryo/fetus, entire pregnancy, occupational exposure of mother; b) Avoid variations in uniform monthly doses; c) Dose is sum of DDE of mother & radionuclides in mother & embryo/fetus; d) Other provisions not stated here.
.1502/Individual Monitoring of External/Internal Occupational Doses
a) Those likely DE 10% of limits; b) Those in high & very high radiation areas; c) Those likely to receive CEDE of 10% from radionuclides; d) Other provisions not stated here.
.1801/Security of radioactive materials
a) Secure from unauthorized removal or access licensed material stored in controlled or unrestricted areas; b) Licensed material not in storage shall have control and constant surveillance.
ALARA: as low as reasonably achievable CDE: committed dose equivalent CEDE: committed effective dose equivalent
DDE: deep-dose equivalent DE: dose equivalent TEDE: total effective dose equivalent
Components of 10 CFR 35 Applicable to All Forms of Brachytherapy (Tables 4A,B) A new term, “Authorized Medical Physicist (AMP)”, and the training thereof, is defined, as well as types [LDR (low dose rate), PDR, HDR] of remote afterloading units (RAU), including medium dose rate (MDR). Mobile services and medical events are new additions. Roles of management, the radiation safety officer (RSO), and authorized users (AU) supervision of individuals are explained. Dose prescriptions, or written directives (WD) details and procedures are enumerated. Table 4B notes source inventories are now at 6-month intervals. Section 35.75 explains new release criteria for patients (NB: See Glasgow (2002b) for a discussion relative to 192Ir seed patients) (USNRC 1997). Some requirements for mobile medical services are in this section, as well as rules for decay-instorage of RAM.
9–An Aperçu of Codes, Directives, Guidances, Notices, and Regulations in Brachytherapy 115 Table 3B. Unchanged Components of 10 CFR 20 (Standards…Protection…Radiation) Section
Major contents of section
.1901/Caution Signs
Radiation symbol (trefoil) color schema (magenta, purple, black) on yellow & design defined.
.1904/Labeling Containers Radioactive Materials
a) Containers of RAM must be marked either “CAUTION” or “DANGER”, RADIOACTIVE MATERIAL; b) Label must identify quantity, date, radiation levels, kind of material; c) Remove/deface labels on empty containers.
.1906/Receiving/Opening Packages
a) Package receipt & monitoring procedures; b) Carrier notified if wipe test or radiation levels exceed limits; c) Package opening procedures; d) Other provisions not stated here.
.1501/Surveys and Monitoring
a) Make necessary surveys; b) Equipment used for surveys calibrated; c) Excluding direct/indirect pocket dosimeters, NVLAP accreditation for badge processor.
.2001/Waste Disposal
a) By transfer to authorized recipient; b) By decay in storage; c) By effluent release within limits; d) Others provisions not stated here.
NVLAP: National Voluntary Laboratory Accreditation Program
Some Components of 10 CFR 35 (F) Applicable to Manual Brachytherapy (Table 4C) One major change is a requirement to decay output or source activities in 1% intervals. Another section adopts AAPM good practices, per various protocols, for quality assurance of therapy planning systems, as a regulation! (NB: To present and future task group members: Be careful what you write lest it become a required regulation!)
Some Components of 10 CFR 35 (H) for Photon-emitting Remote Afterloaders (Tables 4D,E) In the nine sections, the most significant change is the requirements for MDR and PDR units. Physicians other than AUs, trained in MDR and PDR operation, emergency procedures, and source removal, may work under the supervision of an AU [NB: I will denote them as “substitute authorized users (SAU)”]. For the initial treatment, the AMP and AU or SAU must be present; during subsequent (continuation) treatments, the AMP, AU, or SAU must be immediately available. (NB: In the medical world, that’s by pager!) These requirements are less onerous than the prior requirements of the AU always being present during all treatments. These changes may (or may not!) allow PDR to develop in the United States. Another section adopts AAPM good practices, per various protocols, for quality assurance of RAU therapy planning systems, as a regulation!
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Glenn P. Glasgow Table 4A. Components of 10 CFR 35 (A, B) Applicable to All Forms of Brachytherapy Section
.2/Definitions
Major contents of section a) Authorized medical physicist defined; b) LDR, MDR, HDR, PDR defined; c) Mobile medical service defined; d) Medical event (no more misadministrations!) explained; e) Manual prescribed dose (total sources strength & time, or dose per WD) given; f) Remote prescribed dose (total dose & dose per fraction per WD) given.
.24/Authority Radiation Protection Program
a) Defines a stronger management role; b) Defines & strengths RSO role.
.27/Supervision
Explains role of authorized user (AU) & supervised individuals with respect to process & procedures with RAM.
.40/Written Directives (WD)
a) Written directives required or oral directives with 48 h for written; b) HDR: radionuclide; site, fx dose, # fxs, total dose; c) Others; before tmt: radionuclide; site, dose; before finish: # sources, total source strength & time (or total dose); revisions allowed during treatment.
.41/Procedures…written directives
a) ID patient; b) Administration per WD; c) Check manual, computer dose calculations; d) confirm console data.
.51/Training authorized medical physicist
a) Board certifications; b) Degrees + 1 y training + 1 y experience; c) Preceptor’s written statement regarding training.
Requirements for dosimetry systems (DS), full calibrations (FC), and spot-checks (SC) are described, including those for mobile services.
9–An Aperçu of Codes, Directives, Guidances, Notices, and Regulations in Brachytherapy 117 Table 4B. Some Components of 10 CFR 35 (C) Applicable to All Forms of Brachytherapy Section .67/Requirements for possession
Major components of section a) Leak tests ( 1 mSv/y.
.80/Mobile medical services
a) Facility agreement letters; b) On-site, before use survey meter checks; c) Post-treatment surveys; d) Possession licenses required for all sites.
.92/Decay in storage
a) T1/2 < 120 d; decay to background level; b) Remove labels; keep records.
Table 4C. Some Components of 10 CFR 35 (F) Applicable to Manual Brachytherapy Section .404/Surveys after… implant & removal
Major contents of section a) After implant; source accountability; b) After source removal; keep records.
.406/Source accountability
a)…at all times…in storage& use; record.
.410/Safety instructions
a) Initially, annually…to caregivers; b) Size, type, handling, shielding, visitor.
.415/Safety precautions
a) No room sharing with regular patients; b) Post room (RAM) & visitor limits; c) Emergency equipment for source retrieval from or in patient.
.432/Source calibrations (post 10/24/04)
a) Determine output or activity; b) Positioning in applicators per “protocols”; c) Decay outputs/activities at 1% intervals; keep records.
.433/Decay 90Sr sources
Only AMP shall calculate decayed activity & keep records.
.457/Therapy-related computer systems
a) Acceptance testing per “protocols”; b) Source input parameters; c) accuracy of dose/time at points; isodose & graphics plots; d) localization image accuracy.
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Glenn P. Glasgow Table 4D. Some Components of 10 CFR 35 (H) for Photon…Remote Afterloaders Section
Major components of section
.604/Surveys of patients
Before releasing patient…survey patient & RAU to confirm…returned to safe.
.605/Installation,…,repair
a) Certain source work, i.e., install, adjust, etc., by licensed person; b) For LDR RAU, licensed person or AMP can do certain source work; record.
.610/Safety procedures
a) Secure unattended RAU; b) Only approved individuals present in room; c) No dual operations; d) Written procedures for abnormal situations; posted copies; initial/annual instructions with drills; records.
.615/Safety precautions
a) Control access with interlock; b) Area monitors; c) CCTV/audio for all except LDR RAU; d) For MPD/PDR an AMP & AU or operatoremergency response MD present at initiation & immediately available during treatments; e) For HDR an AU and AMP physically present at initiation, but, during continuation, AMP & AU or operator- emergency response MD; f) Emergency equipment for unshielded source or source in patient.
.657/Therapy-related computer system
a) Acceptance testing per “protocols”; b) Source input parameters; c) Accuracy of dose/time at points; isodose & graphics plots; d) Localization image accuracy; e) Electronic transfer to RAU accuracy.
9–An Aperçu of Codes, Directives, Guidances, Notices, and Regulations in Brachytherapy 119 Table 4E. Some Components of 10 CFR 35 (H) for Photon…Remote Afterloaders Section .630/Dosimerty system (DS) equipment
Major contents of section a) Except for LDR RAUs, NIST/ ADCL calibrated DS; b) 2 y & after service; or, c) 4 y, if intercompared with calibrated DS within 18 to 30 mo.& < 2% change.
.633/Full calibrations (FC) of RAUs
a) Before 1st use; at source exchanges &/or repairs to exposure assembly; b) For T1/2 >75 d, excluding LDR RAUs, quarterly; c) LDR RAUs yearly; d) FC: 5% output/1 mm positions, source retraction, timer accuracy/linearity; e) Tube lengths & functions; f) Quarterly autoradiographs of LDR RAU sources; g) Decay outputs/activities at 1% intervals; h) FC & decay by AMP; keep records; for LDR RAU can use manufacturer’s data for FC.
.643/Periodic spot-checks (SC) of RAUs
a) For LDR RAUs, before1st treatment; for other RAUs 1st use daily; b) Per WP by AMP; c) AMP review by 15 d; d) SC includes: interlocks, status lights, audio & CCTV, emergency equipment, source position monitors, timer, clocks, decayed source activity.
.647/Additional requirements…mobile RAUs
a) Survey meter checks; b) Source inventory; c) All .643 checks; d) Interlocks, status lights, radiation monitors, source positioning, before 1st use, simulated treatment at each address.
ADCL: Accredited Dosimetry Calibration Laboratory NIST: National Institute of Standards and Technology
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Glenn P. Glasgow Table 5. Some Components of 10 CFR 35 (L) (Record Retentions) Record Retention Requirement
Section
Duration of license
.2024/RPP (b) RSO authority.
Duration of program (device)
.2610/Safety procedures for device.
Five years
.2041/Procedures for WP; .2026/RPP changes.
Three years
.2040/WDs; .2061/Meter calibrations; .2067/Leak tests & inventories; .2070/Surveys; .2075/Patient release; .2080/Mobile services; .2092/Decay in storage; .2310/Safety instructions; .2404/Implants & source removals; .2406/Source accountability; .2432/Source calibrations; .2433/Sr-90 decays; .2605/RAU installation, repairs; .2632/Full calibrations; .2643/Spot checks; .2647/Additional mobile records.
Some Components of 10 CFR 35 (L) (Record Retentions) (Table 5) Table 5 summarizes the duration (for license, for program, and for 5 and 3 years) requirements for the retention of records.
Some Components of 10 CFR 35 (M) (Reports…Medical Events…Sources) (Table 6) Misadministration is no more! We now have medical events (ME)! Very careful reading is required for this section, as the ME depends, in some cases, on the difference (presumably lower or higher) in delivered dose and prescribed dose (PD), and in other cases, in exceeding the PD. Moreover, the definitions are not in medical physics terms of absorbed dose in gray (Gy); rather, they are in health physics terms of effective dose equivalent (EDE), and shallow dose equivalent (SDE), in sievert (Sv). Recall that in partial organ irradiation in health physics, organ or tissue weighting factors apply in calculating EDE. As a brachytherapy ME will likely involve adjacent organs, some judgment may be required in deciding on the correct EDE in an ME. Table 6 summarizes the reporting of medical events; reporting requirements are similar to pre-2002 regulations.
9–An Aperçu of Codes, Directives, Guidances, Notices, and Regulations in Brachytherapy 121 Table 6. Some Components of 10 CFR 35 (M) (Reports…Medical Events…Sources) Section
Major contents of section
.3045/Report/notification medical event (excluding patient intervention) (1)
Dose differs from PD more than 0.05 Sv EDE, 0.5 Sv organ/tissue & SDE skin, and, TD, and, TD delivered differs from PD by +20% or falls outside PD range; or single fraction delivered dose differs from single fraction PD +50%.
.3045/Report/notification medical event (excluding patient intervention) (2)
Dose exceeds 0.05 Sv EDE, 0.5 Sv organ/tissue & SDE skin, and, TD from wrong: a) byproduct material; b) administration route; c) person; d) treatment mode; e) leaking source.
.3045/Report/notification medical event (excluding patient intervention) (3)
Excluding migrating permanent implant seeds, dose to skin/organ/tissue other than treatment site that exceeds 0.5 Sv organ/tissue and +50% dose expected from WD.
.3045/Report/notification medical event (excluding patient intervention) (3) (b)
Report any patient interventions producing permanent/physiological damage.
.3045/Report/notification medical event (excluding patient intervention) (3) (c, d)
Notify NRC next calendar day after ME with written report in 15 days; notify referring MD & patient unless referring MD chooses not to for medical reasons; details of reports omitted here.
.3067/Report leaking source.
Report >5 nCi removal contamination within 5 days.
EDE: effective dose equivalent SDE: superficial dose equivalent
PD: prescribed dose TD: total dose
Bulletins, Directives, Guidances, Information Notices, Newsletters, and Regulatory Summaries, for Brachytherapy The USNRC issues to licensees bulletins, directives, guidances, information notices, newsletters, and regulatory summaries as new issues not covered in regulations arise and must be addressed. In some cases, these documents endure for many years, and may actually be incorporated by Agreement States into their regulatory statutes.
Bulletins Apparently there are no recent bulletins pertaining to brachytherapy; the last one was “Release of Patients after Brachytherapy with Remote Afterloading Devices” (USNRC 1993).
Directives Directives appear in several forms. FC86-4, Revision 1–Information Required for Licensing Remote Afterloading Devices, a long-standing (1986) policy and guidance directive, explained the contents for NRC license applications for RAUs (USNRC 1986). While it is not currently on the NRC web site, Illinois (and, I imagine, other states) adopted it, with some changes, into their licensing process for RAUs. FC83-20, Revision 2–Facility Interlocks and Safety Devices for High, Medium, and Pulsed Dose-Rate Afterloading Units, is not on the NRC web site. As the title implies, this release clarified the requirements for
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interlocks and safety devices (USNRC 1983). It appears that issues raised are addressed in the 2002 10 CFR 35 revisions.
Guidances Guidances often discuss evolving technologies. For example, as intravascular brachytherapy developed, the NRC issued several guidance documents (Glasgow 2002a, USNRC 2004a). These were necessary as the “new” 10 CFR 35 applies only to photon-emitting RAUs; beta-emitting RAUs fall into the “emergent technology” category evaluated on a “case-by-case” basis.
Information Notices Information Notices advise licenses of recent concerns usually arising from medical events reported to the NRC. A recent notice discussed failures of HDR RAUs (USNRC 2003).
Newsletters Newsletters, notably “Nuclear Materials Safety and Safeguards” (NMSS), announce medical events and enforcement actions against those who violate regulations. A recent one reported on a hospital’s failure “…to secure…licensed material…” (USNRC 2004b).
Regulatory Summaries Regulatory Summaries often clarify issues about the interpretation of regulations, such as the calibration measurements for brachytherapy sources (USNRC 2002b).
Recent NRC Activities—Specialty Boards and Training Requirements The 2002 revisions in 10 CFR 35 did not address personnel training. On March 30, 2005, the NRC published the final rule (USNRC 2005) regarding specialty boards and personnel training. The rule identifies (on the NRC web-site, not in the published rule!) various approved specialty boards and describes pathways for approval of RSOs, AMPs, authorized nuclear pharmacists (ANPs), and physicians using many forms of by-product materials. This flexible rule offers multiple pathways by which individuals may achieve authorization to perform various tasks or assume authorized titles, e.g., RSO, AMP, ANP, or physician authorized user, while maintaining the integrity of the approval process. One major pathway is the educational degree → experience → specialty examination → certification path. Another major pathway is the supervised experience → preceptor statement path. Table 7 (replete with necessary acronyms) shows five ways an individual, depending on education, experience, and certification status, can achieve authorization as a RSO. For non-physicians, the education requirements are either (a) a bachelor’s degree or graduate degree in physical science, or, engineering or biologic science with 20 college credits in physical science, or, (b) a master’s degree or doctorate in physics, medical physics, or physical science, engineering, or applied mathematics. Experience requirements vary from 1 year to 5 years depending on the authorization, and are shorter for those with higher degrees. Generally, experience must be gained under a certified medical physicist (CMP) or authorized individual, and documented. Preceptors must document the successful completion of any structured training programs and attest to the individual’s competencies and abilities to perform learned tasks independently. In some instances, structured didactic training programs include classroom and laboratory training in topical areas. For example, topical areas for an RSO include: radiation physics and instrumentation, radiation protection, mathematics pertaining to use and measurement
9–An Aperçu of Codes, Directives, Guidances, Notices, and Regulations in Brachytherapy 123 of radioactivity, radiation biology, and radiation dosimetry are allowed. Table 8 shows the requirements for approval as an AMP or ANP. [NB. No single-page synopsis with two tables can adequately describe the seven pages of new federal regulations on specialty boards and training requirements; interested readers are advised to study the new regulations in their entirety to fully understand them! (USNRC 2005)] Similar tables (not presented here) describe approval processes for physician authorized users for use of by-product materials. Table 7. Some Components of 10 CFR 35 (J) Requirements for Radiation Safety Officers Person
Degree or Certification
Experience
(1) Radiation Safety Officer
B or GD in PS; or, E or BS with 20 cc in PS;
and 5 or more and Passes yrs in HP Exam including 3 yrs in AHP
Or, (2) Radiation Safety Officer
M or PhD in P, MP, or PS, E, AM
and 2 yrs full-time training in MP under supervision by CMP, or, in CNM, by physician AU
or, (3) Radiation Safety Officer
1 yr full-time RS under supervision by RSO
Certification Examination
Classroom Laboratory Training
Preceptor Statement
Special Training
and Passes Exam
and 200 h in topical areas
or, (4) Radiation Safety Officer
CMP
and applicable experience
and has written attestation by preceptor
and training in RS, regulatory issues, & emergency procedures
or, (5) Radiation Safety Officer
AU, AMP, or ANP on license
and applicable experience
and has written attestation by preceptor
and training in RS, regulatory issues, & emergency procedures
ANP = Authorized Nuclear Pharmacist B = Bachelor’s Degree BS = Biological Science CC = College Credits E = Engineering GD = Graduate Degree M = Master’s Degree PhD = Doctoral Degree AHP = Applied Health Physics
PS = Physical Science CMP = Certified Medical Physicist RS = Radiation Safety AU = Authorized User CNM = Clinical Nuclear Medicine MP = Medical Physicist or Physics RSO = Radiation Safety Officer AMP = Authorized Medical Physicist HP = Health Physics
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Glenn P. Glasgow Table 8. Some Components of 10 CFR 35(J) Requirements for AMPs and ANPs
Person
Degree or Certification
Experience
Certification Examination
(1) Authorized Medical Physicist
M or PhD in P, MP, or PS, E, AM
and 2 yrs under supervision by CMP
and Passes Exam
or, (2) Authorized Medical Physicist
M or PhD in P, MP, or PS, E, AM
and 2 yrs in CRF under supervision by AU eligible physician
and Passes
or, (3) Authorized Medical Physicist
M or PhD in P, MP, or PS, E, AM
and 1 yr fulltime training in MP and 1 yr full-time experience by AMP eligible MP
(1) Authorized Nuclear Pharmacist
Pharmacy; or, passed FPGEC exam
4000 h in nuclear pharmacy
Classroom Laboratory Training
Preceptor Statement
Special Training
and has written attestation of “competency & independency” by MP preceptor
and training in device operation, clinical use, and treatment planning systems
and Passes Exam
or, (2) Authorized Nuclear Pharmacist
ANP = Authorized Nuclear Pharmacist CMP = Certified Medical Physicist RS = Radiation Safety CRF = Clinical Radiation Facility P = Physics AM = Applied Mathematics PhD = Doctoral Degree FPGEC = Foreign Pharmacy Grad Exam Committee
Current, active license
700 h in structured program with 200 h in topical areas
and has written attestation of “competency & independency” by preceptor ANP
PS = Physical Science AMP = Authorized Medical Physicist AU = Authorized User E = Engineering MP = Medical Physicist or Physics MD = Master’s Degree
Conclusions Understanding codes, regulations, and license conditions has to be the least exciting part of a medical physicist’s job! The federal codes are the basis for state codes, but state codes are not identical to federal codes, even in Agreement States. There is no joy being involved in a medical event or discovering a license
9–An Aperçu of Codes, Directives, Guidances, Notices, and Regulations in Brachytherapy 125 violation during an inspection! Compliance with myriad regulations and license conditions is a challenge. However, by knowing the codes and regulations, one can write a better license with which it is easier for one to comply! To be forewarned is to be forearmed! May you always be in compliance with your federal or state license!
References Benedetto, A. R. (1995). “The Brachytherapy Regulatory Environment: Organization of a Radiation Safety Program” and its appendix “Compilation of Selected Rules for Brachytherapy Use of USNRC Licensed Materials” by J. O. Eichling in Brachytherapy Physics. J. F. Williamson, B. R. Thomadsen, and R. Nath (Eds.). AAPM 1994 Summer School Proceedings. Madison, WI: Medical Physics Publishing, pp. 163–173 and 174–183, 1995. Deye, James A. “Codes and Regulations, Radiation” in Encyclopedia of Medical Devices and Instruments. John G. Webster (Ed.). New York: John Wiley & Sons, 1988. Glasgow, G. P. (2002a). “Nuclear Regulatory Commission regulatory status of approved intravascular brachytherapy systems.” Cardiovasc Radiat Med 3:1–11. Glasgow, G. P. (2002b). “Is an Ir-192 permanent seed implant feasible for prostate brachytherapy?” Brachytherapy 1:195–203. Health Physics News and Notices (2002). “IAEA action plan to combat nuclear terrorism.” Health Physics 82:908–909. Health Physics News and Notices (2003). “IAEA and UPU join forces to protect mail.” Health Physics 84:129–130. Higson, G. R. Medical Device Safety: The Regulation of Medical Devices for Public Health and Safety. Bristol, England: Institute of Physics Publishing, 2001. Mossman, K. L. (2003). “Restructuring nuclear regulations.” Environ Health Perspect 111:13–17. U.S. Nuclear Regulatory Commission (1981). Code of Federal Regulations, title 10, chapter 1, part 19 (Notices, Instructions, and Reports to Workers; Inspections). [Online] http://www.nrc.gov/reading-rm/doccollections/cfr/part019. [December 4, 2004]. U.S. Nuclear Regulatory Commission (1983). FC83-20, Revision 2–Facility Interlocks and Safety Devices for High, Medium, and Pulsed Dose-Rate Afterloading Units. Washington, DC. U.S. Nuclear Regulatory Commission (1986). FC86-4, Revision 1–Information Required for Licensing Remote Afterloading Devices. Washington, DC. U.S. Nuclear Regulatory Commission (1991). Code of Federal Regulations, title 10, chapter 1, part 20 (Standards for Protection Against Radiation; Final Rule) [Online] http://www.nrc.gov/reading-rm/doccollections/cfr/part020. [December 4, 2004]. U.S. Nuclear Regulatory Commission (1993). Bulletin 93-01: Release of Patients after Brachytherapy with Remote Afterloading Devices. [Online] http://www.nrc.gov/reading-rm/doc-collections/gen-comm/bulletins/1993. [December 4, 2004]. U.S. Nuclear Regulatory Commission (1997). Regulatory Guide 8.39: Release of Patients Administrated Radioactive Materials. [Online] http://www.nrc.gov/reading-rm/doc-collections/reg-guides/occupational-health/active/ index.html. [December 4, 2004]. U.S. Nuclear Regulatory Commission (2002a). Code of Federal Regulations, title 10, parts 20, 32, and 35 (Medical Use of Byproduct Material: Final Rule). Federal Register, vol. 67, no. 79 (April 24):20250–20397. [Online]. Washington, DC: Nuclear Regulatory Commission. http://www.nrc.gov/reading-rm/doc-collections/ cfr/part020/part032/035. [December 4, 2004]. U.S. Nuclear Regulatory Commission (2002b). NRC Regulatory Issue Summary 2002-20, “Clarification of Requirements under 10 CFR 35.432, “Calibration Measurements of Brachytherapy Sources” [Online] http://www.nrc.gov/materials/miau/med-use-toolkit/reg-issues-sum.html. [December 4, 2004]. U.S. Nuclear Regulatory Commission (2003). NRC Information Notice 2003-21: High-dose rate remote afterloader equipment failure. November 24, 2003. [Online] http://www.nrc.gov/materials/miau/med-use-toolkit/infonotices.html. [December 4, 2004].
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U.S. Nuclear Regulatory Commission (2004a). Licensing guidance Novoste and Guidant Intravascular Brachytherapy (IVB) systems. May 21, 2004. [Online] http://www.nrc.gov/materials/miau/med-use-toolkit.htmlguidance. [December 4, 2004]. U.S. Nuclear Regulatory Commission (2004b). Newsletter NUREG/BR-0117/04-2: Nuclear Material Safety and Safeguards [Online] http://www.nrc.gov/reading-rm/doc-collections/nureg/brochures/br0117/04-2.pdf. [December 4, 2004]. U.S. Nuclear Regulatory Commission (2005). USCFR title 10, part 35 (Medical Use of Byproduct Material— Recognition of Specialty Boards; Final Rule). Federal Register, vol.70, no. 60 (March 30):16366–16367. [Online]. Washington DC: Nuclear Regulatory Commission. http://www.nrc.gov/reading-rm/doccollections/cfr/part35. [2005, April 19].
Chapter 10
Brachytherapy Facility Design Glenn P. Glasgow, M.S., Ph.D., F.A.A.P.M., F.A.C.R. Department of Radiation Oncology Loyola University Chicago Stritch School of Medicine Maywood, Illinois Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 127 Factors in Program Design . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 128 Radiation Control and Shielding Calculations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 129 Manual Procedures . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 129 LDR and Remote PDR Afterloading . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 129 Remote HDR afterloading . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 131 Contained Liquid Radionuclide Therapies . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 131 Manual Procedure Source Room Facilities . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 131 Receipt of Therapeutic Radioactive Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 131 Source Calibration and Quality Assurance Procedures . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 132 Security of TRAM . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 132 Holding TRAM for Decay . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 133 Imaging (Source Localization) Facilities . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 133 Conventional Imaging Rooms . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 133 Portable Imaging Devices . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 133 Computer Tomography and CT/Simulators . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 133 Isodose/Time Computation Devices . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 134 Conventional Treatment Facilities . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 134 Pulsed Remote Afterloading Fixed Facilities . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 136 PDR Room Shielding . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 136 PDR Facility Issues: Sources, Security, Calibration, and Quality Control Procedures . . . . . . . . . . 139 Features of a PDR Treatment Room . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 139 Permanently Installed HDR RAU Facilities . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 141 HDR Facility Issues: Sources, Security, Calibration, and Quality Control Procedures . . . . . . . . . . 142 Features of HDR Treatment Rooms . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 142 Features of HDR RAUs in Operating Rooms . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 142 HDR Room Designs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 144 Mobile HDR Facilities . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 148 Intrafacility Programs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 148 Interfacility Programs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 148 Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 149 Acknowledgments . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 149 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 149
Introduction What’s new in brachytherapy? How do recent changes in the popularity of brachytherapy procedures affect the design of a brachytherapy program or facility? If, as a young medical physicist hired for a new facility, you are asked to make recommendations about brachytherapy programs or services, plan the facility, and implement your plan, what factors would you consider in these processes? Where would you find shielding data? Where would you find information to help you with these tasks? This chapter will provide some, but not all, answers to these and related questions. Beginning with the last question, recent literature searches for articles on brachytherapy facility design yields few new articles (Glasgow 1999, 2002)
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or abstracts (St. Germain 2002). A popular text on shielding techniques (McGinley 2002) offers new information on maze design for remotely afterloaded high dose rate (HDR) brachytherapy facilities. The National Council on Radiation Protection and Measurement (NCRP) has an active workgroup with draft material (NCRP 2004) that includes a section on facility design. Fortunately, older material, discussed later, about facility design (Broadbent 1984; Gitterman and Webster 1984; Glasgow and Corrigan 1995; Glasgow 1995; Glasgow et al., 1993; Houdek et al., 1994; McKenzie et al., 1986; Stedeford, Morgan, and Mayles 1997; Thomadsen et al., 1983) remains relevant even thought some of the material is not so readily available. This is not a shielding tome; we discuss shielding only in the context of facility design. Classical shielding calculation methods still apply (NCRP 1972, 1976; McGinley 2002). So, lacking definitive new literature for guidance, let’s begin anew!
Factors in Program Design Brachytherapy is widely used. From 5% to 15% of radiotherapy patients may be candidates for brachytherapy. Eighteen or more anatomic sites have been treated with some form of brachytherapy. Figure 1, a 3-year review of 45 major journal articles, shows the percentage of the articles reporting results for 10 anatomic sites. Cervix is the most commonly treated; about 16% of cervical patients receive HDR brachytherapy (Eifel et al., 2004). Interstitial prostate implants with 125I and 103Pd seeds are well established (NB: a new 131 Cs source is in clinical test trials (IsoRay 2004). Breast treatment with the MammoSite® balloon applicator is sufficiently popular that a dedicated single-channel HDR remote afterloader unit (RAU) (Varian 2004) is marketed for that specific application. Remote HDR procedures are common for some sites; special three-channel HDR RAUs (Nucletron 2004; Varian 2004) for gynecologic treatments are available. In large urban areas, hospitals in close proximity use shared mobile HDR units to serve multiple locations. Not all brachytherapy succeeds. Intravascular brachytherapy (IVBT), popular for a time, is being abandoned at many facilities. Those designing a brachytherapy program and facility must recognize: (a) treatment programs are highly dependent on specific physicians and their interests and skills; (b) biologically and clinically, the major goal of HDR brachytherapy remains to achieve, with acceptable complication rates, cure rates equivalent to those achieved with LDR brachytherapy; (c) business plan projections (conjectures) about potential patient numbers for specific programs usually are optimistically high by a factor of two to three (NB: often ignoring similar programs in adjacent facilities!); (d) brachytherapy is, for all staff involved, more labor intensive per case than external beam procedures; and (e) a facility will likely last longer than either the
Figure 1. Of 45 articles reporting results for 10 anatomic sites, the percentage of articles by site.
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physicist designing it or the radiation oncologist using it. Often, particularly in smaller facilities, programs develop because radiation oncologists new to the staff have skills in particular types of brachytherapy and want to do what they know! Conversely, radiation oncologists without prior brachytherapy training are unlikely to lead the charge in developing new brachytherapy programs! Hence, selection of particular brachytherapy technologies depends on realistic estimates of the numbers of patient candidates for treatments; the stability of the radiation oncology staff and their interest and prior experience with the proposed technologies; and realistic estimates of the capital costs, annual costs, and support staff (dosimetrists and physicists) available to support the programs. Brachytherapy is a multifaceted discipline; no one brachytherapy technology is best suited for treating all cancer types and anatomic sites. My on-line literature search of “high dose rate brachytherapy” returned 1669 articles. A quick review indicates some forms of brachytherapy, such as HDR 192Ir brachytherapy, may yield acceptable results in one anatomic site but induce too many complications or inferior cure rates in other sites. Pulsed dose rate (PDR), which presents the technical and radiation protection advantages of HDR while retaining the radiobiological advantages of LDR, is unused in the United States; radiation oncologists are reluctant to adopt it even if it appears to offer advantages over conventional technologies. Multiple brachytherapy technologies likely are required; some necessary but specific programs may treat very few patients yearly. The design of the facility, support space, and selection of quality assurance (QA) equipment are, in some ways, the easiest parts of implementing a program.
Radiation Control and Shielding Calculations Each form of brachytherapy offers the potential of unexpected, unnecessary, and potentially harmful excess radiation exposure to both patients and supporting medical personnel. Regulatory (federal or state) controls apply. Proper facility design can enhance radiation control and reduce personnel dose equivalents (DE). Least anyone not know the applicable DE for occupational personnel and members of the public, they are summarized in Table 1, along with the comments regarding their role in shielding calculations. Shielding calculations, using older concepts and terminologies (exposure rate, X, [R/h], activity, A, [Ci], specific gamma-ray constants at 1 m, Γ, [R m2/h/Ci], inverse reduction of radiation exposure rate with the square of the distance, d, in m, from the source; X=A Γ/d2) are straightforward (NB: they can, of course, be performed in more current medical physics notation; surprisingly, one gets the same shielding thickness!). Patient attenuation, discussed later, is commonly neglected.
Manual Procedures Traditional single-session, in-patient gynecology treatment with manually loaded 137Cs sources is the most common method of manual brachytherapy (Eifel et al., 2004). The 137Cs source inventories established in the 1970s and 1980s have decayed significantly. There is sufficient interest in manual afterloading to warrant production of new 137Cs sources (Radiation Product Design 2004) to replace aging inventories. Interstitial prostate implants with 125I and 103Pd seeds are the most common operative procedure, mostly performed on outpatients. Full-service (calibrated, autoradiographed seeds in sterile packages) purchasing eliminates certain physical plant requirements (calibration equipment, etc.) normally required to support these services.
LDR and Remote PDR Afterloading Remote low dose rate (LDR) and PDR afterloading share a common feature; both require an in-patient hospital room especially designed to accommodate the RAU. LDR never achieved, in the United States, the prominence expected. Fewer than 100 LDR RAUs were sold in the United States and Canada. The
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Glenn P. Glasgow Table 1. Dose Equivalents (DE) and Area Designations for Protection against Radiationa Member Designation
Commentb
Dose Equivalent
General Public
1 mSv Annually
5 mSv allowed via prior regulatory authorization with demonstrated need; visitors per authorized user approval
General Public
0.02 mSv in any one hour
In unrestricted area; not 0.02 mSv per hour
Embryo/Fetus
5 mSv over gestation period
Avoid substantial variations above uniform monthly exposure rate
Occupational Adults (External Whole Body)
50 mSv Annually
Occupational Adult (Eye Lens)
150 mSv Annually
Deep DE (DDE) at 1 cm
Occupational Adult (Skin & Extremities)
500 mSv Annually
Shallow DE at 7 µm over 1 cm2; knees & below; elbows & below
Occupational Minors
0.1 DE allowed adults
Deep DE (DDE) at 1 cm; excludes planned specials
////////////////////////////////////////////
////////////////////////////////////////////
///////////////////////////////////////////
Area Designation
Limit
Commentb
Unrestricted area
No restriction to access
Radiation Area
0.05 mSv in one hour at 30 cm from radiation source or surface radiation penetrates
Accessible to individuals; “Caution” signs may be used
High Radiation Area
1 mSv in one hour at 30 cm from radiation source or surface radiation penetrates
Controlled locked access; alarmed area; “Caution” or “Danger” signs may be used
Very High Radiation Area
5 Gy (5 Sv) in one hour at 100 cm from radiation source or surface radiation penetrates
Controlled locked access; alarmed area; deliberate entry; “Grave Danger” sign used
a
See 10 CFR 20 “Standards for Protection Against Radiation” for complete definitions. Limited comments only; for a full description see 10 CFR 20.
b
manufacturer, Nucletron Corporation (Columbia, MD), supports existing units, but will phase out LDR technology by 2010. After 14 years, and 188 treatments, we recently closed our LDR program at Loyola. Our radiation oncologist, however, notes that a few gynecology patients yearly, whose fragile anatomy cannot tolerate multiple fraction HDR applicator insertions, are candidates for single-fraction remote LDR treatments. We treat these patients with conventional single-fraction interstitial 192Ir template therapy rather than multiple-fraction HDR therapy. PDR brachytherapy had limited use (mostly in California and Arizona) in the United States before restrictive federal regulations hampered its growth. Unrealistic federal regulatory requirements for a radiation oncologist to be present every time the source came out of its safe have been removed in the recently revised federal code (USNRC 2002). New, less strident regulations on radiation oncologists’ presence may (or may not) allow PDR to be used in the United States (NB: this topic is discussed fully in chapter 9 on regulatory aspects).
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Remote HDR Afterloading Remote HDR RAUs are now common is the United States; there are over 500 HDR RAUs in use. Most are permanently installed; we will focus on programs and room designs for these devices. Increasingly, mobile HDR RAUs are used as hospitals share a single mobile HDR RAU moving among multiple locations. However, the increased popularly of daily B.I.D. treatments for certain anatomic sites poses scheduling challenges. We will review program features required for these devices.
Contained Liquid Radionuclide Therapies Some medical physicists are involved in cross-department (radiation oncology and nuclear medicine) programs in therapies involving contained liquid radionuclides designed for specific sites, e.g., 125I in intracranial balloons. Technically, these are brachytherapy (in close proximity to tissue) sources. The types and numbers of these therapies are increasing. However, as they are usually handled and prepared in nuclear medicine “hot” labs, we have omitted them from our discussion, as we do targeted radioisotope therapies (166Ho for liver tumors, 89Sr for bone pain, etc.).
Manual Procedure Source Room Facilities There is a major disconnect between the manual source rooms we actually have in our existing facilities and those we would design for a new facility. Generally, the radioactive materials source rooms in existing facilities are rooms of last resort, and are usually small, cramped, and often far removed from patient treatment rooms or areas. The radioactive materials source room at Loyola is a converted stairwell in the sub-basement without air ventilation or water. Current source room activities are little changed from those of years past as described in prior conference (Broadbent 1984) and summer school proceedings (Glasgow 1992). However, let us focus on the activities we perform in these rooms and how we would design the rooms to better perform these activities if we controlled the capital budget committee!
Receipt of Therapeutic Radioactive Materials It is generally recommended that facilities receive therapeutic radioactive materials (TRAM) at a common location, such as the Radiation Safety (Control) Office (RSO), so that, for all departments (radiotherapy, nuclear medicine, etc.), all records (package receipts, required wipe tests, etc.) are kept by RSO personnel at the central receipt location. Following receipt, RSO health physics personnel can distribute packages to departmental source rooms or hot labs. If a central receipt area does not exist, some receipt functions must occur in the source room. A well-designed room, Table 2, has many components. The room should be readily accessible but reasonably isolated away from public and employee work areas. A large room with solid concrete walls and a strong, locked, wide (to accommodate transportable HDR RAU) door with an automatic closer is ideal. Door postings (warning signs and emergency contact numbers) should be permanently mounted. The interior should be well lit. If applicators are cleaned, an appropriate sink with protected (mesh screening) drain and cleaning supply storage is needed. Rooms usually have a well-lit laboratory bench, appropriate electrical outlets, cabinets with drawers, wall pegboard for tool storage, conventional shop tools, a well-lit work area, storage or book shelves for the ever-present 3-ring binders for receipt/shipping papers, a computer terminal for inventories, a telephone for immediate contact with vendors, an area radiation monitor with alarm levels set to respond to low levels of radiation generated by a single misplaced, unshielded source, high-quality flashlights for locating seeds dropped on the floor, and a storage area for shipping cartons with defaced RAM (radioactive material) labels. Table 2 contains a more complete list of features.
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Glenn P. Glasgow Table 2. Features of a Well-Designed Source Room
Physical Features
Accessible but reasonably isolated (away from public and employee work areas) large room with solid concrete walls; strong, locked, wide door (to accommodate transportable HDR RAU) (with warnings and emergency contact number postings) with automatic closer
Communications
Telephone (outside line for immediate vendor contact); computer (and printer); computer jacks; bulletin boards for regulatory postings
Interior Features
Well lit; water and cleaning supplies, sink with protected drain (mesh screening); workbenches with cabinet drawer storage; adequate electrical outlets above the workbenches; wall pegboards with hooks, bulletin boards; adjustable shelves; seamless tile floor; “Radioactive materials” waste can; film storage in lead bins
Tools
Conventional shop tools; flashlights, source handling tools, tweezers, forceps, etc.
Source Storage
Locked lead safe(s) or lead well; lead “L” blocks, lighted magnification lamp; temporary lead “pigs” for source sorting during preparation
QA Equipment
Area radiation monitor; radiation detection (NaI (Tl), GM) counters, survey meters; wipe test equipment (if needed); well-ionization chamber and electrometer; autoradiography equipment; barometer, thermometer
Transport Equipment
Shielded source transport containers; carts
Empty Container Storage
Cabinet, shelves to storage reusable or disposal containers, empty lead “pigs”
Work Area
Computer area; binders, blank forms, office equipment and supplies, calculators
Applicator Preparation Applicators and accessories (by type) stored in large labeled clear plastic boxes and Storage in cabinets or on shelves Everything Else
All those favorite “gadgets” that you use in your source room!
Source Calibration and Quality Assurance Procedures QA procedures, Table 2, such as package surveys, wipe tests, source calibrations, and source autoradiographs, each require their own instrumentation and radiation protection equipment (source storage safes, long-handled forceps, “L” blocks with viewing windows, well-calibration chambers, survey meters, etc.). Physicists are likely to receive most of their radiation exposures during QA procedures so attention to shields without voids or cracks and processes to reduce exposure are important. One design (Chavaudra 1997) features a mobile “L” shield that slides along the front of a laboratory bench to provide protection at each workstation on the bench. Each QA function requires its own equipment setup and local area shielding so that required tasks can be efficiently and quickly performed in compliance with the ALARA (as low as reasonably achievable) principle.
Security of TRAM Interim storage of TRAM is required. The outer door (with automatic door closer) will be secured (computer locked, dead-bolt lock, etc.); preferably the TRAM in the room should be locked in lead shields or safes. Devices (carriers, carts, etc.) for transport of TRAM and application tools to patient rooms or treatment areas are essential. Often applicators, accessories, and sterilized trays ready and prepared for operating room procedures can be stored in the source room.
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Holding TRAM for Decay Adequate space and shielding for holding TRAM for decay and subsequent waste disposal is required. In some cases the source room will suffice. At universities, large on-site waste disposal facilities are common.
Imaging (Source Localization) Facilities Source localization for isodose computations is common for interstitial operative procedures and manual and remote afterloading procedures. In some instances, if sources or applicators are inserted in an operating theater, source localization can occur there. However, radiographs with portable x-rays devices often exhibit poor image quality and may fail to achieve radiograph geometries (orthogonal pair, stereo shift, multiple radiographs for random seed entry, etc.) required for isodose computations. Generally, greater accuracy is achieved if source localization occurs in an imaging area designed for that purpose. Whatever imaging facility is used, it is imperative to obtain high-quality accurate images required for accurate isodose computations. We omit discussion of magnetic resonance imaging (MRI) and ultrasound imaging, evolving modalities not widely used for isodose planning.
Conventional Imaging Rooms With planning, conventional x-ray rooms in a diagnostic department can be used for controlled geometry (orthogonal, stereo, etc.) radiographs. Conventional x-ray simulators in radiation therapy facilities exhibit excellent isocentricity and, if available, generally provide excellent localization images. However, both modalities require patient transport and setup that potentially can cause undesired applicator movement between the time of imaging and time of treatment.
Portable Imaging Devices Portable x-ray “C-arms” can be used in the radiation oncology department inside a dedicated treatment room (e.g., an HDR vault). They offer the advantage of imaging the patient in the treatment position without moving the patient. However, without proper planning, they exhibit the same problems that occur in C-arms used in operating rooms. Geometrical reconstruction devices can alleviate some of these problems (Glasgow 1998). Liu et al. (2003) report an improved method of C-arm fluoroscopy for isodose computations.
Computed Tomography and CT/Simulators Computed tomography (CT) and CT/simulators in radiation oncology departments increasingly are used for source localization (if CT-compatible applicators are used) as they provide complete anatomic data relative to the source distribution. It is important to image the patient in the treatment position to keep applicators in their correct position relative to the anatomy. Again, use of a CT/simulator requires patient transport and setup that potentially can cause undesired applicator movement between the time of imaging and the time of treatment. Hence, in facility design, any CT/simulator should be physically close to any dedicated brachytherapy treatment room in the department. Indeed, integration of HDR and CT/simulation in one room offers several advantages but could present scheduling difficulties in a truly busy department.
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Isodose/Time Computation Devices Software required for isodose and treatment duration calculations generally runs on a dedicated computer platform or runs as a special program on an isodose computation computer that includes software for external beam isodose computations. In either case, the hardware (computer, digitizer, light boxes, etc.) requires allocated space, electrical outlets, and a proper work environment for bioware (personnel; hopefully your dosimetrist!) to operate software on hardware. With the exception of interactive “real-time” prostate seed planning programs used in operating rooms during prostate implants, and as remote afterloaders can receive their treatment instructions via electronic data transfer, there is no physical requirement to locate the remote HDR adjacent to the HDR RAU. Brachytherapy isodose computation centers should conform to the generally good design principles of any well-designed computer workspace. Even high-end portable “laptops” now have the capacity for brachytherapy isodose planning.
Conventional Treatment Facilities The conventional treatment facility is a common hospital room without special shielding for brachytherapy. Such rooms normally house gynecology patients treated with manually afterloaded 137Cs sources, patients treated with 192Ir seeds in nylon ribbons, and permanent implant patients treated with 125I or 103Pd who require overnight observation post-implant. Larger rooms to accommodate afterloading carts, portable bedside shields, and positioning visitors’ chairs far from the patient are ideal. Usually, by inverse-square law considerations, adjacent hallway exposure rates around large rooms are less than around small rooms. A room adjacent to low occupancy areas (stairwells, elevator shafts, soiled linen storage, etc.) is ideal. Carefully consider the thickness of the floor and ceilings and the nature of the occupied areas directly below and above to determine if supplemental shielding is required. Supplemental overhead lighting may be required as some hospital room lights have inadequate local area illumination needed during manual afterloading procedures. Dedicated (partially or totally shielded) rooms afford the best control of radiation levels to adjacent areas. Older reports describing features of shielded rooms are still relevant and offer excellent discussions of the myriad issues of shielding brachytherapy rooms (Thomadsen et al. 1983; Gitterman and Webster 1984; McKenzie et al., 1986; Stedeford, Morgan, and Mayles 1997; McGinley 2002). Obviously, the thickness of wall shielding material needed is facility dependent, and depends on room size, adjacent area use, and number of patients treated. Lead thicknesses from 0.5 cm to 3.5 cm (to shield an adjacent bed in the opposing room) are typical. Shielded, direct entry doors are sufficiently heavy to require an automatic door closer. For those considering shielding a new room or retroshielding an existing room, older material shielding data, Table 3, for brachytherapy radionuclides are applicable. These data are reasonably consistent for all radioisotopes except 192Ir where some variation (1.2 to 2.0 cm) in tenth-value-layer (TVL) is noted. The newest shielding data (Rivard, Waid, and Wierzbicki 1999) are mass attenuation coefficients for Clear-Pb™ for several radionuclides. These data may be useful for those designing facilities with either direct (window) or indirect (window plus parabolic or convex mirrors) for patient observations. While it is common to neglect patient self-attenuation in shielding calculations the degree of safety gained, Table 4 has been described for gynecology patients treated with manually afterloaded 137Cs sources and 192Ir (Glasgow, Walker, and Williams 1985; Glasgow 2002b). For 137Cs a gynecology patient with a 40-cm lateral pelvic width provides about 50% self-attenuation. For either type of room, radiation levels inside the room can be reduced with localized portable shields. Institute of Physics and Engineering in Medicine (IPEM) Report 75 (Stedeford, Morgan, and Mayles 1997) provides a useful discussion of bedside shields. Figure 2(a) shows a homemade (used with some success), fixed-height, slant shield with a step to accommodate both short and tall personnel; Figure 2(b) shows a conventional adjustable-height radiation shield.
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Table 3. Comparative Selected Broad-Beam TVL (cm) Shielding Data Radioisotope
Concrete
Steel
Lead
Cobalt-60 NCRP 40 & 49a IPEM 75b IPSM 46c Boutroux-Jaffré
20.6 20.6 No data given 22
6.9 No data given No data given 6.7
4.0 4.0 4.6 4.2
Cesium-137 NCRP 40 & 49a IPEM 75b IPSM 46c Boutroux-Jaffré
15.7 15.7 No data given 17.5
5.3 No data given No data given 5
2.1 2.1 2.2 2.2
Ir-192 NCRP 40 & 49a IPEM 75b IPSM 46c Boutroux-Jaffré
14.7 11.3 No data given 14.7
4.3 No data given No data given 4.3
2.0 1.5 1.2 1.6
Au-198 NCRP 40 & 49a IPEM 75b IPSM 46c Boutroux-Jaffré
13.5 13.5 No data given No data given
No data given No data given No data given No data given
1.1 1.1 1.0 No data given
a
Approximate values obtained with large attenuation. No explicit statement that data is broad beam data. c McKenzie et al. (1986). b
Table 4. Effective Transmission Factors for192Ir and 137Cs versus Patient Lateral Half-Pelvic Widths Patient lateral half-pelvic (r) (cm)
Effective transmission factor K(r) for 192Ir (Glasgow 2002b)
Effective transmission factor K(r) for 137Cs (Glasgow 1985)
10
0.72
0.71
12
0.66
0.66
14
0.64
0.62
16
0.55
0.59
18
0.50
0.55
20
0.44
0.50
22
0.40
0.45
24
0.36
0.41
25
0.34
0.38
Generally, properly designed shields reduce bedside radiation exposure levels by more than one-tenth, and proportionally increase bedside work times, Table 5. However, gaining personnel acceptance for use is difficult and requires training; there often are issues associated with shield storage.
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Glenn P. Glasgow Table 5. Radiation Survey and Working Times Per Unit Dose Equivalent for 192Ir
Location Description
•
X = exposure rate in (mR/h); t = time (minutes) •
Behind 1 in. bedside shield, 6 in. from bed
X (mR/h) t mins
6 in. from bedside with shield removed
X (mR/h) t mins
•
Working Time Per Unit Dose Equivalent Iridium Activity in mgRaeg 20
40
1.4 43
2.9 21
29 2
60 1
The work times per unit dose equivalent are the times in which personnel would receive 1 mrem whole body exposure from the exposure rates at each location. The work times are guides to aid in planning your duties so you may minimize your radiation exposure during your shift. For example, for 20 mg Ra, working 6 in. from bedside with the shield removed for 2 minutes and working behind the shield for 43 minutes would result in 2 mrem exposure.
(a)
(b)
Figure 2. (a) A homemade portable bedside shield featuring a slant design and an adjustable step. (b) A commercial radiation shield with adjustable height.
Pulsed Remote Afterloading Fixed Facilities PDR Room Shielding As PDR is currently little used in the United States, there are few facility descriptions (Speiser and Hicks 1994). A remote PDR RAU room, Figure 3, is similar to a LDR RAU room, except types and quantities of radionuclides are different. Generally, a PDR source would be ~1 Ci of 192Ir while an LDR RAU would be about 0.35 Ci of 137Cs (Glasgow 1995). As the lead TVL for 137Cs (2.1 cm) and 192Ir (2.0 cm) are nearly identical, shielding comparisons are easily made (NCRP 1972). St. Germain notes that, in principle, that PDR brachytherapy requires no additional shielding above that required for corresponding continuous
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Figure 3. Floor plan of a pulsed low dose rate (P-LDR) microSelectron-PDR remote afterloading facility [Reprinted from “Safety Programmes for Remote Afterloading Brachytherapy: High Dose Rate and Pulsed Low Dose Rate” by B. L. Speiser and J. A. Hicks in Brachytherapy from Radium to Optimization, R. F. Mould, J. J. Batterman, A. A. Martinez, and B. L. Speiser (eds.). Veenendaal, The Netherlands: Nucletron B. V. © 1994, with permission from Drukkers. Courtesy B. L. Speiser.]
LDR implant if the PDR average hourly exposure and total treatment dose are equivalent to those the continuous LDR treatment (NCRP 2004). Williamson offers a similar discussion, but concluded that, as a practical application, additional wall shielding (1.9 cm lead) was required to satisfy regulatory requirements on allowed exposures in adjacent areas (Williamson, Grigsby, and Meigooni 1995). Figure 4 shows one design of an LDR room (Wilson et al. 1986; Glasgow et al., 1993) and Figures 5(a) and 5(b) show our LDR room at Loyola (Glasgow and Corrigan 1995). Could our shielded LDR room, now closed, be
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Figure 4. A dedicated LDR remote afterloading room with a small maze and viewing window. (Modified from B.M. Wilson et al., Med Phys 13: 608. © 1986, with permission from AAPM. Courtesy J. D. Bourland.]
used as a PDR room? No, the 1 Ci of 192Ir (about 558 mgRaeq) is about four times the activity of 0.35 Ci of 137Cs (about 140 mgRaeq). For gynecology treatments, 70 mgRaeq is a typical loading. For 192Ir, for an equivalent absorbed dose prescription, treatments times would presumably be one-eighth those of 137Cs, e.g.. 1/8 hour. However, during this interval, the instantaneous hallway exposure rate would approach the 0.02 mSv in one-hour limit. Considering a 2.5 m distance from patient to hallway (0.16 inverse square reduction) and 0.5 in. lead (about 0.65 TVL; 0.22 reduction) in the wall, the total reduction is about 0.035. For 558 mgRaeq the exposure rate at 1 m would be 460 mR/h. Applying the reduction factor gives an instantaneous hallway exposure rate of about 16 mR/h. For 1/8 hour, this yields the limit of 2 mR or 0.02 mSv, with no margin of error. While some factors, such as patient attenuation in gynecology patients could reduce the exposure rate, other anatomic sites offer no attenuation and could receive higher absorbed dose prescriptions. Hence, this superficial analysis indicates that additional retroshielding would likely be needed to convert this LDR room to a PDR room. Careful analysis of proposed PDR treatment regimes, duty factors, workloads, and average hourly and weekly exposures must be done in any PDR facility shielding analysis.
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Figure 5a. A small second floor hospital room renovated to house an LDR remote afterloading device for gynecologic treatments. This room features an internal storage closet in which the LDR unit is stored when not in use. A 1.3 cm (1/2 in.) thick lead projection shield beneath the bed shields the area below; a 0.6 cm (1/4 in.) thick lead shield (not shown) is suspended from the ceiling above. Note the compressed air supply, dedicated electric outlet, radiation monitor, remote control and telephone, power-assisted door opener, 1.3 cm (1/2 in.) lead walls shields, and supplemental lights overhead.
PDR Facility Issues: Sources, Security, Calibration, and Quality Control Procedures Per licensure, PDR RAU source security, calibration, and QA procedures are essentially those required for HDR RAUs. The major issue is that the HDR facility likely is in the radiation oncology department, while the PDR room is in the hospital. This presents some issues with respect to required quality control (QC) procedures. If the PDR hospital room is in new construction, then a small adjacent locked room or vestibule could be designed to accommodate new and spent source storage, calibration equipment, and the other QC equipment. However, it is far more likely that a conventional hospital room may be modified to accommodate a PDR RAU and that adjacent support facilities are not available. Hence, sources and calibration and QC equipment must be transported to the PDR room from some remote location. The best one can do is to recognize the procedural issues when the room is being modified, and try to provide
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Figure 5b. Adjacent closet houses the Selectron-LDR, a low dose rate remote afterloading device for gynecologic treatments. Note the emergency source retrieval container.
as much space in the room as possible during modification. Obviously, some plan must be made to store and secure the PDR sources as well as the necessary calibration and QC equipment.
Features of a PDR Treatment Room Features of a PDR treatment room (electrical source-retraction door interlocks, radiation area monitor, brachytherapy bed, emergency equipment, nurse communication system, visual/audible status indicators at nurse station, PDR locked storage, supplemental lighting) are similar to those reported for a remote LDR room (Glasgow and Corrigan 1995). Note that 10 CFR 35.615 (USNRC 2002) requires “…viewing and intercom systems to permit continuous observation…” for a remote PDR room, but does not require the same for remote LDR rooms.
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Permanently Installed HDR RAU Facilities The major topics associated with permanently installed HDR RAUs are: (a) the interim storage (and security) of the HDR source prior/post installation; (b) source calibration and QA procedures (discussed in the context of facility design and space for the equipment); and (c) whether the HDR facility is in a room shared with other equipment (simulator, linac), is in a dedicated HDR treatment room, or is in an operating room for intraoperative procedures. Intrafacility mobility is discussed in the last section on mobile HDR RAUs. Many factors determine which program or programs one develops. Table 6 lists features of a dedicated HDR minor procedures suite. Table 7 describes estimates of facility cost ranges for different programs. Table 8 describes some HDR program options and estimates of their costs.
Table 6. Equipment List: Dedicated HDR Remote Afterloading/Minor Procedures Suite Remote afterloader storage and treatment location; source guide tubes; in-room radiation detector; check source
Anesthesia area: Medical gases and vacuum; designated location and electrical power for patient monitoring equipment; remote displays
Remote afterloader control console
Operating room/procedures light
Visual communications: 2 systems
Applicator storage and medical and nursing supplies
Audio communications
Sink/scrub/gown/glove area
Secure locked door; door interlock
Treatment applicators
Multiple position patient procedure table; x-ray compatible
Applicator positioning clamp: Integral with procedures table
Overhead track-mounted radiography with X-ray generator and X-ray control console; or,
Emergency off buttons at console, in maze, and in room
Mobile C-arm x-ray unit with fluoroscopy; fiducial marker device
Emergency lighting; wall or ceiling mounted
Emergency recovery equipment and container; “Open source” sign, etc.
Emergency power for selected equipment: Audio, video, anesthesia patient monitoring, lighting, radiation detectors and indicators; remote afterloader
Survey meters
Patient monitors
Optional ultrasound imaging system
Optional treatment planning workstation
Table 7. Facility Construction Costsa Program
a
Square Foot Costs
New construction costs for medical building without shielded areas, i.e., offices, etc.
$100
For an HDR suite for minor procedures, i.e., Class B operating room, add $400 per sq ft
$500
$300,000
For an HDR suite for major procedures, i.e., Class C operating room, add $100 per sq. ft.
$600
$360,000
Calendar Year 2004. Assume 400 sq ft for treatment area and 200 sq ft console and support areas.
b
Facility Costs for 600 sq ftb
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Mobile
Low End
High End
HDR Device
$150K–$300K
$250K–$350K
$250K–$350K
Applicators
$100K–Dupl
$50K–Basic
$100K–Many
Table
$35K
$35K
$75K
Van
$25K
Physics QA
$25K
$25K
$50K
Image System
Use Existing
$120K C-Arm
$600K IBU
Vault
Use Existing
$175K–Sm
$235K–Lg
Totals
$310K–$460K
$655K–$815K
$1325K–$1525K
HDR Facility Issues: Sources, Security, Calibration, and Quality Control Procedures HDR interim source storage and security parallels the PDR discussion; in a dedicated facility it may be reasonable to have a secure HDR source storage area adjacent or inside the dedicated room, as this is where the source transfers occur. Prior comments about source room receipt records and inventory records apply. Similarly, as HDR source calibration and QA procedures must occur inside the room where the HDR RAU is located, it is reasonable to provide either space inside the room for calibration and QA equipment, or, a special mobile cart holding all calibration and QA equipment that can be moved to the room.
Features of HDR Treatment Rooms Certain safety features, Table 6, of HDR treatment room (area monitor, door interlocks, emergency equipment, etc.) are likely to be required for licensure in all types of facilities (Glasgow 1996). However, as noted later, for HDR rooms accommodating mobile HDR RAUs it may be desirable to implement these features in a manner different than those with a permanently installed HDR RAU. Note that 10 CFR 35.615 (USNRC 2002) requires “…viewing and intercom systems to permit continuous observation…” for remote HDR rooms.
Features of HDR RAUs in Operating Rooms In addition to the safety features required for permanent installations, operating rooms (ORs) present two unique challenges: shielding staff and maintaining sterile conditions. Two options exist: (1) Locate the OR in the radiation therapy facility, Figures 6 and 7; or (2) locate the HDR RAU in an existing OR. If the former option is used, it may be possible to combine into the design, if desired, both linac intraoperative procedures and HDR intraoperative procedures, as shown in Figure 6. Intraoperative radiation therapy (IORT), which describes external beam IORT, does not describe brachytherapy IORT; however, Dobelbower and Abe (1989) address the general issues of the use of complex equipment in the OR. Local shielding for HDR in ORs can be used effectively (Sephton et al., 1999). St. Germain has addressed the issue of maintaining operating room staff sterile conditions while staff step outside the inner OR while the source is out treating the patient (St. Germain 2002). She suggests an inner/outer room design. During the portion of the procedure when the source is out, the attending staff in sterile gear steps out of the inner operating room into the outer room when sterile conditions are still maintained. She also has designed an adjacent “bail-out” room for storage of the HDR RAU in the event the source sticks in the device and it must be removed from the OR so other procedures can be performed in the OR.
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Figure 6. Dedicated intraoperative radiotherapy facility at the Regional Cancer Center at Baptist Hospital of Miami. The facility is also used as an HDR suite. [Reprinted from “Design and Implementation of a Program for High Dose Rate Brachytherapy” by P. V. Houdek, G. P. Glasgow, J. G. Schwade, and A. A. Abitbol in High Dose Rate Brachytherapy: A Textbook. S. Nag (ed.). Armonk, NY: Futura Publishing Co., © 1994, with permission from Blackwell Publishing.]
Figure 7. Dedicated HDR brachytherapy suite and OR at the U. of Miami’s Sylvester Comprehensive Cancer Center. The suite can be used for intraoperative procedures. [Reprinted from “Design and Implementation of a Program for High Dose Rate Brachytherapy” by P. V. Houdek, G. P. Glasgow, J. G. Schwade, and A. A. Abitbol in High Dose Rate Brachytherapy: A Textbook. S. Nag (ed.). Armonk, NY: Futura Publishing Co., © 1994, with permission from Blackwell Publishing.]
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HDR Room Designs HDR room designs obviously take many forms depending on the program designs and the facility sites. The major difference is between those facilities renovating rooms for HDR RAUs and those building new facilities. We offer a few examples from published literature. Often existing unused or underused vaults (orthovoltage, 60Co, etc.) can be modified, at substantial cost savings relative to new construction costs, as an HDR vault. Usually, existing shielding is adequate and the only costs are those of modifying the vault to accept the HDR RAU and its associated electrical interlocks and communication systems. Figure 8 shows an HDR vault in a modified orthovoltage room (Stedeford, Morgan, and Mayles 1997). In this figure note the additional of a partition in the room to create a short-length entry maze to better shield the door. In contrast, Figure 9 shows a dedicated HDR RAU vault with a long maze and minimum entry door shielding (Stedeford, Morgan, and Mayles 1997). Note in this figure a small storage area for the microSelectron™ HDR and the shielded area inside the vault, which is described as a shielded operating theater. McGinley discusses HDR maze exposure distributions, noting that, in the maze described, exposure rates in the direction from the inside of the maze to the entry door “…fall off slower than indicated by inverse square of exponential behavior” (McGinley 2002).
Figure 8. HDR treatment room in a modified orthovoltage room. [Reprinted from Stedeford, B., H. M. Morgan, and W. P. Mayles, “Brachytherapy Room Design” in The Design of Radiotherapy Treatment Room Facilities, Institute of Physics and Engineering in Medicine. © 1997. Reproduced with permission.]
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Figure 9. HDR treatment room (shielded operating theatre). [Reprinted from Stedeford, B., H. M. Morgan, and W. P. Mayles, “Brachytherapy Room Design” in The Design of Radiotherapy Treatment Room Facilities, Institute of Physics and Engineering in Medicine. © 1997. Reproduced with permission.]
At Loyola, our dedicated HDR facility (Figure 10) was planned as a Class C (full feature OR), but became a Class B limited procedure room. The room accommodates a dedicated C-arm (General Electric Model 9800 Digital Mobile Imaging System) and an ultrasound unit (B & K) for prostate volume studies. The room lacks anesthesia equipment. Patients requiring anesthesia have applicators placed in second floor ORs in the facility; they are then transported to the HDR facility in the basement. Biodex Medical Brachytherapy Table makes an excellent treatment couch that, thus far, has accommodated gynecologic, breast, and head and neck procedures. The room has outer dimensions of 22 ft width by 17 ft depth, inner dimensions of 17.5 ft width by 13 ft depth, with 15 in. thick ordinary density concrete wall and 1.75 in. lead in the power-assisted door. As we are in the basement, the distance from couch to the floor above was about 12 ft with slightly more than 1 ft concrete. The workload was estimated as 50 patients yearly, each an average of 4 fractions for 200 procedures per year. Assuming a constant 10 Ci source (Γ=0.46 R/h per Ci at 1 m) used 2 hours weekly, the exposure rate at 1 m (neglecting patient attenuation) is about 9.2 R/week or 460 R/year. The distance from patient to the door and console area is about 3.75 m, for an inverse square reduction of about 0.07, yielding about 0.64 R/week or about 32.2 R/year. Limiting occupational DE to 5 mSv (0.5 rem) yearly, ALARA, 0.1 mSv/week, requires a wall reduction of 0.1 mSv/week/0.64 R/week, or 0.15, somewhat more than 1 TVL. As the concrete TVL is about 15 cm (about 6 in.), our 15 in. walls (about 2.5 TVLs) provide an adequate margin of safety. The 1.75 in. lead door (about 4.5 cm or 2.25 TVL) provides the same margin of safety. These data are consistent with McGinley’s comment that “Typical room shielding consists of 0.35 m (about 14″) to 0.61 m (about 24″) concrete or 4 to 5 cm lead.” (McGinley 2002). Because adjacent areas were all occupational, the 0.02 mSv in one-hour limit was not a concern around our facility. However, we note, from experience, only one patient can be treated in any one hour.
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Figure 10. A dedicated “Class B” (minor procedures) HDR suite. (Loyola University Medical Center Outpatient Clinic, 2003)
Overhead, the reduction for 3.65 m (about 0.075) and 30.5 cm (about 2.1 TVL; 0.008) was 0.00067. For four procedures weekly, conservatively assuming one full-hour treatment of 2.3 R at 1 m, in one hour the exposure above is 2.3 R (0.0006) or 0.00014 R in one hour, which is 0.14 mR or 0.014 mSv in one hour, which satisfies the 0.02 mSv in one-hour requirement. This yields a safety margin of about 42%. Preferably, one would like a 100% safety margin. Practically, our 42% is acceptable. Due to many assumptions, the actual radiation survey is a more reliable measure than calculation results. If your facility has space for HDR, but no vault, NELCO offers a pre-fabricated vault, Figure 11, that can be installed in a space about 100 sq. ft. (7 ft deep by 10.5 ft wide by 8.5 ft high) (NELCO 2004). The 50,000-pound unit, designed for 2-hour weekly source exposure time, has all necessary features of an HDR vault and can be installed in 3 weeks. Nucletron features an integrated brachytherapy unit (IBU), Figure 12, designed to perform patient preparation, applicator insertion, imaging, isodose computation, delivery, and verification in the unit (Nucletron 2004). The imaging equipment allows fluoroscopy from all directions. The IBU is designed for those centers with large brachytherapy services and is relatively expensive.
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Figure 11. A pre-fab HDR suite. [Courtesy Rick LeBlanc, NELCO, Woburn, MA.]
Figure 12. An integrated brachytherapy unit. [Courtesy E. van ‘t Hooft, Nucletron International B.V.]
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Mobile HDR Facilities A mobile HRD RAU could be an intrafacility device (used at multiple locations within the same facility) or interfacility device [used at multiple (client) locations outside a main (central) facility], Figure 13.
Intrafacility Programs Intrafacility use under a single license is reasonably straightforward. Protocols for the same transport of the HDR RAU through the hospital are required. Each location in the facility where the HDR unit is used must satisfy the safety features of permanent locations previously discussed.
Interfacility Programs For interfacility use, two types of vans are available: The “mobile vault” full-service van that contains not only the HDR RAU and ancillary equipment but also the fully shielded treatment vault, and the van that transports only the HDR RAU and limited ancillary equipment. We focus on the latter as it is more commonly used. For interfacility use both the main user and the client facilities each must hold a license. Those licenses may differ in content but together they cover all aspects of use of the mobile HDR RAU. Federal technical requirements for a mobile service (USNRC 2002) are described in 10 CFR 35.647. Applicable transport regulations must be rigorously adhered to and those transporting must have received hazard materials (HAZMAT) training. Docking (unloading) facilities (tailgate lifts, dock lifts, etc.) may be necessary for the safe transfer of the HDR RAU to and from the van at each treatment facility. The features of the treatment sites may vary depending on the program. Some mobile HDR services transport survey meters, radiation area monitors, and door interlock switches that can be quickly installed and used at each site. They also transport laptop planning computers, small film digitizers, and other necessary QA equipment that can be used at each site. Others design treatment site facilities with all necessary physical plant accommodations that we previously described at permanent facilities. Security of the mobile
Figure 13. “On the road, again!” A mobile HDR van. [Courtesy Plato Lee.]
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HDR RAU at each site and during transport is of great concern. If stored at a client facility for short times, it must be kept in a secure, locked area, such as a source room. The van must have antitheft alarms; the HDR RAU cannot be left in the van overnight. The van must satisfy either applicable state or federal Department of Transportation requirements for the transport of radioactive materials.
Conclusions All brachytherapy programs, manually afterloaded sources, interstitial permanent implants, remote LDR, MDR, HDR, and PDR RAUs have many similar features and the design of the facilities for each program are reasonably straightforward. Key in facility design is a full understanding of how the brachytherapy sources and services are to be used. I trust that this chapter offers those unfamiliar with facility design a reasonable overview of the topic. May you never lose a source!
Acknowledgments I dedicate this chapter to two prominent medical physicists, deceased, with whom I have co-authored articles. Pavel Houdek, Ph.D., exhibited excellent design skills as exhibited in the room layouts I have included in this presentation. Dale Kubo, Ph.D., had many interests, one of which was ensuring well-designed procedures and processes were used in remote afterloading programs. I regret they are not with us at this 2005 AAPM Summer School; each would have contributed excellently to our facility and design discussions. My thanks to Jean St. Germain for sharing draft NCRP materials, Patton McGinley for his material on HDR mazes, and David Keys and Plato Lee for discussions about their mobile HDR programs.
References Boutroux-Jaffré, F. “Photon Emitting Sources,” Chapter 1 in A Practical Manual of Brachytherapy. B. Pierquin and G. Marinello (eds.). Madison, WI: Medical Physics Publishing, pp. 3–21, 1997. Broadbent, M. V. “Brachytherapy Source Storage, Room Design, and Shielding” in Radiotherapy Safety. B. Thomadsen (ed.). New York: American Institute of Physics, pp. 99–116, 1984. Chavaudra, J. (1997). Chapter 6: “Radiation Protection,” Chapter 6 in A Practical Manual of Brachytherapy. B. Pierquin and G. Marinello (eds.). Madison, WI: Medical Physics Publishing, pp. 83–106, 1997. Dobelbower, R. R., and M. Abe (eds.). Intraoperative Radiation Therapy. Boca Raton, FL: CRC Press, Inc., 1989. Eifel, P. J., J. Moughan, B. Erickson, T. Iarocci, D. Grant, and J. Owen. (2004). “Patterns of radiotherapy practice for patients with carcinoma of the uterine cervix: A pattern of care study”. Int J Radiat Oncol Biol Phys 60:1144–1153. Gitterman, M., and E. W. Webster. (1984). “Shielding hospital rooms for brachytherapy patients: Design, regulatory, and cost/benefit factors.” Health Physics 46:617–625. Glasgow, G. P. “Radiation Safety Program for Radiation Oncology” in Advances in Radiation Oncology Physics. J.A. Purdy (Ed.). 1990 AAPM Summer School. AAPM Monograph No. 19. New York: American Institute of Physics, pp. 596–622, 1992. Glasgow, G. P. “Principles of Remote Afterloading Devices” in Brachytherapy Physics. J. F. Williamson, B. R. Thomadsen, R. Nath (eds.). Proceedings of the AAPM 1994 Summer School. Madison, WI: Medical Physics Publishing, pp. 149–161, 1995. Glasgow, G. P. (1996). “Radiation control, personnel training, and emergency procedures for remote afterloading unit.” Endocuriether/Hypertherm Oncol 12:67–79. Glasgow, G. P. “Isodose Planning: Brachytherapy” in Treatment Planning in Radiation Oncology. F. M. Khan, and R. A. Potish (eds.). Philadelphia, PA: Williams & Wilkins, 1998. Glasgow, G. P. “Brachytherapy” in The Modern Technology of Radiation Oncology. J. Van Dyk (ed.). Madison, WI: Medical Physics Publishing, pp. 695–752, 1999.
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Glasgow, G. P. “Equipment Selection and Facility Design” in Proceedings of the American College of Medical Physics 19th Annual Meeting and Workshops. American College of Medical Physics, June 3–4, 2002, Jackson Hole, WY, pp. 263–279, 2002a. Glasgow, G. P. (2002b). “Is an Ir-192 permanent seed implant feasible for prostate brachytherapy?” Brachytherapy 1:195–203. Glasgow, G. P., and K. W. Corrigan. (1995). “Radiation design and control features of a hospital room for a low dose rate remote afterloading device.” Health Physics 69:415–419. Glasgow, G. P., S. Walker, and J. D. Williams. (1985). “Radiation exposure rates near brachytherapy patients containing 137Cs sources.” Health Physics 48:97–104. Glasgow, G. P., J. Daniel Bourland, P. W. Grigsby, J. A. Meli, and K. A. Weaver. “Remote Afterloading Technology.” AAPM Report No. 41. New York: American Institute of Physics, 1993. Houdek, P. V., G. P. Glasgow, J. G. Schwade, and A. A. Abitbol. “Design and Implementation of a Program for High Dose Rate Brachytherapy” in High Dose Rate Brachytherapy: A Textbook. S. Nag (ed.). Armonk, NY: Futura Publishing Co., 1994. IsoRay (2004). Cs-131 Seed. [Online] IsoRay. http://www.isoray.com. [2005, March 30]. Liu, L., D. A. Bassano, S. C. Prasad, B. L. Keshler, and S. S. Hahn. (2003). “On the use of C-arm fluoroscopy for treatment planning in high-dose-rate brachytherapy.” Med Phys 30:2297–2302. McGinley, P. H. Shielding Techniques for Radiation Oncology Facilities, 2nd ed. Madison, WI: Medical Physics Publishing, 2002. McKenzie, A. L., J. E. Shaw, S. K. Stephenson, and P. C. R. Turner (eds.). “IPSM Report 46: Radiation Protection in Radiotherapy.” London, Institute of Physical Sciences in Medicine, 1986. National Council on Radiation Protection and Measurements (NCRP). NCRP Report No. 40. Protection Against Radiation from Brachytherapy Sources. Bethesda, MD: NCRP, 1972. National Council on Radiation Protection and Measurements (NCRP). NCRP Report No. 49. Structural Shielding Design and Evaluation for Medical Use of X-rays and Gamma Rays of Energies up to 10 MeV. Bethesda, MD: NCRP, 1976. National Council on Radiation Protection and Measurements (NCRP). NCRP Report SC 91-1. Draft Report; Untitled. Bethesda, MD: NCRP, 2004. (Courtesy J. St. Germain; Private Communication, Dec. 10, 2004). NELCO (2004). (No Date). HDR Shielded Booth. [Online] NELCO. http://www.nelco-usa.com. [2005, March 30]. Nucletron B. V. (2004). (July 7, 2004). Remote Afterloaders. [Online]. Nucletron B.V. http://www.nucletron.com/ content. [2005, March 30]. Radiation Product Design (2004). 137Cs Sources. [Online]. Radiation Product Design.http://www.rpdinc.com. [2005, March 30]. Rivard, M. J., D. S. Waid, and J. G. Wierzbicki. (1999). “ Mass attenuation coefficients of Clear-Pb for photons from I-125, Pd-103, Tc-99m, Ir-192, Cs-137, and Co-60.” Health Physics 77:571–578. Sephton, R., K. R. Das, J. Coles, W. Toye, and P. Pinder. (1999). “Local shielding of high dose rate brachytherapy in an operating theatre.” Australas Phys Eng Sci Med 22:113–117. Speiser, B. L., and J. A. Hicks. “Safety Programmes for Remote Afterloading Brachytherapy: High Dose Rate and Pulsed Low Dose Rate” in Brachytherapy from Radium to Optimization. R. F. Mould, J. J. Batterman, A. A. Martinez, and B. L. Speiser (eds.). Veenendaal, The Netherlands: Drukkers, 1994. Stedeford, B., H. M. Morgan, and W. P. Mayles. “Brachytherapy Room Design” in The Design of Radiotherapy Treatment Room Facilities. York, England: Institute of Physics and Engineering in Medicine (IPEM), pp. 77–89, 1997. St. Germain, J. (2002). “Shielding of HDR, IVB, and PET/CT facilities.” Health Physics 82:S183 (Abstract). Thomadsen, B., J. van de Geijn, D. Buchler, and B. Paliwal. (1983). “ Fortification of existing rooms used for brachytherapy patients.” Health Physics 45:607–615. U.S. Nuclear Regulatory Commission (2002a). Code of Federal Regulations, title 10, parts 20, 32, and 35 (Medical Use of Byproduct Material: Final Rule). Federal Register, vol. 67, no. 79 (April 24):20250–20397. [Online]. Washington, DC: Nuclear Regulatory Commission. http://www.nrc.gov/reading-rm/doc-collections/ cfr/part020/part032/035. [December 4, 2004]. Varian Medical Systems (2004). (July 7, 2004). News and Events. [Online]. Varian Medical Systems. http://www. varian.com/comp/040707.html. [2005, March 30].
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Williamson, J. F., P.W. Grigsby, and A.S. Meigooni. “Clinical Physics of Pulsed-Dose-Rate (PDR) Remotely Afterloaded Brachytherapy” in Brachytherapy Physics. J. F. Williamson, B. R. Thomadsen, and R. Nath (eds.). Madison, WI: Medical Physics Publishing, pp. 577–616, 1995. Wilson, B. M., et al. (1986). Presentation at the AAPM 28th Annual Meeting. Abstract. Med Phys 13:608.
Further Reading Gifford, D., and T. J. Gooden. (1990). “ An analysis of personnel dose records which justifies the application of costbenefit analysis techniques in the design of an afterloading facility and the use of controlled areas and systems of work within suite to control occupational exposure.” Br J Radiol 63:214–218. Klein, E. E., P. W. Grigsby, J. F. Williamson, A. S. Meigooni. (1993). “Pre-installation empirical testing of room shielding for high does rate remote afterloaders.” Int J Radiat Oncol Biol Phys 27:927–931. Kubo, H. D., G. P. Glasgow, T. D. Pethel, B. R. Thomadsen, and J. F. Williamson. (1998). “High dose-rate brachytherapy treatment delivery: Report of the AAPM Radiation Therapy Committee Task Group No. 59.” Med Phys 25(4):375–403. Also available as AAPM Report No. 61. Nag, S., R. Dobelbower, G. P. Glasgow, G. Gustafson, N. Syed, B. Thomadsen, J. F. Williamson. (2003). “Intersociety standards for the performance of brachytherapy: A joint report from ABS, ACMP, and ACRO.” Crit Rev Oncol/Hematol 48:1–17.
Chapter 11
Calibration for Brachytherapy Sources Larry A. DeWerd, Ph.D., F.A.A.P.M. University of Wisconsin Madison, Wisconsin Introduction to the Calibration of Sources and Measurement Devices . . . . . . . . . . . . . . . . . . . 153 Brachytherapy Source Standards . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 153 Importance of Calibration and Standards in the Clinical Situation . . . . . . . . . . . . . . . . . . . . . . . . . . 154 Calibration Concepts and Quantities . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 156 Calibration Quantities and Units . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 156 Photon Sources . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 156 Beta Particle Sources . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 157 Primary Standards . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 158 Photon Sources for Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 158 Beta Sources for Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 161 ADCL Standards . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 162 ADCL Source Calibrations and Ophthalmic Applicators . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 162 Characteristics of Well Chambers for Brachytherapy Sources . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 163 Well-Chamber Characteristics for Photon Sources . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 166 Well-Chamber Characteristics for Beta Sources . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 166 Clinical Calibration Using Well Ionization Chambers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 167 Clinical Calibration of Photon Sources . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 168 Clinical Calibration of Beta Sources . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 169 Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 169 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 169
Introduction to the Calibration of Sources and Measurement Devices Brachytherapy uses encapsulated radioactive sources to deliver radiation dose to tissues near the source. These sources need calibration before clinical use; this activity is usually accomplished with a well ionization chamber. This chapter discusses source calibration as determined by primary standards in use, the procedures for measuring air kerma strength of brachytherapy sources for both low dose rate (LDR) and high dose rate (HDR) sources, and the absorbed-dose-to-water at 2 mm for beta sources. The use of well chambers for routine clinical source calibration is also discussed. Characterization of brachytherapy sources requires good initial measurements so that the patient dose is known with confidence.
Brachytherapy Source Standards Radiation emissions from brachytherapy sources involve either γ, β, or x-rays, or a combination of these radiations. The majority of sources in use emit γ rays. The type of radiation and energy has an effect on the calibration methodology. The calibration of the source can greatly enhance the precision of the treatment. This is a rapidly changing field, especially concerning intravascular brachytherapy where the use of gamma-emitting 192Ir sources and the 32P beta-emitting source from Guidant used for intravascular brachytherapy applications have come and gone in the past few years. Most of the instrumentation standards for brachytherapy sources are at the U.S. National Institute for Standards and Technology (NIST) or are directly traceable to NIST standards. Brachytherapy sources can be divided into three classes. The first is low energy—LDR sources generally used for prostate implants. The second is high energy—both LDR and HDR sources used for a number
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of treatment modalities. The third is beta sources—used for ophthalmic applicators and for intravascular brachytherapy. Depending on the energy of the source, different reference points are used. For example, for low-energy LDR sources, the reference distance is 1 cm; for beta sources, it is generally 2 mm or 5 mm. A short distance from the source may lead to greater uncertainties in the calibration because of the high-dose gradients of these sources.
Importance of Calibration and Standards in the Clinical Situation A common clinical brachytherapy dosimetry practice of the past was to accept the manufacturer-specified source strength without verification as a basis of treatment planning. Acceptance of the manufacturer’s stated source strength without measurement by the clinical physicist is an unwise approach as discussed by AAPM Task Group 40 (TG-40) (Kutcher et al., 1994). To quote from this report, “Each institution planning to provide brachytherapy should have the ability to independently verify the source strength provided by the manufacturer.” It is also recommended by TG-40 that a minimum of 10% of the LDR seeds be calibrated. As has become evident, it is essential that each new HDR 192Ir source be calibrated and that it have a periodic quality check at least once in its clinical lifetime (Glasgow et al., 1993). The need for individual source calibrations is further underscored when the differences in source construction for the same radionuclide, e.g., 125I, are considered. These differences greatly affect the response per unit source strength of calibration instruments. The radiological characteristics of a brachytherapy source are strongly dependent on the source encapsulation and the radioactivity distribution within the source, as discussed in Rivard, Kirk, and Leal (2005). The radioactive material may be continuously distributed or divided in sections within the encapsulation. Figure 1 shows a contact autoradiograph of two 125I sources; Figure 1a shows the 125I on a solid wire and Figure 1b shows the 125I deposited on two balls. Note the difference in the “isodose lines” of Figure 1. The uniformity and precision of the radiation from the brachytherapy source can be greatly affected by the distribution and encapsulation so that individual sources need to be measured. The TG-40 report (Kutcher et al., 1994) elaborates on these ideas. In addition, it has been recommended (Collé 1999) that the uniformity of seed and line sources be evaluated in terms of absorbed-dose-rate-towater at a distance of 2 mm from the source center, both longitudinally and perpendicular to the source axis (equatorial) in a tissue equivalent medium. In LDR applications of 125I and 103Pd seeds, there has been a proliferation of new seeds. To maintain consistency in seed dosimetry, procedures have been published (DeWerd et al., 2004) to assure the calibration and dosimetry values. Included in this publication are quality assurance (QA) procedures for maintaining the air-kerma strength standards for a given model seed. On each Accredited Dosimetry Calibration Laboratory (ADCL) calibration report there is a date to which the given seed is traceable to NIST. This assures that the correct air kerma strength is associated with the correct published dose-rate constant. The American Association of Physicists in Medicine (AAPM), American Brachytherapy Society (ABS), and U. S. Food and Drug Administration (FDA) recognize the need for accurate calibration in terms of well-defined physical quantities. The report of AAPM TG-56 states that the medical physicist should independently measure 10% of the sources to be implanted in the patient (Nath et al., 1997). Recently, other third parties have offered services to measure seeds. Sterilization procedures and insertion of the seeds into sterilized needles have given incentive to this activity. The medical physicist is still responsible for the dose given to the patient so the medical physicist should independently validate the output of seeds. Manufacturers assign uncertainties for their sources; however, many times these uncertainties are not met or are larger than desired for clinical dosimetry. Details on these measurements were given in the 1994 AAPM summer school publication (DeWerd and Thomadsen 1995) and in a course at the Radiological Society of North America (RSNA) (DeWerd 1997) where it is demonstrated that some calibrations actually fell outside ±10%. There are a number of anecdotal examples of the manufacturer stated activity being discrepant by 7% or greater for all source types for both LDR and HDR measurements. While
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Figure 1a. Surface contour plot of an Amersham model 6711 LDR 125I source made with an autoradiograph.
Figure 1b. Surface contour plot of a North American Scientific model MED3631-A/M LDR 125I source made with an autoradiograph.
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the ±10% limit is workable for shipping or diagnostic purposes, each clinic should realize the importance of calibrating each source, since a 10% error can have deleterious and unacceptable effects on patients. Just as with 60Co teletherapy sources, individual users should calibrate each brachytherapy source as recommended by AAPM TG-41 (Glasgow et al., 1993). The calculation of dose to a patient depends on the accuracy of the well-type chamber or source calibration. Therefore, the accepted practice is to obtain a well chamber calibration from an ADCL. The ADCL is directly traceable to NIST and each maintains a high degree of precision in the transfer of calibration values to the clinical field instrument. The Calibration Coefficient provided by the ADCL is used to correct an instrument’s reading to obtain the reference value, which is the value of a quantity that has a known accuracy related to the primary value from a national laboratory. Calibration factors should be obtained for all instruments used to measure source outputs, e.g., for the ionization chamber and the electrometer used to read the charge accumulated during the measurement by the chamber. AAPM TG-40 (Kutcher et al., 1994) recommends that brachytherapy sources (or well chambers) have calibrations with direct or secondary traceability to national standards.
Calibration Concepts and Quantities Calibration Quantities and Units Source description quantities fall into one of two categories: measures of source output as emitted through the encapsulation and measures of radioactivity actually contained in the encapsulation. The source output is the quantity that is generally measured. When a given source is measured, it then reflects the radiation emitted which accounts for the encapsulation thickness variation and other parameters. This output is given in different units but the endpoint for dosimetry is absorbed-dose-in-water, with the unit being Gy [J/kg]. Apparent activity has been used in the past but it is not a quantity with a NIST-traceable standard. The quantity includes self-absorption or reduction of source output by its encapsulation and is obtained by dividing the measured source output by a known output value for the “bare” output for the radionuclide. This quantity should not be used for medical dosimetry as explained in appendix E of the update of the AAPM TG-43 report (Rivard et al., 2004). Photon Sources The quantity used to measure the output for photon-emitting sources is air kerma strength. The quantity kerma describes the first step in energy dissipation by indirectly ionizing radiation, transferring energy through one of the interactions to charged particles. Thus, the quantity kerma is defined as the energy transferred to charged particles per unit mass in a volume, as the volume limits to a point. Note that the medium (e.g., air, water, or tissue) must be specified. In modern brachytherapy dosimetry, air kerma strength is used, which includes the distance, and the TG-43 dosimetry equation is used to calculate absorbed dose. Air kerma strength is the quantity endorsed by the AAPM (AAPM 1987; Williamson and Nath 1991) and is similar to reference air kerma rate, which is endorsed by many European and international advisory bodies (IAEA 1999; ICRU 1997). Air kerma strength (AAPM 1987), SK, is defined as the product of airkerma rate (ICRU 1980) in free space, K" (d )[Gy / h ] , corrected for air attenuation and scatter, measured along the transverse bisector of the source, and the square of the measurement distance, d [cm],
SK = K" ( d ) ⋅ d .2
(1)
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Generally, the distance of measurement is corrected to a “reference distance” of 1 m. As recommended by the AAPM TG-32 report (AAPM 1987), the units for air kerma strength, represented by U, are µGy • m2 • h–1 or cGy • cm2 • h–1. Notice that in any of the systems of specification, the reference distance, d0, need not be equal to the measurement distance, d. The ADCL provides calibrations of sources and well chambers in terms of air kerma strength. Beta Particle Sources Beta sources are absorbed in much smaller distances and thus need calibration in closer proximity to where they are used. Thus, the standard distance of 2 mm in water has been chosen as the calibration and delivery distance. This is the distance as chosen by AAPM TG-60 (Nath et al., 1999). In addition, for beta sources a calibration of absorbed-dose-to-water at a reference distance is the quantity of choice. This is accomplished by calibration after passing through a thickness of solid material that approximates water to have similar attenuation properties. NIST has chosen A150 plastic for this purpose (Soares, Halpern, and Wang 1998). In the application of these sources to intravascular situations, lesions in the coronary arteries are on the order of 2 to 4 cm in length in arteries with diameters of 3 mm to 5 mm. Therefore, radioactive line sources of a very narrow diameter, less than 1 mm, are used to deliver dose. Typical geometries include encapsulated seed trains or equivalent type line sources. The physical length of these sources vary but are generally between 3 and 6 cm to adequately cover the lesions undergoing treatment. The uncertainty involved in measurements of reference absorbed dose rate for betas at NIST is large, being given at 15% for k=2 or the 95% confidence level. However, recent communications from NIST have stated that this value has been reduced to 8% because of improvements in geometry and other methodology. However, since it was so large, an attempt was made to quantify the contained activity for one 90Sr source model. The importance of contained activity as a metric for source strength is to provide comparison between model predictions and dosimetry measurements. Monte Carlo calculations predict dose per history, where a history represents the interactions undergone by a single emitted photon or electron. The number of histories can be related to contained radioactivity by disintegration probabilities and branching ratios for complicated decay structures. Thus, Monte Carlo models predict dose rate per unit contained activity. Contained activity for beta-emitting sources can best be determined from a destructive measurement, which involves dissolving a source in a liquid medium that captures all of the contained activity into an aqueous solution (Collé 1999). By a suitable dilution of this solution, contained activity can be determined with a high degree of accuracy by the liquid scintillation technique. A contained activity calibration of a seed or wire beta-emitting source can be transferred to a well-type ionization chamber. This transfer then results in a method to specify a source in terms of contained activity rather than reference absorbed dose rate. The preferred quantity is absorbed dose to water at 2 mm (Nath et al., 1999) and thus, the quantity of contained activity needs conversion to this reference absorbed dose. Note that the problem that could be associated with the contained activity for routine clinical use is the dependence on the generalization of this quantity from the few sources that have been dissolved. Thus, variations in the cladding of the seed sources or other differences between dissolved sources and the source being measured are ignored. The recommended and most generalized quantity for beta source strength is the reference absorbed-dose-ratein-water. Ophthalmic sources also are calibrated to absorbed-dose-to-water; however, these sources are calibrated in terms of surface absorbed dose rate to water (Soares 1991). Since these sources have a distributed dose rate across their surface, the quantity is specified for the central 4 mm diameter, and the variation over the surface of the applicator becomes important.
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Primary Standards The primary standard at NIST is obtained either from calorimeters, an extrapolation chamber, standard chambers with precisely known volumes for high energies, or from free-air chambers (FACs) at low energies. Each of these would be considered a primary measurement device. For example, a FAC measures parameters that are directly in the definition of air kerma, and thus, it falls into the class of an absolute or primary dosimeter. These primary measurements are then transferred to the ADCL, which then transfers the calibration to the clinic through use of a reference class well chamber. The uncertainty in value is increased at each step, but the physicist at the clinic can expect that the uncertainty in the value for brachytherapy measurements from the ADCL is within 2% to 5%, depending on the radionuclide. Photon Sources for Brachytherapy NIST has various means to calibrate LDR photon sources, either with the Wide Angle Free Air Chamber (WAFAC) for low-energy sources or with an ionization chamber for high-energy sources. Currently, there is no direct NIST calibration for HDR 192Ir sources. However, there are primary calibrations on thimble or spherical chambers with external beams, which then can be used for a “primary” calibration on HDR sources. This technique is explained further below. LDR Sources. NIST developed a standard for 192Ir sources (Loftus 1980) using graphite cavity ionization chambers. The procedure has also been described in the literature (Verhaegen et al., 1992). The NIST procedure was to use a large number of LDR sources to provide adequate signal for the relatively insensitive cavity chamber used. The calibrated seeds were then measured individually in a 3.44-liter aluminum re-entrant ionization chamber to transfer the calibration in the future. Thereafter, the re-entrant chamber was used (Weaver, Loftus, and Loevinger 1988). This calibration was only made for two types of LDR 192 Ir seeds; additional types of seeds would require a redundant calibration of the type described by Loftus. The uncertainties at k=2 (95%) in these measurements at NIST is 2.7% for LDR 192Ir sources. The aluminum spherical reentrant chamber has a consistency of 0.2% to 0.3%. NIST uses the WAFAC to collect the radiation from low-energy, photon-emitting brachytherapy sources, such as 125I and 103Pd (Seltzer et al., 2003). For short-lived LDR brachytherapy sources, such as 125 I, a well-type chamber is used that was standardized by using sources calibrated against NIST standards, such as the WAFAC. This well chamber is used for consistency measurements. Figure 2 is a schematic
Figure 2. Schematic of the NIST WAFAC. [Modified from Seltzer et al. (2003).]
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of the WAFAC, modified from Seltzer et al. (2003). The front window of the WAFAC is located 30 cm from the source, which is rotated. There are two volumes used to determine the source strength for a solid angle of acceptance. The aluminum filter between the seed and the entrance window of the WAFAC is generally used to filter out titanium K-edge characteristic x-rays at 4 keV. With the use of the WAFAC, the standard for these low-energy photon sources was established in January 1999. The history of the standard for low-energy photon brachytherapy sources is outlined below for clarity. In 1984 Loftus calibrated 125I sources using the Ritz Free air chamber (Seltzer et al., 2003). This calibration included contributions from titanium K-edge x-rays. In 1993, Loevinger introduced the WAFAC and added the aluminum filter to eliminate the titanium x-rays. NIST completed testing of the WAFAC in 1998 and developed a plan with the ADCLs and AAPM for implementation of the new standard. The Amersham model 6711 125I source was used for this testing as its calibration has been well characterized. On January 1, 1999, NIST introduced the new standard; Amersham models 6702 and 6711 were based on measurements performed in 1998. Through the model 6711 source, it was found in 2000 that measurements made in 1999 on various seed models had a 3% to 6% error for all 125I and 103Pd sources. For example, the difference in calibration coefficients for the seeds measured in 2000 versus the standard given in 1999 for the University of Wisconsin ADCL is given for 125I in Table 1 and for 103Pd in Table 2. This calibration change in the WAFAC standard resulted in calibration reports having dates of source calibration in 2000 or later as explained further in DeWerd et al. (2004). All the seed models are now based on a calibration value in 2000 or later. Since the dose rate constant for the source was determined from the NIST WAFAC air kerma strength, the old value of the dose rate constants have been changed accordingly to reflect the value from 2000 or after. Finally, the well chambers can be calibrated for LDR sources by the ADCL using a NIST-calibrated source. HDR Sources. As mentioned previously, NIST does not calibrate HDR 192Ir sources. The traceability of these sources is through NIST-calibrated ionization chambers at cesium and x-ray beam, M250, energies. These two beams are presently used to determine an interpolated value at 397 keV for calibration as given in Goetsch et al. (1991). Then a calibration method at multiple distances or with a shadow
Table 1. UW ADCL Observed Change in 125I Calibration for 2000 from 1999 NIST Standard Manufacturer and Model
Percent change in 2000 calibration to 1999 standard
North American Scientific MED3631A/M
–2.4%
Best model 2301
–3.3%
Mills 125SH/125SL
–8.6%
Syncor Pharmacy model
–4.4%
Draximage model LS-1
–9.2%
Table 2. UW ADCL Observed Change in 103Pd Calibration for 2000 from 1999 NIST Standard Manufacturer and Model
Percent change in 2000 calibration to 1999 standard
North American Scientific MED3633
–4.5%
Theragenics TheraSeed model 200
–4.2%
Int Brachy 1031L
–6.1%
Best model 2335
–6.5%
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shield is used with the above-interpolated value. A recent paper has expanded this concept to using a calibration coefficient for each energy line of the HDR 192Ir spectra, but the results agree within 1% (van Dijk, Kolkman-Deurloo and Damen 2004). David Rogers has found (private communication) that the differences are within 0.5%. The standard Goetsch technique has been improved and a comparison made between HDR manufacturers and sources (Stump et al., 2002). This work found that in the calibration of all sources, both the old and new Varian and Nucletron sources in various catheters agree in their calibration within 0.7%. In addition, an intercomparison was made with the Henri Becquerel Laboratory in Saclay, France, and agreement was found to less than 1% (Douysset et al., 2005). The uncertainty analysis for all of these calibrations is well outlined in Stump et al. (2002) to be ±2.1% at the coverage factor of k=2 or the 95% confidence level. The air kerma strength determination requires a calibration of the ionization chamber for the energy spectrum of 192Ir. The spectrum of an HDR 192Ir encapsulated source can be exposure-weighted to obtain an effective energy of 397 keV. This energy falls within 2% of the halfway point between 137Cs (662 keV) and the effective energy (146 keV) of a 250 kVcp medium filtration x-ray beam (HVL = 3.2 mm Cu). NIST provides air kerma calibration coefficients for both of these beam qualities. The ionization chamber wall must be thick enough to block all electrons emanating from the source or capsule, and to provide charged particle equilibrium (CPE) for the highest energy secondary electrons present in the 137Cs beam. Thinner-walled chambers require a buildup cap of at least 0.3 g/cm2; in the absence of a 137Cs buildup cap, a 60Co cap suffices. To establish the response of the chamber, the attenuation of the cap (Ax) must be eliminated. The response of the chamber alone, Nch, to a beam, is given by Nch = Nx Ax ,
(2)
where Nx is the NIST-traceable calibration coefficient for the chamber and cap. The calibration coefficient for 192Ir can then be obtained by interpolation (Stump et al., 2002) between the correction factors for the two bracketing energies from the following equation. (AwNx)Ir = [(AwNx)x-ray + (AwNx)Cs]/2 ,
(3)
where (Nx)Ir, (Nx)x-ray, and (Nx)Cs are the exposure calibration coefficients for 192Ir, 250-kVcp x-rays, and Cs respectively, and (Aw)Ir, (Aw)x-ray, and (Aw)Cs are the corresponding wall effects. From equation (3) the calibration coefficient (Nx)Ir can be determined. The air kerma strength can then be measured in free space along the transverse bisector of the source at a large distance relative to the dimensions of the source and detector. Specifying “free space” requires corrections for absorption and scattering in air, ionization chamber walls, or the room physical environment present in the calibration geometry. This methodology has been explained in previous publications (DeWerd and Thomadsen 1995; Goetsch et al., 1991; Stump et al., 2002). An improved jig giving less scatter and a better setup is described by Stump et al. (2002). Multiple distances are used to account for scatter and error in the setup distance. A comparison of measurements done at one distance versus those done at multiple distances show that the precision for the single distance measurements is ±3%, whereas the precision for multiple distances is ±1.5%. At typical brachytherapy treatment distances, ranging from a few millimeters to a few centimeters, conventional ionization chambers cannot be treated as point-like detectors. In addition, at these short distances, air kerma measurements are extremely sensitive to positional uncertainties. Both of these issues contribute a major part to the overall calibration uncertainty. Corrections for the finite size of idealized chambers (Dove 1959; Kondo and Randolph 1960) become large 137
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for distances less than 5 cm. A method used to avoid distance uncertainties is to shield the chamber with a lead block to measure the scatter. Of course at the larger distance used, the scatter is a larger percentage of the signal. The shielded readings are subtracted from the unshielded reading to yield the net readings. This method results in very small readings on the ionization chambers and leakage must be accounted for or a larger volume chamber must be used. The method using the series of distances with precisely measured separations can be used to account for scatter and distance setup error. The actual distance for a reading is d′ = d + c ,
(4)
where d′ is the true center to center source chamber distance, d is the apparent center-to-center source chamber distance, and c is the error is setup distance (c can be positive or negative). The contribution of the room scatter radiation, K" S , is included in the measured air kerma rate, K" (d ) . Therefore the air kerma value from the primary radiation only, K" p (d ), is given by K" p (d ) = K" ( d ) − K" s .
(5)
The independence of the room-scatter air kerma rate contribution with respect to distance has been shown to be the case over distances of 10 cm to 40 cm (Goetsch et al., 1991). The air-kerma strength, SK, can then be written based on the inverse square law relationship, as
[
SK = K" p ( d ) i d ' = K" ( d ) − K" s 2
]( d + c )
2
(6)
for any distance. Three measurement groups at three distances can be used to determine the three unknowns, SK, K" S , and c. In actuality, the University of Wisconsin (UW) ADCL uses this technique with seven distances of 10, 15, 20, 25, 30, 35, and 40 cm. The seven distances above redundantly determine the scatter and error in distance since there are 3 unknowns with 35 solutions (DeWerd, DeWerd, and Attix 1993). A computer-generated solution then can be used to average the solutions. Thus, finally, the value of SK is determined without scatter and without distance error. This value then can be used by the ADCL to calibrate well chambers for clinical use. Beta Sources for Brachytherapy For beta sources, NIST uses an extrapolation chamber to measure either a single source or the ophthalmic applicator. In both cases, the chamber volume is extrapolated to a theoretical zero gap to measure the dose at the window surface. For intravascular sources, the absorbed-dose-to-water is specified at 2 mm. For ophthalmic applicators, it is specified as a surface dose. The extrapolation chamber is a primary standard for the determination of absorbed-dose-to-water at 2 mm for beta-emitting sources. Constructional details and operational performance of extrapolation chambers are given in Soares et al. (1998). The extrapolation chamber is basically an air-filled, plane-parallel chamber where the distance between the high-voltage and collecting electrodes (air gap) can be varied. The absorbed dose rate is determined from current measurements at a number of precise air gaps. The current values as a function of air gap are used to determine the slope of the data at the limit of zero air gap. The absorbed-dose-rate-in-water is then given by the Bragg-Gray relationship
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Dw =
water (W / e) ⋅ Sair ( ∆I / ∆$)$→0 kback , ρ0 a
(7)
where (W/e) is the average energy in joules needed to produce one coulomb of charge of either sign in water is the ratio of the mean mass collision stopping power of water to that dry air (33.97 ± 0.05 JC–1), Sair of air, ρ0 is the density of air at the reference temperature and pressure (T0, p0), a is the area of the collecting electrode, ( ∆I / ∆$)$→0 is the rate of change of corrected current (normalized to a reference temperature and pressure) with extrapolation chamber air-gap thickness as the thickness approaches zero, and kback is a correction factor that accounts for the difference in backscatter from the collecting electrode compared to that of water. The area, a [cm2], of the collecting electrode used is critical. The dose rate is determined from the measured currents within this area. The measured dose rate is averaged over this area. The area of the collecting electrode must be smaller than the radiation field being measured, so that the collecting electrode determines the measurement area rather than the radiation field area. The extrapolation chamber has been used at NIST to determine reference absorbed dose rate from a beta-emitting seed or a line of seeds (Soares, Halpern, and Wang 1998). For these measurements the source is inserted in a hole in a tissue-equivalent plastic block with the center of the source at a distance of 2 mm from the surface. At this depth, the radiation field from a seed or line source is such that a collecting electrode diameter of 1 mm can be used to measure absorbed-dose rate. There are problems with this method, mainly due to uncertainties in the effective collecting area of the extrapolation chamber and the divergence effect of the small source/collector geometry. This results in an uncertainty of ±15% (k=2). Taylor and Kuyatt (1994) explain evaluating uncertainties and coverage factors. When corrections are made for the geometric effect and smaller gaps are used, the calibration uncertainty has been decreased to ±8% (k=2).
ADCL Standards The primary air kerma strength calibrations for all brachytherapy sources has been transferred from NIST to the ADCLs. The HDR 192Ir source calibration is determined by the ADCLs as mentioned previously in the section HDR Sources, and is traceable to NIST through external beam calibrations of thimble or spherical chambers. The absorbed-dose-to-water calibrations performed by NIST via the extrapolation chamber have also been transferred to the ADCLs. The ADCLs maintain all these standards to a very high precision. These primary measurements are then transferred to the clinic via the ADCL through use of a reference class well chamber. The calibration precision is demonstrated by proficiency tests with NIST or round robins with each other. For all photon brachytherapy sources, all ADCLs are within ±0.6% of NIST during proficiency tests. For beta sources, agreement among the ADCLs and NIST is ±2%. The UW ADCL has done proficiency tests with NIST for ophthalmic applicators and was within 5%. In all cases, these proficiency tests fall within the stated criterion. In addition to the standards maintained for brachytherapy sources, which are then transferred to well chambers, the ADCLs calibrate electrometers also. Intercomparisons between ADCLs for electrometers are always within ±0.2%.
ADCL Source Calibrations and Ophthalmic Applicators The ADCLs will calibrate sources in addition to well chambers. This involves having a NIST-traceable calibration coefficient on a reference well chamber, since many sources have a relatively short half-life. The source to be calibrated is then inserted in the well chamber, the signal measured, and the air kerma strength determined. Because of the short half-life of many sources, it is generally more useful for the clinic to have a calibration on their well chamber instead of having sources calibrated.
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Ophthalmic applicators generally use 90Sr, which has a half-life of 28 years. The UW ADCL has a NIST-calibrated source, which is then used with the source to be calibrated. Both sources are placed on radiochromic film at the same time for various exposures. This allows the surface dose on the film to be registered. Net optical density versus dose can then be plotted for the NIST-calibrated source, and the dose in the central 4 mm of the applicator can be determined. In addition, this allows an isodose plot to be made from the contact autoradiograph. The calibration of these sources is about 30% or more different from the prior calibrations done before 1991 (Soares 1991). In addition, the exposure surface is not uniform. For example, Figure 3 shows the autoradiograph of a uniform source. A dose profile in the two orthogonal directions also illustrates its uniformity. Many applicators are not as uniform as shown in Figure 3. For example, Figure 4 shows the autoradiograph of a very non-uniform source. The dose profile in the two orthogonal directions shows the dip in the center, which is shorter in the vertical direction than in the horizontal direction. Note the offset or shift of the dose-weighted isocenter from the physical source center and the non-uniform dose distribution. In some cases, ophthalmic applicators have to be rejected from calibration because of extreme dose profiles or damage of the sources.
Characteristics of Well Chambers for Brachytherapy Sources Well ionization chambers are a common piece of equipment for measurement of brachytherapy source strength. It is highly recommended that the well chamber be dedicated to brachytherapy uses and not be shared with nuclear medicine. The major reasons for this are twofold. First, when the chamber is not under control of the therapy medical physicist, “situations” can develop such as contamination from an unsealed liquid nuclear medicine source. Second, the brachytherapy source calibration is highly dependent on the insert used and subsequent geometry. Calibration at an ADCL is done with a specific insert and it is part
Figure 3. Contour plot of a uniform 90Sr ophthalmic applicator.
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Figure 4. Isodose contour plot of a non-uniform 90Sr ophthalmic applicator.
of the calibration coefficient. The general characteristics for well chambers are given below. The medical physicist should be aware of the equipment used and characteristics that need to be kept in mind. Many well chambers have a “sweet spot” where the output is highest for a given source. Figure 5 shows a typical example for a photon source, in this case an HDR 192Ir source. From this figure it can be noted that a longer source, e.g., a 60-mm train, would result in a 6% decrease in signal at the ends. This falloff is the problem with measurement of source trains because of the axial geometry limitations of the sweet spot of well chambers. Therefore, many manufacturers have developed new chambers or modified existing ones to cover longer sources. These chambers have an extended region of uniform axial response, or a sweet length. An example of such a chamber is shown in Figure 6. As part of the calibration, the ADCL provides a determination of the ion collection efficiency for well chambers. Ion collection efficiency, Aion, is a measurement of the percentage of charge measured by the chamber versus the total charge, which is a measure of the ionization recombination of the chamber. To know the dose absorbed by the air in the ionization chamber requires collecting all the ions formed by the radiation. For many well chambers for these sources, Aion is equal to 1.000 and so can be ignored. However, Aion may not be unity for HDR sources. Well chambers can be open to the atmosphere or pressurized. If they are open to the atmosphere, correction for air density should be made. If they are pressurized, a quality control measurement with time (quarterly) should be made to check for loss of pressure caused by leakage. Air density corrections (temperature and pressure) are calculated according to equation (8). , kTp =
(273.15 + T ) 101.3 (273.15 + T ) 760 • • = (273.15 + T ) p (273.15 + T ) P 0 0
(8)
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where T is the temperature in °C, p or P is the pressure in kPa or mm Hg, respectively, and T0 is the reference temperature at time of calibration (22.0 °C in the United States). As with all ionization chambers, periodic consistency tests with a long-lived source (137Cs for example, or an external 60Co beam) should be part of the routine QA program. An acceptable method for these checks in addition to those mentioned above is a redundancy check or intercomparison with one or two other chambers on the same source in the same time period. In particular, having two or three chambers that have been calibrated at an ADCL, which then would be used to measure the same source within a short time period is acceptable. If necessary, account should be made for decay of the source. Either the same electrometer for each chamber or independent dosimetry systems can be used for this exercise
Figure 5. Normalized response of a well chamber along its axial length.
Figure 6. An example of a well-chamber response from the bottom of the insert showing the axial sweet length for 192Ir.
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applying the calibration coefficients for each. The chambers or the systems, whether at a given center or from a neighboring center, should be used at the same time and a ratio of results kept. This method relies upon the stability of the systems with respect to each other. This method is suggested in the AAPM TG-56 document (Nath et al., 1997). In addition, the concept of redundant measurements has been proven for external beam chambers in the paper by Rozenfeld and Jette (1984). Such tests not only verify the calibration of the chamber, but also check that sealed chambers maintain their pressure. The response of the chamber should remain constant to within ±0.5% (Goetsch et al., 1992). The ADCL also calibrates electrometers for use with the well chambers. Both the chamber and electrometer correction coefficients must be used to achieve correct results. The electrometer and the well chamber then operate as a system. The scale of the electrometer should be low enough so that good precision and an adequate signal-to-noise are achievable. This is especially crucial for LDR sources where the noise can be a significant amount of the signal for some electrometers. For this reason many manufacturers have made electrometers with pA scales.
Well-Chamber Characteristics for Photon Sources Note in each case, the appropriate source holder must be used since it is part of the calibration. For HDR sources, the ion recombination should be checked, since not all chambers have full collection efficiency. Recombination correction, krecom, may be determined using the two-voltage technique for gamma-emitting sources. If the ratio of the voltage used in this technique is exactly 2 (e.g., if 150V and 300V are used as is often the case with well type chambers) then the recombination correction can be determined from (Attix 1986) 1 krecom
=
4 1 Q300 . − 3 3 Q150
(9)
The ADCL has determined Aion previously upon calibration; the user will have to determine for the source on site. Good-quality chambers generally exhibit negligible recombination effect for gamma-emitting brachytherapy sources. Many well-type chambers have proven their stability with time. Evaluations of commercially available well chambers show a long-term reproducibility of 0.5% or better. If during the constancy checks of the well type chamber it is observed that its response changes more than ±1%, a recalibration is recommended.
Well-Chamber Characteristics for Beta Sources The beta sources presently used for intravascular brachytherapy are 90Sr-90Y sources. Much of the well chamber characteristics were reported on in the previous summer school of 2002 (DeWerd 2002). The source trains are sent to the end of the catheter by pressure from de-ionized water that runs along the inner lumen, past the pellets, and back through the outer lumen. It is important that the catheter be purged with water prior to each measurement. Void regions in the catheter can produce significant variation. Measurements of each source should be done at 90° rotations in the well chamber to average the positioning of the source and the lumen. In this case, the insert used in the chamber is a major part of the calibration, since it has a direct effect on the chamber system. For calibration of long intravascular brachytherapy source trains, the response curve or sweet length (variation of chamber output with position within the well chamber) must be of sufficient length to accommodate the source. DeWerd (2002) elaborates on the measurements made for various chambers. The sweet lengths of chambers for these source trains extend from 50 mm to over 100 mm.
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For a well chamber to be useful in measuring a beta source, it is necessary to have accurate calibration coefficients by which the dose rate of the source can be determined from the signal generated in the well chamber. Since this source is beta emitting, a number of questions were involved in the calibration of these chambers in addition to the axial response or sweet length. DeWerd (2002) addressed four areas of concern, the first being the axial response. In addition, the type of radiation best measured, e.g., create bremsstrahlung radiation or measure beta particles; catheter uniformity for calibration; and calibrations for the various lengths of seed trains were also addressed. For the type of radiation measured it was found that the best compromise with respect to signal-to-noise ratio, rotational consistency and accuracy were obtained by filtering out low-energy beta particles without the creation of additional bremsstrahlung. For these reasons a 2-mm acrylic holder was chosen for use with the well chambers. A study was also done for any variations among catheters. The results showed that the calibration coefficient for various catheters was the same within ±0.1%. While it is imperative to match the clinically relevant catheter to the specific source train length, there appears to be no concern regarding use of multiple catheters for well chamber measurement of Novoste source trains. Finally, an investigation was conducted for calibrations for the various lengths of seed trains. Each Novoste train was inserted within the sweet length of the chamber and measured through a rotation of 360°, in 90° steps, for a total of five measurements of each train. The relationships between the 30 mm, 40 mm, and 60 mm trains were found to be in proportion to their length or number of seeds. However, some chambers tested were up to 4% different from the expected ratio. DeWerd (2002) showed that the various train calibration coefficients are related in proportion to the length of the train. The ADCL will thus calibrate one seed train length for a chamber that follows the relationship as given in equations (10). D40 = 0.75D30 (10) D60 = 0.50D30 where D30 is the 30 mm calibration coefficient, D40 is the 40 mm absorbed dose-to-water calibration coefficient, and D60 is the 60 mm absorbed dose-to-water calibration coefficient.
Clinical Calibration Using Well Ionization Chambers Well-type chambers provide a reliable method for calibrating brachytherapy sources before clinical use and are recommended for use for both brachytherapy photon sources, and beta sources. The preferred method for traceability in the source calibrations is to have the well-type chamber calibrated against the primary standard at an ADCL. This calibration should be carried out for each source model to be used with the appropriate source holder made by the manufacturer for that source. The source holder is part of the calibration. The accuracy of the source calibration limits the final accuracy of the dose calculation for the patient in the clinic. Therefore, the clinic needs to determine the air kerma strength, SK, for photon sources or the absorbed-dose-to-water for beta sources. In the clinical situation this calibration is accomplished with a calibrated well chamber. The user’s well chamber should come to equilibrium with its surroundings before beginning clinical calibration. The minimal time necessary for this is generally 30 minutes. Care should be taken that the temperature measured is that of the chamber volume and not the room temperature. A minimum of three
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significant digits should always be obtained for charge accumulated or current measurements. Thus charge should be accumulated for a set time depending on the activity of the source. A minimum of three measurements for each source insertion that are neither monotonically increasing nor decreasing should be obtained.
Clinical Calibration of Photon Sources When using chambers for calibrations, the chamber should be located in a “scatter free” location (Podgorsak et al., 1992) .The calibration coefficient of the chamber assumes that the source will occupy the same position as during calibration by the ADCL. This position is usually that of peak response or the center of the uniform response plateau. The user should verify the position before proceeding. See the section, Well-Chamber Characteristics for Beta Sources, and Figure 5 for further information on the sweet spot. The calibration coefficient for the chamber gives air kerma strength per unit reading. Measurements for the HDR source generally use the current mode whereas for LDR sources, it may be either charge or current. For the current measurements, SK = I • Ne • kTP • NSK • Aion • Pion ,
(11)
where: SK is the source air kerma strength, I is the current measured from the well chamber, Ne is the correction coefficient for the electrometer scale, kTP is the correction for temperature and pressure [see equation (8)]; this factor is not used for pressurized chambers, NSK is the ADCL provided air kerma strength calibration coefficient for the well chamber, Aion is the correction for collection efficiency at the time of calibration, from the ADCL calibration report, and Pion is the correction for collection efficiency at the time of measurement. Note that equation (11) applies to LDR photon-emitting sources as well, with two differences. The first involves Aion and Pion, which are generally equal to 1.000 and thus ignored. The second is kTP, which is given by equation (8). Recently, deviations from the standard air density correction have been found for low-energy, LDR photon sources (Griffin et al., 2005; Bohm et al. 2005). These deviations are found at various pressures that enter at different altitudes. There is an additional correction, PA that can be applied after applying the standard kTP given by the relation: PA = k1 ( P ) 2, k
(12)
where P is the pressure and the constants, k1 and k2, relate to the type of source, 103Pd, 125I with a silver base or without a silver base, and the type of well chamber. These constants are given in Griffin et al. (2005) and reproduced in Table 3. In addition, for LDR sources, the signal-to-noise and the accompanying selection of an electrometer may be crucial; see Characteristics of Well Chambers for Brachytherapy Sources.
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Table 3. Constants k1 and k2 for three chambers and three types of seeds Type of LDR source
HDR 1000 and IVB 1000
PRM WC 2 chamber
k1
k2
k1
k2
Pd
0.0241
0.562
0.171
0.266
I with silver base
0.0490
0.455
0.082
0.378
I without silver base
0.0573
0.431
0.069
0.403
103 125 125
Clinical Calibration of Beta Sources For beta-emitting sources, the calibration procedure is much the same as above for photon sources; however, it is of much greater importance that close attention is paid to the source holder and setup of the chamber. The sweet length and source insert are very important as described previously in the section, Well-Chamber Characteristics for Beta Sources. When performing clinical measurements of the beta sources, make sure the correct settings are used on the insert for the appropriate length source. Measurements should be made at various orientations of the source about its cylindrical axis and the results averaged. Multiple insertions of the sources should also be made for loose seed trains delivered in a catheter. Since there is a polarity effect for beta particles, the calibration coefficient is only valid for the polarity stated in the calibration certificate. The same polarity as used in the ADCL calibration should be used in the clinic. The ion collection efficiency measurements are ignored for beta-emitting source measurements. The correction factor below is for a given train length, e.g., 30 mm. D2mm = Mu • Ne • N2mm • kTp ,
(13)
where the quantities are the same as equation (11) except that D2mm is the absorbed-dose-in-water at 2 mm and N2mm is the calibration coefficient for the beta source provided by the ADCL. Again, if a pressurized chamber is used here, kTP is not used. However, there is a 1% effect at lower air density for beta sources for pressurized chambers (Griffin et al., 2005).
Conclusion The use of brachytherapy sources for brachytherapy is increasing. It is important that the patient receive the maximum benefit from this technique by good physics practices. Critical to this application is the calibration of the source. The present procedures have increased the dosimetric accuracy greatly for the clinic. Well chambers have simplified the measurement problems at the clinical level. This chapter has demonstrated that calibration of sources traceable to national standards, is important.
References American Association of Physicists in Medicine (AAPM). Report No. 21. “Specification of Brachytherapy Source Strength.” New York: American Institute of Physics, New York, 1987. Attix, F. H. Introduction to Radiological Physics and Radiation Dosimetry. New York: John Wiley and Sons, 1986. Bohm, T. D., S. L Griffin, P. M. DeLuca, and L. A. DeWerd. (2005). “The effect of ambient pressure on well chamber response: Monte Carlo calculated results for the HDR 1000 Plus.” Med Phys 32:1103–1114.
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Collé, R. (1999). “Chemical digestion and radionuclide assay of TiNi-encapsulated 32P intravascular brachytherapy sources.” Appl Radiat Isot 50:811–833. DeWerd, L. A. Brachytherapy Dosimetric Assessment: Source Calibration, Radiological Society of North America, Categorical Course in Brachytherapy Physics. B. Thomadsen (ed.). Oak Brook, IL: RSNA, 1997. DeWerd, L. A. “Source Standardization and Calibration for Intravascular Brachytherapy” in Intravascular Brachytherapy/Fluoroscopically Guided Interventions. S. Balter, R. C. Chan, and T. B. Shope, Jr. (eds.). American Association of Physicists in Medicine 2002 Summer School proceedings. Medical Physics Monograph 28. Madison, WI: Medical Physics Publishing, pp. 423–443, 2002. DeWerd, L. A., and B. R. Thomadsen. “Source Strength Standards and Calibration of HDR/PDR Sources” in Brachytherapy Physics. J. F. Williamson, B. R. Thomadsen, and R. Nath (eds.). American Association of Physicists in Medicine (AAPM) 1994 Summer School. Madison, WI: Medical Physics Publishing, 1995. DeWerd, L. A., S. M. DeWerd, and F. H. Attix. (1993). Solution to Inverse Square Equations Involving Distance Error and Scatter Correction. Wisconsin Medical Physics Report #197 Madison; WI. (Available from the University of Wisconsin.) DeWerd, L. A., M. S. Huq, I. J. Das, G. S. Ibbott, W. F. Hanson, T. W. Slowey, J. F. Williamson, and B. M. Coursey. (2004). “Procedures for establishing and maintaining consistent air-kerma strength standards for low-energy, photon-emitting brachytherapy sources: Recommendations of the Calibration Laboratory Accreditation Subcommittee of the American Association of Physicists in Medicine.” Med Phys 31:675–681. Douysset, G., J. Gouriou, F. Delaunay, L. DeWerd, K. Stump, and J. Micka. (2005). “Comparison of dosimetric standards of USA and France for HDR brachytherapy.” Phys Med Biol 50:1961–1978. Dove, D. B. (1959). “Effect of dosemeter size on measurements close to a radioactive source.” Br J Radiol 62:202–204. Glasgow, G. P., J. D. Bourland, P. W. Grisby, J. A. Meli, and K. A. Weaver. “Remote Afterloading Technology.” AAPM Report No. 41. New York: American Institute of Physics, 1993. Goetsch, S. J., F. H. Attix, D. W. Pearson, and B. R. Thomadsen. (1991). “Calibration of 192Ir high dose-rate afterloading systems.” Med Phys 18:462–467. Goetsch, S. J., F. H. Attix, L. A. DeWerd, and B. R. Thomadsen. (1992). “A new well ionization chamber for the calibration of iridium-192 high dose rate sources.” Int J Radiat Oncol Biol Phys 24:167–170, Griffin, S. L., L. A. DeWerd, J. A. Micka, and T. D. Bohm. (2005). “The effect of ambient pressure on well chamber response: Experimental results with empirical correction factors.” Med Phys 32:700–709. International Atomic Energy Agency (IAEA) IAEA-TECDOC-1079. “Calibration of Brachytherapy Sources.” Vienna, Austria: IAEA, 1999. International Commission on Radiation Units and Measurements (ICRU). Report No. 33. “Radiation Quantities and Units.” Bethesda, MD: ICRU, 1980. International Commission on Radiation Units and Measurements (ICRU). Report No. 58. “Dose and Volume Specification for Reporting Interstitial Therapy.” Bethesda, MD: ICRU, 1997. Kondo, V. S., and M. L. Randolph. (1960). “Effect of finite size of ionization chambers on measurements of small photon sources.” Radiat Res 13:37–60, American Association of Physicists in Medicine (AAPM). (1994). “Comprehensive QA for radiation oncology. Report of AAPM Radiation Therapy Committee Task Group 40.” Med Phys 21:581–618. Also available as AAPM Report No. 46. Loftus, T. P. (1980). “Standardization of 192Ir gamma-ray sources in terms of exposure.” J Res Nat Bur Stand 85:19–25. Nath, R., L. L. Anderson, J. A. Meli, A. J. Olch, J. A. Stitt, and J. F. Williamson. (1997). “Code of practice for brachytherapy physics: Report of the AAPM Radiation Therapy Committee Task Group No. 56.” Med Phys 24(10):1557–1598. Also available as AAPM Report No. 59. Nath, R., H. Amols, C. Coffey, D. Duggan, S. Jani, Z. Li, M. Schell, C. Soares, J. Whiting, P. E. Cole, I Crocker, and R. Schwartz. (1999). “Intravascular brachytherapy physics: Report of the AAPM Radiation Therapy Committee Task Group No. 60.” Med Phys 26:119–152. Also available as AAPM Report No. 66. Podgorsak, M. B., L. A. DeWerd, B. R. Thomadsen, and B. R. Paliwal. (1992). “Thermal and scatter effects on the radiation sensitivity of well chambers used for high dose rate Ir-192 calibrations.” Med Phys 19:1311–1314.
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Rivard, M. J., B. M. Coursey, L. A. DeWerd, W. F. Hanson, M. S. Huq, G. S. Ibbott, M. G. Mitch, R. Nath, and J. F. Williamson. (2004). “Update of AAPM Task Group No. 43 report: A revised AAPM protocol for brachytherapy dose calculations.” Med Phys 31:633–674. Rivard M. J., B. L. Kirk, and L. C. Leal. (2005). “Impact of radionuclide physical distribution on brachytherapy dosimetry parameters.” Nucl Sci Engr 149:101–106. Rozenfeld, M., and D. Jette. (1984). “Quality assurance of radiation dosage: Usefulness of redundancy.” Radiology 150:241–244. Seltzer, S. M., P. J. Lamperti, R. Loevinger, M. G. Mitch, J. T. Weaver, and B. M. Coursey. (2003). “New national air-kerma-strength standards for 125I and 103Pd brachytherapy seeds.” J Res Nat. Instit Stand Tech 108:337–358. Soares, C. G. (1991). “Calibration of ophthalmic applicators at NIST: A revised approach.” Med Phys 18:787–793. Soares, C. G., D. G. Halpern, and C.K. Wang. (1998). “Calibration and characterization of beta-particle sources for intravascular brachytherapy.” Med Phys 25:339–346. Stump, K. E., L. A. DeWerd, J. A. Micka, and D. R. Anderson. (2002). “Calibration of new high dose rate 192Ir sources.” Med Phys 29:1483–1488. Taylor, B. N., and C. E. Kuyatt. (1994). “Guidelines for Evaluating and Expressing the Uncertainty of NIST Measurement Results.” NIST Technical Note 1297. Washington, DC: U.S. Dept. of Commerce, 1994. van Dijk, E., I. K. K. Kolkman-Deurloo, and P. M. G. Damen. (2004). “Determination of the reference air kerma rate for 192Ir brachytherapy sources and the related uncertainty.” Med Phys 31:2826–2833. Verhaegen, F., E. van Dijk, H. Thierens, A. Aalbers, and J. Seuntjens. (1992). “Calibration of low activity 192Ir brachytherapy sources in terms of reference air kerma rate with large volume spherical ionization chambers.” Phys Med Biol 37:2071–2082. Weaver, J. T., T. P. Loftus, and R. Loevinger. “NBS Measurement Services: Calibration of Gamma-Ray-Emitting Brachytherapy Sources.” National Bureau of Standards (NIST) Special Publication 250-19, Washington, DC: NIST, 1988. Williamson, J. F., and R. Nath. (1991). “Clinical implementation of AAPM Task Group 32 recommendations on brachytherapy source strength specifications.” Med Phys 18:439–448.
Chapter 12
Localization, Part I: Radiographic Methods and Accuracy Eugene P. Lief, Ph.D. Maimonides Comprehensive Cancer Center Brooklyn, New York Introduction: Requirements for Localization Accuracy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 173 Use of Various Imaging Modalities for the Source Localization . . . . . . . . . . . . . . . . . . . . . . . . . 174 Conventional Radiographs and Autoradiographs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 174 Orthogonal Techniques . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 174 Stereo-Shift Method . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 175 Automatic Reconstruction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 176 HDR Dwell Positions Localization Using Autoradiographs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 178 CT Scanning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 179 Catheters and Applicators . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 179 Prostate Seed Post-Implant Localization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 179 Automatic CT-based prostate Seed Detection . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 181 Electronic Portal Imaging Devices (EPIDs), Fluoroscopy, and Cone-Beam CT . . . . . . . . . . . . . . . 181 Source Localization for Various Brachytherapy Procedures . . . . . . . . . . . . . . . . . . . . . . . . . . . 182 LDR Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 182 HDR Brachytherapy Procedures . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 182 Intravascular Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 183 Other Methods of the Source Localization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 183 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 184
Introduction: Requirements for Localization Accuracy By definition, “brachytherapy” (“brachy” means “close” in Greek) is radiotherapy at small distances. Such distances (from zero to a few centimeters from the source of radiation) at present can be achieved utilizing small, sealed radioactive isotopes. Due to the inverse square law, which is valid for radiation from small (point) sources, a small inaccuracy in the distance measurements can result in significant difference in dose (Smith, Meigooni, and Chiu-Tsao 1990). For example, if we want to determine radiation dose at a distance of 1 cm from a small radioactive source, inaccuracy of 1 mm would result in a factor of (1.1 cm/1 cm)2 = 1.21 in dose calculation. That means more than 20% error in dose administration, which is a significant discrepancy for radiotherapy. When the dimensions of the sources are comparable with the distance to the tumor or the source radiation is not isotropic, there is an additional consideration of the source orientation, which could be established using plain radiographs (Lief, Davis, and Wallner 1998; Davis et al., 2003). Due to these considerations, source localization was one of the most important parts of brachytherapy. It can be achieved by using x-rays, ultrasound, magnetic resonance imaging (MRI), or other nonradiographic technique. In this chapter we will focus on radiographic localization, just briefly mentioning alternative approaches. Radiographic techniques can be further subdivided into plain radiographs, fluoroscopy, computed tomography (CT), and autoradiographs. By “plain radiographs” we mean taking radiographic films (or digital images obtained by electronic portal imaging devices (EPIDs) using external diagnostic x-ray source. Fluoroscopy utilizes continuous x-ray irradiation that is detected electronically, with the image being continuously updated on the screen. CT methods are based on performing a computer tomography scan and subsequently localizing (manually or automatically) radioactive or “dummy” sources as well as source applicators. One of the latest developments is utilization of a cone-beam CT feature of the Acuity simulator (Varian Medical Systems, Palo Alto, CA) for catheter localization in high dose rate
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(HDR) treatment of prostate (Prestidge 2004). Autoradiographs are performed by registering the source location on attached radiographic film, using radiation from the radioactive isotope instead of external source. Historically, plain radiographs with external x-ray source and autoradiographs were more commonly used when the CT was less available. Currently, the CT has become the most efficient imaging modality for prostate seeds post-implant localization and also for localization of HDR 192Ir applicators made of special CT-compatible materials. Autoradiographs are predominantly used for HDR 192Ir and intravascular source localization.
Use of Various Imaging Modalities for the Source Localization Conventional Radiographs and Autoradiographs Two or more plain radiographs provide the most accurate way of determining the source position as well as orientation. The most important tasks in the radiograph-based localization are matching the images of the same sources on all the films and correctly determining magnification factors of the films. Another challenge is to find the source location with respect to the tumor. Usually, the treatment target visualization can be performed using the CT easier than with the radiographs; however the CT is generally less convenient for the source localization. Orthogonal Techniques The simplest two-film technique is the one with the film planes orthogonal to each other. Typically, these are anterior-posterior (AP) and lateral films or two films taken at any directions orthogonal to each other. Sometimes these directions have to be different than AP/Lateral to avoid prostheses or other objects which are difficult to image. In such cases, the imaging could be done using paired left anterior oblique (LAO) and right anterior oblique (RAO) radiographs. To correctly determine the magnification factor, a magnification ring (or an opaque circle, e.g., a quarter coin) may be placed as close to the implant area as possible without occluding the view. The magnification factors for both films 1 and 2 in the vicinity of the ring M1(r) and M2(r) can be calculated then as a maximum distance between any two points of the ring on the images 1 and 2, respectively, divided by the actual ring diameter. Usually it is reasonable to assign the same magnification factor to all the seeds on the image, as the size of the implanted domain is usually much smaller than the distance from the radioactive sources to the x-ray tube. Nevertheless, since the location of the ring is typically different from the implanted domain, film magnification factors for the seeds are generally different from M1(r) and M2(r) and could be found using the following formulas (Smith, Meigooni and Chiu-Tsao 1990):
M1 =
(F F + x 1 2
(r )
1
)
F1 F2 + x1 z2
(
(1)
)
z2 / M 1( r ) + F2 z2( r ) − z2 / M 2( r )
and similar for M2 (see Figure 1): M2 =
F1 F2 + x1 z2
(F F + x z ) / M (r )
1 2
1 2
(r ) 2
(
)
− F1 x1( r ) − x1 / M 1( r )
,
(2)
where F1 and F2 are source to film distances for beams 1 and 2, respectively, and x1 and z2 are projections of a point deviation from the isocenter on the planes of the films. F1 and F2 are negative if the
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Figure 1. Radioactive source localization using two orthogonal radiographs.
corresponding beam is directed opposite the coordinate axis. In equations (1) and (2), x1(r) and z2(r) refer to x- and z-coordinates of the magnification ring. Using M1 and M2, one can find the actual spatial coordinates of each source. The advantage of the orthogonal method is its simplicity and accuracy. Film technique offers sharp images of the seeds implanted. This is especially noticeable for AP films with better contrast than lateral films. Using films, each coordinate of nonoverlapping seeds can be typically digitized with the accuracy of 0.5 mm. The disadvantage of the method is increasing difficulty of the seed matching with the increase of their quantity. The important factor in seed matching is identity of y-coordinates of the same seed on both projections. Other difficulties include: 1. Possible patient movement between the exposures: if the patient moves between the exposures, the source matching would be much more difficult and much less accurate, and 2. Inability to distinguish different possible coplanar seed locations: an example of the uncertainty of the seed locations would be the registration of two nonexisting “sources,” shown as open circles in Figure 1, instead of two actual seeds (S and I) shown in solid. Stereo-Shift Method As mentioned above, the orthogonal technique becomes inconvenient if there is a large number of sources implanted and especially if many of them lie in the same plane. Another technique, called “stereo-shift,” could be utilized in this case (Smith, Meigooni, and Chiu-Tsao 1990). The idea of this method is to take two images from slightly different positions. It is usually achieved by moving the x-ray tube with respect
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to the patient between the two exposures and using a fiducial marker attached to the patient at a predefined location near the implant. The magnitude of the shift should be small enough with respect to the source-to-seeds distance, so the radiographs would look similar and make seed matching easier. On the other hand, the distance of the shift should be significantly larger than the difference in heights of the seeds and the fiducial marker to make the method more accurate. A typical shift is about 20 cm. Figure 2 shows the scheme of the z-coordinate determination for a seed moving with the patient from location S1 to S2, with respect to the x-ray tube. The idea of the stereoshift method is utilizing the fact that as a small object (a radioactive source S or a fiducial marker F) is moved by a fixed distance s with respect to the x-ray source, its “shadow” on the film will move by a distance x which depends upon the object’s height z. The shift s, applied to the source, moves it from position S1 to S2, while the fiducial marker after the same shift moves from F1 to F2. These identical shifts change the distance between x-ray images of the objects S and F on the film from x1 to x2 (Figure 2). Using the magnitude s of the stereoshift and a height of the fiducial marker zf, one can arrive at the following expression for the unknown height of the seed zs (Smith, Meigooni, and Chiu-Tsao 1990):
zs = SFD ⋅
( SFD − z ) ⋅ ( x − x ) − s ⋅ z , ( SFD − z ) ⋅ ( x − x ) − s ⋅ SFD 2
f
f
1
2
f
(3)
1
where SFD is the source-to-film distance. The magnification factor M for each seed can be found, using a formula:
M =
SFD − zs SFD
.
(4)
The advantage of the stereo-shift method is ease of seed identification. However, the accuracy of the stereo-shift method is lower than for the orthogonal films, since a small error in the source coordinate measured from a film results in a parallax with a much larger error in the seed height. For example, if the magnitude of the stereo-shift s = 20 cm, SFD = 140 cm, and zs = 40 cm, then an error ∆x = 0.5 mm in seed digitization translates into a larger error ∆zs in the seed z-coordinate calculation: ∆ zs ≈
∆x ⋅ ( SFD − zs )
s SFD ⋅ tan SFD − zs
≈ 1.8 mm .
(5)
In this formula, the value of the angle in tangent function is expressed in radians. The error ∆y in ycoordinate found from seed digitization is not magnified because of the parallax but simply scaled by the magnification factor. Numbers chosen for this example can vary depending on the setup. Automatic Reconstruction There are a number of various approaches to automatic reconstruction using plain radiographs (Amols and Rosen 1981; Altschuler, Findlay, and Epperson 1983; Biggs and Kelley 1983; Rosenthal and Nath 1983; Stock et al., 1995; Todor et al., 2002). Using the same idea of imaging the implant from two or more different angles, one can use the couch translation or the gantry rotation to obtain those images.
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Figure 2. Scheme of the source S localization using stereoshift s and fiducial m ark F.
By matching the source images on the radiographs, measuring their coordinates with respect to the center of the images, and entering the coordinates in the computer, one can reconstruct the spatial distribution of the sources. One of the methods is based on two conventional radiographs; one is typically in AP direction, and the other, at a specific angle (usually, 15° to 45°) with respect to the first one. The two-film method requires identifying seeds on the first film and finding corresponding seeds on the second film. The problem becomes more difficult with an increasing number of seeds. It also becomes more tedious to correlate seed images on both films for larger angle between the gantry positions at which the films were taken. On the other hand, the reduction of this angle significantly reduces the accuracy of spatial seed localization in the AP direction. The compromise is usually achieved at an angle of about 20°. For the two-film method, as for the stereo-shift method above, the largest uncertainty in coordinate determination is in the z-direction. Similar to equation (5), one can find that an error ∆x = 0.5 mm in the seed digitization translates into a larger error ∆zs in the seed z-coordinate calculation:
∆ zs ≈
∆x ⋅ ( SFD − zs ) SFD ⋅ sin ( 20° )
≈ 1.0 mm .
(6)
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In this formula, the value of the angle in sine function is expressed in degrees. As in equation (5), the numbers are chosen just for the illustration. They will vary depending on the geometry of each particular case. As mentioned in the orthogonal method description, two-film methods can result in registration of nonexisting seeds instead of the real ones. One way to eliminate this is to use the third radiograph (Amols and Rosen 1981; Rosenthal and Nath 1983). In this approach, there is an AP film and two other films symmetrically taken at the gantry angles ±α, where the angle α could range from 8° to 45°. The 45° approach allows more accurate determination of the z-coordinates of the sources. To make matching easier, one could compare y-coordinates of the seeds. For the same seed, these coordinates should be identical on all three images. If the accuracy of the seed digitization in the three-film method is limited, two seeds with close z-coordinates could be interchanged. This and other difficulties in seed matching can leave a certain number of sources on both films unmatched. In this case, the missing activity can be reassigned among the matching sources. It has been suggested that up to 5% of seeds could be mismatched for the dose calculations to be performed with the acceptable accuracy (Rosen et al., 1982). HDR Dwell Positions Localization using Autoradiographs An important example of brachytherapy source localization is determination of actual coordinates of HDR Ir source dwell positions. Usually, the consistency of the dwell positions is checked using a plain endobronchial catheter of a known length. This can be done by attaching a 15 to 30 cm length of catheter at the closed end to a sheet of a plain ready-pack film, using masking tape. Ready-pack “XV” or “EDR”type films (Eastman Kodak Company, Rochester, NY), which contains a single sheet of film in a lightproof paper envelope, can be used for this purpose. Some other types of film with higher speed, “XTL” or “PPL,” are too fast for the HDR exposures, and the autoradiograph would have excessive blurring from the source motion. If there is a noticeable motion of the film sheet within the paper envelope, the film can be stapled through the envelope at two corners 5 cm or more from the catheter. First, a metal wire with non-active x-ray opaque, nonradioactive markers 10 mm apart (“dummy wire”) is inserted in the empty catheter. The markers are indicating intended dwell positions of the HDR source. The wire has to go all the way into the catheter and be securely fastened by a special cup at the open end. The film with the wire located in the taped empty catheter is exposed to 80 to 120 kVp diagnostic x-rays. The side of the envelope with the catheter should be closer to the x-ray tube. This exposure leaves on the film (after processing) sharp light spots at the markers’ locations. After the first exposure, the catheter remains taped to the film envelope, while the “dummy wire” is removed. The empty catheter is connected to the HDR, and the source is preprogrammed to stop at the positions corresponding to the locations of the “dummy” markers. The dwell time at each position depends on the source strength and the film speed. During the HDR run, the film receives high doses from the source at the dwell locations. After the HDR exposure, the film is processed. The developed film exhibits sharp light spots from the markers overlaid with the dark blurry spots from the HDR source. For the analysis, one should find centers of the dark spots and mark them on the film. Centers of light and dark spots should ideally coincide. The distance between the centers shows the magnitude of the misalignment. Another method, based on autoradiography, allows checking of the HDR dwell positions using only one HDR exposure (Anderson et al., 1995). The idea is to compare the actual dwell positions of the HDR source with predefined stationary markers associated with the empty catheter. A special plastic phantom in the form of a slab has an endobronchial catheter permanently attached to its flat external surface. Several thin metal wires are embedded in the plastic, perpendicular to the catheter, at a small distance from it. The wires are located at regular 1-cm intervals, at the expected HDR dwell positions. The phantom has a thin slot for a film parallel to the catheter and located at a small distance from it. During the HDR exposure the dwell positions are preprogrammed to coincide with the wire locations. The HDR source irradiates the film and the wires simultaneously. Direct film irradiation results in dark
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blurry spots on the film, like in the previous method. Irradiation of wires generates a stream of photoelectrons, which leaves sharp dark marks on the film. After the HDR run is finished, the film is pulled out of the phantom and processed. During the analysis, one finds centers of each dark spot indicating the center of the HDR source at the dwell positions. In this method, the center of the dark spot should coincide with the dark mark from the wire. If the HDR alignment tests are correctly performed, one could expect an accuracy of 1 mm in the source dwell positions determination. Autoradiographs can be also used for the HDR source localization in some other catheters and applicators (Jones 1989). For Nucletron applicators (Nucletron B.V., Veenendaal, The Netherlands), ready-pack X-OMAT “V” film from Kodak can be used with lead foil as a marker. Radiation-sensitive yellow Detex paper (Detex Corporation, New Braunfels, TX) instead of ready-pack film (Jones 1989). For “V”-type film used with a 370 GBq (10 Ci) 192Ir source, it is recommended to use the following dwell times: bronchus catheter and implant needle = 0.1 s, uterine tube = 0.3 s, esophageal applicator = 2 s. Autoradiographs can be also performed using GAFchromic film (International Specialty Products, Wayne, NJ) (Chiu-Tsao et al., 2004a,b). This film has several advantages: it is not light-sensitive, does not require processing, and has lower speed, which enables longer dwell times and reduces the blurring of the autoradiograph due to the exposure from the moving source.
CT Scanning Catheters and Applicators CT becomes a major imaging modality not only in teletherapy, but also in brachytherapy. Its main advantage is the ability to obtain in a matter of minutes all necessary information about the locations, form, and orientations of the “dummy” sources, catheters, and applicators. The disadvantages of the CT are: lower spatial resolution of small sources and difficulty in reconstruction in the presence of large metal objects, like applicators or even clusters of seeds. Large seed clusters, which sometimes appear after permanent prostate implants, and regular metal applicators can create star artifacts in the form of multiple straight lines intersecting at the metal object location. Such CT-image distortions make it difficult to localize the applicator and invalidate the Hounsfield numbers and dose computation in the vicinity of the metal objects. To use the CT as an imaging modality in the vicinity of small metal sources or applicators, special reconstruction algorithms should be employed to minimize the artifacts. These algorithms make it possible to image small metal seeds used for the permanent prostate implant, but bulky metal objects would still give significant artifacts. HDR 192Ir vendors had to provide special CT-compatible applicators made of solid materials with low diagnostic x-ray attenuation. Modern brachytherapy treatment-planning systems are capable of catheter reconstruction and threedimensional (3-D) dose calculation based on the CT images. The CT image set is imported into the treatment-planning system and associated with the patient. Manually digitizing “dummy” seed markers with a computer mouse, one can reconstruct the form, location, and orientation of a catheter or an applicator (Figure 3). Alternatively, some brachytherapy planning systems can reconstruct the sources, the catheters, and the applicators automatically, based on predefined Hounsfield numbers. Prostate Seed Post-Implant Localization Seed localization after the permanent prostate implant is an important application of CT-scanning in brachytherapy (Roy et al., 1993). It allows for quickly collecting all the information of the seed locations and orientations. Seeds appear as distinct white spots on the CT-images (Figure 4); however, there are two major challenges in the CT-based localization. First, some seeds in the same slice are located so close to each other that they merge into a single white spot in the image. Sometimes, it is possible to distinguish visually the “double seed” from a “single” one judging by the size of the spot; however, it is not always
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Figure 3. Reconstruction of HDR catheters and patient’s anatomy by VariSource HDR planning software package from Varian Medical Systems, Inc. (Palo Alto, CA). [Courtesy of Anil Sharma, Ph.D., Long Beach Memorial Medical Center.]
Figure 4. Automatic seed finder module for a commercial implant planning software package “VariSeed v.7.1” from Varian Medical Systems, Inc. (Palo Alto, CA). Different sizes of light spots, representing seeds, make it difficult to distinguish visually between a single source and several sources clustered together, as well as sources that fully or partially lie in a given CT-slice. Note also small reconstruction artifacts in the form of lines connecting different sources. [Courtesy of Ted Jackson, Ph. D., Varian Medical Systems, Inc.]
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obvious. There is also a possibility of getting three or more seeds together, but it is much less probable. One of the ways of handling such clusterization is to assign unaccounted seeds to larger spots, which are likely to contain more than one. The second problem is that some seeds appear on both adjacent contiguous slices as two full-size spots. It happens because a part of such a seed lies in one slice, and the other part in the other. For 4.5-mm sources this is especially noticeable if the slice thickness is 3.75 mm or smaller. For instance, if 3-mm helical contiguous scanning is used, the double-registration could be as high as 40% (Lief, Davis, and Wallner 1998). This number could be slightly reduced by eliminating white spots of obviously smaller size that result from the small protruding of a seed into a CT-slice and partial volume effect. Another way to recognize the double-registration is to correlate the locations of the white spots in adjacent images and eliminating one of those that coincide. Unfortunately, this does not work all the time either because most seeds implanted into the prostate are tilted by a certain angle with respect to the CT-axis. In this case, an oblique seed protruding into both adjacent slices can register at two different locations there. One more way to reduce the double-registration is to use the appropriate CT-slice thickness and spacing (Lief, Davis, and Wallner 1998). With all these measures taken, the CT-scanning becomes a powerful tool for the post-procedure localization in permanent prostate implants. CT is the imaging modality recommended by the American Brachytherapy Society (ABS) (Nag et al., 2000) for post-implant dosimetric analysis. Automatic CT-based Prostate Seed Detection There have been many approaches to automatic CT-based seed localization (Brinkmann and Kline 1998; Li et al., 2001; Liu et al., 2003; Holupka et al., 2004). The main purposes of automatic localization are: to save time, to determine seed location more accurately than using the manual methods, and exclude interand intra-observer variability. The idea of the threshold method (Brinkmann and Kline 1998) was to search for 3-D connected objects in the thresholded CT-volume by rastering through the volume voxel by voxel. The method was tested on a special phantom with the seeds implanted according to a typical plan. Using the software written by the authors, about half of all seeds were detected with 0.5 mm accuracy, while for the other half the error was within 2 mm. In a multiple threshold approach (Li et al., 2001), the computer program enabled a graphic user interface allowing varying multiple thresholds for better performance. The purpose of the parameter variation was to localize single seeds and also to distinguish individual seeds in a cluster. The method was tested on two phantoms: one with 20 seeds at discrete locations, and the other with 100 seeds located close to each other. The algorithm was able to identify the seeds within 1 mm of their physical locations for discreet seed localization. It was further able, without the operator intervention, to separate seeds at close proximity to each other with an average localization error of less than 2 mm. Another approach to the automatic seed localization was based on the Hough transform (Holupka et al., 2004). The computer program based on this approach works in fully automatic regime without the operator input. The method has been tested on phantoms and clinically used for post-implant dosimetric assessment on 1000 patients. Some modern treatment-planning systems have a capability of automatic seed detection. Figure 4 shows a Seed Finder™ module in software package Variseed 7.1 (Varian Medical Systems, Inc., Palo Alto, CA).
Electronic Portal Imaging Devices (EPIDs), Fluoroscopy, and Cone-Beam CT Fluoroscopic images, available during the permanent prostate implant procedure, can be used for periodic checking of the seed location during the implant delivery. Several groups (Tubic et al., 2001; Narayanan, Cho, and Marks 2002; Todor et al., 2003; Su et al., 2004) explored fluoroscopy for the seed localization and use of multiple projections for eliminating the source overlap problem. Intra-operative
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localization and dynamic dosimetric analysis provide an opportunity to compensate (during the implant delivery) for possible differences between the actual and planned seed positions. EPIDs became a much more used imaging modality in recent years, especially with the introduction of new amorphous silicon detectors. One of the capabilities of these detectors is to enable cone-beam imaging (Cho, Johnson, and Griffin 1995). This capabilities were used for intra-operative CT-scanning (Prestidge 2004), using cone-beam CT with an amorphous silicon detector in the Acuity simulator. Potential of this imaging modality includes intra-operative localization, reconstruction, and real-time dosimetry.
Source Localization for Various Brachytherapy Procedures LDR Implants Low dose rate (LDR) temporary interstitial implants are usually performed using catheters that are surgically inserted in the tumor. The treatment is performed using strings of low-dose sources separated by regular intervals. In order to plan the procedure and order sources of appropriate activities, “dummy wires” are inserted in the catheters exactly in the same positions as active wires. For the localization, two radiographs are taken at two angles, about 20° apart. Special care should be taken to obtain usable images with a clear view of sources and no overlap. Sometimes, a few gantry angles have to be checked, using fluoroscopy, before a suitable projection is chosen. After the films are taken and processed, the dummy markers on the films have to be identified and matched. After digitizing two-dimensional (2-D) coordinates of dummy sources, the 3-D reconstruction takes place, and all coordinates of the sources, which coincide with coordinates of the markers, are obtained. This information is used for planning brachytherapy treatment and choosing activities of the real sources that are located at the spots indicated by the dummy markers. Historically, many permanent prostate implant procedures were performed using fluoroscopy for visualization. These days, permanent prostate implants are usually done with ultrasound guidance that better shows the prostate borders than fluoroscopy. Although, some work has been performed to enhance the seed visibility in ultrasound imaging (Davis et al., 2003), still the accurate seed digitization requires xrays. If the imaging takes place in the operating room, it is typically performed by a C-arm fluoroscopy, which allows obtaining digital images at various angles. The images obtained can be used to reconstruct the implanted seed positions before all the planned seeds have been used. Dosimetric analysis performed in the middle of the procedure provides the dose distribution based on the actual, rather than planned, seed locations. In the case of possible deviations from the original plan, the intended positions of the remaining seeds can be recalculated before the implant delivery is resumed (Todor et al., 2003). After the implant delivery is finished, the patient undergoes a CT scan. Timing of CT varies from institution to institution. Some groups suggest that possible resolution of edema and tumor shrinkage makes dosimetry more realistic if the CT scanning is performed a few weeks after the implant. In addition to the CT, some institutions take a few more radiographs (typically, AP, lateral, and 20° from vertical for the supine setup) for the dose reconstruction in AP and lateral views (Roy et al., 1993). At many institutions the chest radiograph is also taken after the implant. The purpose of this is to check if there are any seeds that were carried away from the prostate with the bloodstream and ended up in the lungs.
HDR Brachytherapy Procedures For HDR treatments of prostate, catheter reconstruction has to be performed using a CT-set (Figure 3). Thin needles used for the catheter insertion can be imaged by either conventional or cone-beam CT (Prestidge 2004). Using dummy markers, one can locate HDR dwell positions that can be used for planning. Although the catheter positions could also be reconstructed using conventional radiographs,
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these radiographs would not show the prostate location. Therefore, CT use is necessary to visualize the target. The MammoSite® applicator (Proxima Therapeutics, Alpharetta, GA) is used for breast cancer treatment. The applicator consists of a spherical balloon with the diameter of 4 to 5 cm and a straight HDR catheter passing through its center. The empty balloon is inserted into a cavity in breast and filled with a diluted x-ray contrast agent (3%) and saline. There is only one HDR dwell position in the geometrical center of the balloon. Besides determination of the center position with respect to the patient’s anatomy, it is necessary to make sure that the balloon has a spherical form and is not flattened or squeezed to the shape of a football. The balloon deformation would lead to the distortion of the intended dose distribution. All these tasks can be performed by a CT-scan or a set of plain radiographs. Localization of the conventional tandem and ovoids applicator can be performed in the operation room using a fluoroscopy C-arm. “Dummy wires” are inserted in the applicator, and AP and lateral images are taken. Digital images taken with the use of the C-arm can be copied on a floppy disk or transferred over the computer network to the brachytherapy treatment-planning computer, where the positions of the dummy markers are digitized for the applicator reconstruction. Alternatively, if the applicators are CTcompatible, the CT scan with the applicator in place can be performed, so the applicator with the catheters are automatically reconstructed by the brachytherapy treatment-planning system. Instead of conventional CT, a cone-beam CT can also be used (Prestidge 2004).
Intravascular Brachytherapy Intravascular brachytherapy is used for treatment of benign disorders of coronary arteries. One of the radionuclides used for this purpose is a beta-emitter 90Sr. A small radioactive source is normally located inside a well-shielded container. If a special flexible catheter is attached to the source and a hydraulic pressure is applied using a special syringe, the source can leave the container and move along the catheter to the position. The dwell time is typically about 3 to 4 minutes, depending on the source strength and vessel diameter. After the dwell time elapses, a special manual switch is flipped, and the source returns back into the container driven by the water pressure created by the same syringe. Localization of the dwell position within a patient’s heart is critically important for the success of the procedure. The last U.S. vendor providing intravascular applicators (Novoste Corporation, Norcross, GA) suggested two methods of the source localization. The first method, which was mostly used with the catheter 5-French, involves a special dummy source in a separate container similar to the one containing the active source. After the flexible catheter is inserted to the proper depth, the “dummy” container is attached to it, and the non-active source is sent into position. The interventional cardiologist, using fluoroscopy, performs the test run for two purposes. First is to make sure that the non-active source can negotiate all the turns inside the catheter, and second, that it dwells in the right location. In 2003, Novoste introduced new types of sources used with new catheters 3.5 French. These catheters are coming with a dummy wire with a marker at the end, precisely in the same position to which the active source is supposed to go. When the catheter was initially inserted, the cardiologist would check the position of the dummy marker. If it was acceptable, the wire would be removed, and the catheter would be connected to the container with the active source for the treatment.
Other Methods of the Source Localization: In addition to the methods described above, there are other less common ways to localize radioactive sources. An electromagnetic tracking device was utilized for the intra-operative source localization and real-time treatment planning (Watanabe and Anderson 1995). The idea was to substitute x-rays for localization because of large, metal surgical retractors, which precluded radiographic exams. The
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nonradiographic localization suggested was performed by an electromagnetic tracking device, which consisted of a transmitter coil, a receiver coil, and a signal-processing unit. There was an RMS (root mean square) error of only 0.7 mm in determining the distances between points 2 cm apart, thereby demonstrating the feasibility of the method. Another imaging modality used for the prostate seed localization is MRI (Dubois et al., 1997; Dubois, Bice, and Prestidge 2001). The idea behind the method is to utilize an imaging modality that can adequately show both the sources of radiation and its target, i.e., the prostate gland. Generally, MRI gives the best picture of the prostate, allowing definition of its borders. Therefore, MRI was evaluated as a tool for localizing the seed locations. It was found that MRI yields an average localization error of (1.67±0.76) mm, about the same as the CT, which had (1.59±0.64) mm. Furthermore, the difference in isodose volumes for all calculated isodose levels was not statistically significant at 95% confidence levels (Dubois et al., 1997; Dubois, Bice, and Prestidge 2001). This observation was made for both 103Pd and 125I permanent prostate implants. For the HDR source localization, besides radiographic, one can utilize mechanical and optical methods. Thus for the microSelectron HDR 192Ir unit from Nucletron, there is a mechanical fixture which allows direct measurement of the source location (Jones 1989). The device allows the 192Ir source to pass into a transparent channel and to displace a plastic marker along a calibrated scale. This device, which is easy to use, provides direct, accurate information about the machine function. Another type of the device by Varian is utilized for the same purpose. To dertermine the active cable length, the source cable is sent into a special channel with a precise ruler at the end. When the source reaches its end position in this channel, a special digital camera takes a picture of the source side-by-side with the ruler. From the picture it is easy to determine the last dwell position of the active source and to calculate the position correction, if necessary.
References Amols, H. I., and I. I. Rosen. (1981). “A three-film technique for reconstruction of radioactive seed implants.” Med Phys 8:210. Anderson, L. L., F. W. Mick, K. Zabrouski, and Y. Watanabe. (1995). “Photoelectrons facilitate autoradiography for 192 Ir remote afterloaders.” Med Phys 22(11):1759–1761. Altschuler, M. D., P. A. Findlay, and R. D. Epperson. (1983). “Rapid, accurate 3D location of multiple seeds in implant radiation treatment planning.” Phys Med Biol 28:1305–1318. Biggs, P. J., and D. M. Kelley. (1983). “Geometric reconstruction of seed implants using a three-film technique.” Med Phys 10:701–704. Brinkmann, D., and R. W. Kline. (1998). “Automated seed localization from CT datasets of the prostate.” Med Phys 25(9):1667–1672. Chiu-Tsao, S.-T., T. L. Duckworth, N. S. Patel, J. Pisch, and L. B. Harrison. (2004a). “Verification of Ir-192 near source dosimetry using GAFCHROMIC film.” Med Phys 31(2):201–207. Chiu-Tsao, S.-T., T. Duckworth, C. Zhang, N. S. Patel, C. Y. Hsiung, L. Wang, J. A. Shih, and L. B. Harrison. (2004b). “Dose response characteristics of new models of GAFCHROMIC films: dependence on densitometer light source and radiation energy.” Med Phys 31(9):2501–2508. Cho, P. S., R. H. Johnson, and T. W. Griffin. (1995). “Cone-beam CT for radiotherapy applications.” Phys Med Biol 40(11):1863–1883. Davis, B. J., R. R. Kinnick, M. Fatemi, E. P. Lief, R. A. Robb, and J. F. Greenleaf. (2003). “Measurement of the ultrasound backscatter signal from three seed types as a function of incidence angle: application to permanent prostate brachytherapy.” Int J Radiat Oncol Biol Phys 57(4):1174–1182. Dubois, D. F., W. S. Bice, and B. R. Prestidge. (2001). “CT and MRI derived source localization error in a custom prostate phantom using automated image coregistration.” Med Phys 28(11):2280–2284.
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Dubois, D. F., B. R. Prestidge, L. A. Hotchkiss, W. S. Bice, and J. J. Prete. (1997). “Source localization following permanent transperineal prostate interstitial brachytherapy using magnetic resonance imaging.” Int J Radiat Oncol Biol Phys 39(5):1037–1041. Holupka, E. J., P. M. Meskell, E. C. Burdette, and I. D. Kaplan. (2004). “An automatic seed finder for brachytherapy CT postplans based on the Hough transform.” Med Phys 31(9):2672–2679. Jones, C. H. “MicroSelectron HDR Source Localisation Techniques.” Proceedings of the 5th International SELECTRON Users’ Meeting 1988. The Hague – The Netherlands. R. F. Mould (ed.). Leersum, The Netherlands: Nucletron International B. V., 1989. Khan, F. The Physics of Radiation Therapy. Baltimore, MD: Williams and Wilkins, pp. 380–381, 385–387, 2003. Li, Z., I. A. Nalcacioglu, S. Ranka, S. K. Sahni, J. R. Palta, W. Tome, and S. Kim. (2001). “An algorithm for automatic, computed-tomography-based source localization after prostate implant.” Med Phys 28(7):1410–1415. Lief, E. P., B. J. Davis, and K. E. Wallner. (1998). “CT scanning for post-operation seed localization in permanent prostate implants.” Med Phys 25:A172 (Abstract). Liu, H., G. Cheng, Y. Yu, R. Brasacchio, D. Rubens, J. Strang, L. Liao, and E. Messing. (2003). “Automatic localization of implanted seeds from postimplant CT images.” Phys Med Biol 48:1191–1203. Nag, S., W. Bice, K. DeWyngaert, B. Prestidge, R. Stock, and Y. Yu. (2000). “The American Brachytherapy Society recommendations for permanent prostate brachytherapy postimplant dosimetric analysis.” Int J Radiation Oncol Biol Phys 46:221–230. Narayanan, S., P. S. Cho, and R. J. Marks II. (2002). “Fast cross-projectional algorithm for reconstruction of seeds in prostate brachytherapy.” Med Phys 29:1572–1579. Prestidge, B. P. “Early Experience with Cone-Beam CT for Brachytherapy.” Materials of 8th Annual International Conference and Workshop “New And Future Developments In Radiotherapy.” Sponsored by Wayne State University, School of Medicine, Detroit, MI. November 12–14, 2004, San Diego, CA, 2004. Rosen I. I., K. M. Khan, R. G. Lane, and C. A. Kelsey. (1982). “The effect of geometric errors in the reconstruction of Iridium-92 seed implants.” Med Phys 9:220. Rosenthal, M. S., and R. Nath. (1983). “An automatic seed identification technique for interstitial implants using three isocentric radiographs.” Med Phys 10(4):475–479. Roy, J. N., K. E. Wallner, P. J. Harrington, C. C. Ling, and L. L. Anderson. (1993). “A CT-based evaluation method for permanent implants: application to prostate.” Int J Radiat Oncol Biol Phys 26(1):163–169. Smith, V., A. S. Meigooni, and S.-T. Chiu-Tsao. “Orthogonal Techniques,” “Stereo-Shift Technique and Automatic Reconstruction Techniques,” and “Specialized Techniques” in Interstitial Brachytherapy: Physical, Biological, and Clinical Considerations. Interstitial Collaborative Working Group. New York: Raven Press, Ltd., pp. 239–252, 1990. Stock, R. G., N. N. Stone, M. F. Wesson, and J. K. DeWyngaert. (1995). “A modified technique allowing interactive ultrasound-guided three-dimensional transperineal prostate implantation.” Int J Radiat Oncol Biol Phys 32:219–225. Su, Y., B. J. Davis, M. G. Herman, and R. A. Robb. (2004). “Prostate brachytherapy seed localization by analysis of multiple projections: Identifying and addressing the seed overlap problem.” Med Phys 31(5):1277–1287. Todor, D. A., G. N. Cohen, H. I. Amols, and M. Zaider. (2002). “Operator-free, film-based 3D seed reconstruction in brachytherapy.” Phys Med Biol 47:2031–2048. Todor, D. A., M. Zaider, G. N. Cohen, M. F. Worman, and M. J. Zelefsky. (2003). “Intraoperative dynamic dosimetry for prostate implant.” Phys Med Biol 48:1153–1171. Tubic, D., A. Zaccarin, J. Pouliot, and L. Beaulieu. (2001a). “Automated seed detection and three-dimensional reconstruction. I. Seed localization from fluoroscopic images or radiographs.” Med Phys 28:2265–2271. Tubic, D., A. Zaccarin, J. Pouliot, and L. Beaulieu. (2001b). “Automated seed detection and three-dimensional reconstruction. II. Reconstruction of permanent prostate implants using simulated annealing.” Med Phys 28:2272–2279. Watanabe,Y., and L. L. Anderson. (1997). “A system for nonradiographic source localization and real-time planning of intraoperative high dose rate brachytherapy.” Med Phys 24(12):2014–2023.
Chapter 13
Localization II: Volume Imaging Techniques and Accuracy for Brachytherapy Dosimetry Jason Rownd, M.S. Medical College of Wisconsin Milwaukee, Wisconsin Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 187 Volume Imaging . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 187 Volume-based Simulation and Planning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 189 Volume Dosimetry and Plan Evaluation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 198 Quality Assurance . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 200 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 200
Introduction Brachytherapy is inherently a three-dimensional treatment for a three-dimensional disease, and one should use as much information as available to correctly identify, localize, plan, and treat the patient. Concepts and definitions will be presented to provide a foundation for detailed discussions on volume-based brachytherapy, with specific case examples from the Medical College of Wisconsin.
Volume Imaging In two-dimensional (2-D) imaging modalities, the smallest image element is the pixel with a length and a width. In three dimensions, this concept expands to become a voxel, or volume element. Inherently, one can think of a three-dimensional (3-D) image as an array of 2-D planes collected together to describe the original object. Thus the three dimensions of a voxel can be thought of as the two dimensions of a pixel in the plane with the slice thickness, or slice separation, of adjacent planes as the third dimension. There are physical lengths that correspond to each side of an imaged voxel. The relationship between the physical length and the voxel unit can be described by its resolution, e.g., how many mm per voxel side. This is usually separated into the area (length and width) within a single imaged plane and the slice thickness, assuming no gaps in the imaged data between slices. In CT scans, each 2-D slice of the volume is mapped onto an X × Y pixel grid. This relationship is the same for all CT scanners, but the value of the resulting in-plane resolution is dependent on the field of view and the number of pixels used to image each plane. For example, with a 512 × 512 pixel array and a 50-cm diameter field of view, each imaged pixel corresponds to 0.98 mm on a side, or roughly 1 mm2/pixel resolution in the plane of the CT scan. The number of pixels and the field of view are machinedependent parameters that can and should be adjusted for the best results. The in-plane resolution places idealized limits on the ability to identify any object within the plane, specifically 1 pixel. A mistake in imaging of 1 pixel would correspond to a localization mistake in the patient of 1 mm. In addition to the in-plane resolution, there is slice separation and slice thickness. The slice thickness is a measure of how much of the real object volume is described by the imaged plane. The slice separation is a measure of how far apart each imaged plane is from an adjacent slice. Ideally, there should be no loss of information because of imaging gaps. The slice thickness should exactly correspond to the center-to-center slice separation of the 2-D views. Typical slice separations can be as small as 1 mm or as large as 1 cm. The idealized limits on voxel resolution apply when there are no intrinsic faults in the imaging system. Any number of mechanical and electronic faults can contribute to decreased resolution. The AAPM Task
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Group 66 Report No. 83 (Mutic et al., 2003), AAPM Report No. 1 (Judy et al., 1977), and AAPM Report No. 39 (Lin et al., 1993) detail a list of routine quality assurance (QA) tests and phantoms that can be used to identify problems with a CT-scanning system. In addition to machine-based concerns, the patient themselves can present problems for accurate volume reconstructions. High-Z materials, such as prosthesis or even brachytherapy applicators, can generate imaging artifacts. These artifacts can all but eliminate any useful volume information (see Figure 1). If the artifacts are sufficiently severe, then it can be impossible to achieve the ideal resolution of 1 pixel or voxel. One constraint on collecting all CT scans with the best resolution possible is the amount of memory each scan uses to store this information. The more information you require from the scanned volume, the larger the resulting data file. Fortunately, with modern computing systems, this is rarely a problem. The biggest problem seems to be a limit on the number of CT slices useable by a brachytherapy treatmentplanning system, either because of software-coded limits or because of hardware-constrained planning speed. In most cases, the more images manipulated by the planning system, the slower any updates are handled. There may not be a physical limit, but there could be a practical limit if the system slowdown is too severe. Ideally, the planner should acquire only as many slices as necessary to describe the volumes of interest with a resolution suitable to identify accurately structures of interest with the volume. When magnetic resonance imaging (MRI) studies are discussed, there are additional concerns. There are still the concepts of pixel and voxel resolution for MR scans. There are also concerns about imaging artifacts, but more specifically, imaging distortions. With CT scans, the resulting images are assumed to have nearly ideal reconstruction, where artifacts merely mask the original objects with streaks of highcontrast pixels. The problems with MR scans are more subtle and geometrical in nature. The AAPM Task Group 1 report (Price et al., 1990) gives some detailed QA procedures that can reduce machine-specific problems. The basic theory behind MRI is that mobile protons inside objects can be excited by the application of magnetic fields and ultimately detected. The magnitude of the excitation and the various associated relaxation times are partially dependent on the detailed composition of the object and the applied magnetic fields. The specific variations of tissues within the human body can produce imaged responses that vary
Figure 1. Axial view of a metallic tandem and ring applicator showing common streak artifacts in the CT scan.
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by orders of magnitude. Thus MR images are very good at resolving tissue-type differences, one of its key reasons for being used in imaging studies. This high sensitivity can also result in imaging artifacts when strong gradients in the magnetic susceptibility are present within an imaged object, specifically at structure interfaces, i.e., bone-air, tissue-metal. Instead of the streaking artifacts one might see in CT scans, one could see distorted geometry in the imaged object. The size and shape of the imaged object can vary by several millimeters from its actual dimensions. Reducing the possibility of such artifacts and distortions involves the proper choice of the field of view to exclude such interfaces whenever possible. In brachytherapy, this can also mean choosing applicators that are closer to tissue-equivalent with respect to their MRI responses, specifically nonmetallic applicators. Another key advantage of MR imaging studies is the variability in reconstructed views. The imaged plane in MR is controlled by the application of three orthogonal magnetic fields. With proper control of the various fields, it is possible to image specifically the axial, sagittal, coronal, or even oblique planes inside the imaged object. With this advantage, MRI can be controlled to image the specific planes that best describe the particular tissue being imaged. This better visualization can lead to better localization of the necessary structures, but may not improve applicator localization. In many treatment-planning systems, only axial views may be used in the planning process. Oblique images, while useful to the physician for imaging the tumor, are often ignored by the planning system. This limitation can be partially offset by comparison of axial planning views with oblique diagnostic views at the time of planning.
Volume-based Simulation and Planning Historically, brachytherapy planning was done based upon standardized rules and tables. Before the advent of modern imaging modalities, brachytherapy dosimetry was limited to standardized plans and calculated point doses. Even now, much of brachytherapy planning is done with radiographic films and planar, 2-D dosimetry. Normal tissue point doses are based upon radiographic films with contrast injected to accentuate the approximate location of normal tissues relative to the brachytherapy applicators (see Figures 2 and 3). Applicator reconstruction is based upon x-ray markers visible on the radiographs. Current planning for external beam radiotherapy treatments relies on the volume information provided by CT and MR scans. Brachytherapy treatment planning can take advantage of the same information these volume scans provide. Accurate brachytherapy treatment planning requires correct reconstruction of the patient anatomy, applicators, and source positions. In most cases, this is a matter of effective scan techniques with the recognition of the image limitations, appropriate applicator choices for the imaging modality, and estimates of the source positions within these applicators. The basic volume concern for any imaging modality is that it correctly describes the scanned object, in this case the patient anatomy with applicators in place. The choice of the imaging modality will affect the correct identification of the target structures. CT scans can be used to identify target structures when the target tissue is sufficiently distinct or when other methods are used to identify the target, such as surgical clips or even the applicator itself. MR scans are often better at imaging gynecological tumors (see Figures 4 and 5). In either modality, reducing the imaging artifacts is necessary for proper visualization of the tissues as well as the applicators. A single, large volume image study can contain hundreds of megabytes of computer data. The number of imaged slices, and the corresponding required data space, are functions of the desired scan volume and slice separation. Imaging only the volume that is needed for treatment planning reduces the burden on the computerized planning system as well as limiting patient exposure during the CT scan. As mentioned before, the imaging resolution for CT/MR scans within an imaged cross section is approximately 1 mm per pixel, based upon the field of view and the scan dimensions. The field of view must contain the entire applicator, tumor volume, and relevant organs at risk. Because CT images may not have sufficient soft tissue contrast, additional injected contrast may be necessary to differentiate between tumor and normal
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Figure 2. AP radiograph of a tandem and ring applicator with x-ray markers and injected contrast.
Figure 3. Lateral radiograph of the same applicator and contrast enhanced anatomy.
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Figure 4. Axial view of a CT-/MR-compatible tandem and ring insertion imaged by CT with clearly identified applicator and contrast enhanced tissues.
Figure 5. Axial view of the implant imaged by MR with the enhanced image of the tumor location now visible.
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tissues, i.e., rectum and bladder (see Figure 6). Contrast injected only into the Foley balloon is not sufficient to identify bladder proximity to the brachytherapy applicators (see Figure 7). MR scans offer better tissue differentiation without the aid of additional contrast media, but are not as commonly available for brachytherapy planning. In both MR and CT scans, the applicator choice can influence the image quality. Stainless steel applicators, suitable for radiographic film-based treatment planning are poor choices for volume studies. In CT scans, these applicators can cause enough streaking and artifacts to obscure completely the applicator location as well as limit the physician’s ability to visualize patient anatomy (Figure 1). Even in MR scans, nonferromagnetic steel applicators can still distort the shape and location of tissue relative to the applicator because the magnetic susceptibility gradient is so large between the metal applicator and the soft tissue of the patient. Carbon-fiber-based or plastic applicators can safely be used in both imaging modalities to reduce imaging artifacts and distortions. The slice separation and orientation of the applicators affects the accuracy of the reconstruction of the applicator within the scanned volume. Large slice separation values mean more interpolation of applicator geometry, and in some cases, unreasonable extrapolation. When the end of an applicator must be correctly identified, slice separations of 2.5 mm means an uncertainty of 1.25 mm in applicator tip position, and 5 mm of uncertainty if the slice spacing is 1 cm. As mentioned before, artifacts can completely obscure the applicator in cross-sectional views. Additional uncertainties in the accuracy of the reconstructed applicator can also occur when the applicator insertion is oblique to the cross-sectional imaging axis. If only cross-sectional views are used to identify an otherwise continuous applicator, the reconstructed applicator may not resemble the actual applicator. Whenever possible, the best imaging planes should be used to identify the applicator sections. Figure 8 shows the result of a poor reconstruction of the ring section of a tandem and ring applicator, while Figure 9 shows a better reconstruction of the true shape of the applicator when the ring section was reconstructed
Figure 6. Axial view of a tandem and ring insertion imaged by CT with bladder and rectal/sigmoid tissues enhanced by inject contrast.
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Figure 7. Axial view of the Foley bulb imaged by CT. The isodose lines going through the Foley bulb are clearly less than those going through other sections of the bladder in the previous figure.
Figure 8. Axial view of the ring section of a tandem and ring applicator taken with CT next to the poorly reconstructed result.
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Figure 9. Obliquely reconstructed view of the ring section of a tandem and ring applicator taken with CT and the resulting reconstructed applicator.
on an oblique imaged slice. In some cases, the orientation of the applicator is nearly orthogonal to the imaging access, such as the tandem section of a tandem and ovoid/ring applicator or the needles in a perineal implant. In these cases, the cross-sectional axis is a good choice for reconstruction. Adapting the scan views to visualize the applicator in the imaging studies is crucial to getting the best reconstruction of the applicator. The last part of the reconstruction is the trickiest, correctly identifying where the source positions fall within the actual applicator. In radiographic films, x-ray marker dummies spaced uniformly 1 cm apart can be used to identify a subset of source positions. In CT images, these same markers can be effectively used. The amount of distortion and artifacts produced by these markers is minimal and provide sufficient identification of the necessary source positions. Figure 10 illustrates this feature with respect to the tandem and ring applicator. The uncertainty in the source positions depends on the intrinsic pixel dimensions of the scan and the small amount of artifacts produced by the markers, as well as the proper placement of the markers. Unfortunately, the excellent tumor recognition of MR scans is offset by the lack of a suitable marker to identify possible source positions explicitly. In the case of MR-based treatment planning, other methods for identifying source positions are required. In some cases, it is possible to estimate the source positions based upon knowledge of the applicator geometry. In other cases, some other method must be used to reconstruct source positions. Image fusion of the CT scans and the MR scans is often used in external beam treatment planning and can also be used in brachytherapy treatment planning. CT scans tend to be better for identification of the applicator and source positions for reasons mentioned above, while MR scans can be better for imaging tumors relative to organs at risk. In the best of both worlds, contouring of the necessary targets can be done on the MR scans and fused onto the CT scans with the resulting treatment planning completed on the CT scans. 3-D applicators should be reconstructed by 3-D imaging studies. In single-catheter applications, such as endobronchial treatments, an orthogonal pair of radiographic films can suitably reconstruct the necessary
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Figure 10. Obliquely reconstructed views of a CT scan showing x-ray marker dummies inside the tandem and ring applicator used to identify possible source positions.
information for treatment planning. Vaginal cylinder applicators, with their simpler design and treatment goals, are also easily reconstructed from radiographic films with sufficient information about the organs at risk (bladder and rectum) being entered based upon injected contrast. Even in the single-catheter applicator treatments for breast brachytherapy, MammoSite® (Proxima Therapeutics, Alpharetta, GA), one could argue that all the planning information, dwell positions, and prescriptions could be obtained from films (see Figure 11). As the brachytherapy applicators get more complex and recognition of tumor and organs at risk become more necessary, accurate volume imaging and reconstruction becomes more necessary. Brachytherapy applicators can be broken up into two varieties: rigid and flexible catheters. Rigid applicators use metal, plastics, or carbon-fiber materials to maintain their shape during the procedure. The simplest rigid applicator is the hollow steel needle. Because of their metallic nature, these applicators can create imaging artifacts in volume studies. They are typically inserted into the patient with the help of a template and held in place by this template for the entire treatment. A 3-D imaging study of needle-based treatments has some benefits that overcome the inherent artifacts produced by steel needles. The main benefit is that the needles tend to run normal-to-axial image views and are easily reconstructed, providing a very accurate picture of the implanted volume. The main drawback is that the metallic needles produce image streaking that makes tissue delineation difficult (see Figure 12). Removing the additional metal of the needle stylets or obturators before taking the CT scan can minimize these artifacts. Additional contrast agents can also be used to identify the nearby rectum and bladder, while the treated volume is identified by needle locations. Another rigid applicator is the tandem and ovoids applicator used in gynecological treatments. This applicator consists of one catheter in the tandem, used to treat the cervix into the uterus and two catheters on either side in the ovoids, used to treat the vaginal surface and nearby tissues. The stainless steel variations can be imaged easily with radiographs, and CT-/MR-compatible versions exist for volume imaging. Similar to the tandem and ovoid applicator is the tandem and ring applicator. The pair of ovoids is replaced by a single ring section that rests up against the cervical os during treatment. The imaging
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Figure 11. En face radiograph of a balloon-based treatment for breast brachytherapy showing x-ray marker dummies and contrast-filled balloon.
Figure 12. Axial view of a cervix implant using steel needles and the resulting artifacts in the CT image.
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concerns are nearly identical between the tandem and ovoids applicator and the tandem and ring applicator. The straight line of the tandem is easily reconstructed in both. As mentioned before, the shape and placement of the ring section of the applicator make axial views insufficient to correctly reconstruct the applicator. To a lesser extent, the ovoid angle relative to the axial view can also make it difficult to reconstruct the applicator. A simpler applicator is the tandem and cylinder, which is really just a single catheter applicator, where the center catheter of the cylinder is extended by the tandem into the uterus. The relative simplicity of this applicator makes reconstruction trivial and the imaged volumes contain minimal artifacts. Flexible applicators can also be as simple as plastic needles instead of steel needles. The plastic tips of these needles can make their insertion more difficult and the plastic needles may not stay as rigid during the treatment as the steel needles. However, the amount of artifacts produced in scans is greatly diminished. If the template used to hold the needles in places is also MR compatible, then the entire applicator can be imaged as necessary using either MR scans or CT scans. Prostate (Figure 13) and cervix implants can be done this way. Identifying source positions requires knowing the dimensions of the plastic needles and the necessary offsets to be applied to the imaged catheters. More complex implants can be handled using flexible catheters inserted with the aid of hollow needles. The hollow steel needles are removed, leaving the flexible plastic catheters behind in the tissue held in place by sutures or adjustable guides pulled against the skin. Post-resection soft tissue sarcoma beds can be irradiated using a single plane of these flexible catheters, and probably require little if any volume imaging. Orthogonal radiographs and x-ray marker dummies can be used to identify the implanted catheters and treatment region sufficiently.
Figure 13. Axial view of a prostate implant using plastic needles and the negligible artifacts in the CT image.
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Post-lumpectomy volume implants of the breast usually require more than a single plane of catheters. The multi-planar nature of the implant, the irregular nature of the target volume, and the proximity of organs at risk, such as the skin, make imaging crucial for proper coverage and treatment.
Volume Dosimetry and Plan Evaluation Once the brachytherapy applicator, of whichever type, has been implanted into the patient and imaged for treatment planning, a suitable prescription must be developed. Many different methods of prescription are available for brachytherapy treatments. The prescription goal may simply be the dose to an individual point or group of points. It may also be based upon careful selection of the target volume and the percentage of that target covered by the prescription dose. Brachytherapy treatments of the prostate use various scans (ultrasound/MR/CT) to determine target volumes with prescriptions based upon dose volume histogram (DVH) analysis to get the prescribed dose to a specified percent of the target volume. There are two common methods for partial breast irradiation using brachytherapy: the MammoSite balloon and multicatheter volume implants. The prescription goal of the balloon implant is to deliver a desired dose to a distance of 1 cm from the balloon surface. Imaging the balloon within the breast after surgery is done to obtain the balloon diameter and tissue conformance around the balloon (see Figure 14). When using multiple catheters to implant the breast for brachytherapy treatment, CT scans are necessary to identify both the catheter positions and the target volumes for planning (see Figure 15). The prescrip-
Figure 14. Axial view of a breast brachytherapy treatment using the MammoSite® balloon imaged by CT showing exact size and tissue conformance.
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Figure 15. Axial view of a breast brachytherapy treatment using multiple plastic catheters imaged by CT showing catheters, lumpectomy cavity, and target volume for treatment planning.
tion goals for this type of implant can be more complex. Maximizing target coverage and minimizing normal tissue doses require accurate catheter reconstruction. This process begins with the placement of the catheters and their position relative to axial planes of the CT scan. Catheters that run perpendicular to the axial slices are easiest to reconstruct, but this is usually not the most efficient way to cover the target volume. Catheter placement exactly parallel to the axial slices of the CT scan is often easiest for the physician but provides the most difficult reconstruction since lengths of the catheters may be only partially imaged if they are placed between axial slices as imaged. Compromising by placing catheters at an oblique angle gives enough information to reconstruct the implant accurately while allowing the physician to use the patient’s anatomy, i.e., chest wall, to guide the catheter placement. The plastic catheters create minimal artifacts with the only problem being the identification of possible source locations. Observing how the x-ray marker dummies rest within the plastic catheters, one can determine where the effective endpoint of each catheter is for source loading and plan accordingly (Figure 16). In gynecological implants, e.g., tandem and ovoids applicators and vaginal cylinders, it is common to prescribe a given dose to a particular point relative to the applicator, such as point A for cervix cases, or to the surface of the applicator, such as the surface of a vaginal cylinder applicator. In these cases, it is less common to define specific target volumes and desired treatment volumes. Various weights or prescription percentages can be prescribed to different sections of the applicator, but this is usually achieved by individually modifying the prescription points on a case-by-case basis. As more of these cases are imaged and planned with CT and MR scans, more prescription options can be developed.
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Figure 16. En face radiograph of a multiple plastic catheter breast brachytherapy simulation showing the location of the most distal source position near the end of each catheter within the support ball on the skin.
Quality Assurance Given that the scanners and their output are accurate with respect to the anatomy and applicator imaging, the final dosimetry should still be evaluated for accuracy and compliance with the prescription goals. The prescription compliance can be addressed by detailed inspection of the isodose distributions and DVHs. Additional inconsistencies can also be caught by comparing current plan results with previously treated cases or by explicit point-dose calculations using another treatment-planning system. In all treatment plans, the initial limitations on the imaging accuracy combined with the uncertainties of the applicator reconstructions and target volume delineations result in final uncertainties in the dosimetry which should be evaluated case by case.
References Judy, P.F., S. Balter, D. Bassano, E. C. Mccullough, J. T. Payne, and L. Rothenberg. (1977). “Phantoms for Performance Evaluation and Quality Assurance of CT Scanners.” AAPM Report No. 1. Chicago: AAPM, 1977. Lin, P.-J. P., T. J. Beck, C. Borras. G. Cohen, R. A. Jucius, R. J. Kriz, E. Nickoloff, L. Rothenberg, K. J. Strauss, and T. Villafana. “Specification and Acceptance Testing of Computed Tomography Scanners.” AAPM Report No. 39. New York: American Institute of Physics, 1993. Mutic, S., J.R. Palta, E. K., Butker, I. J. Das, M. S. Huq, L. D. Loo, B. J. Salter, C. H. McCollough, and J. Van Dyk. (2003). “Quality assurance for computed-tomography simulators and the computed-tomography simulation process: Report of the AAPM Radiation Therapy Committee Task Group No. 66.” Med Phys 30:2762–2792. Also available as AAPM Report No. 83. Price, R. R., L. Axel, T. Morgan, R. Newman, W. Perman, N. Schneiders, M. Selikson, M. Wood, and S. R. Thomas. (1990). “Quality assurance methods and phantoms for magnetic resonance imaging.” Med Phys 17(2):287–295. Also available as AAPM Report No. 28.
Chapter 14
Semiempirical Dose-Calculation Models in Brachytherapy Jeffrey F. Williamson, Ph.D. Department of Radiation Oncology Virginia Commonwealth University Richmond, Virginia Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 201 Biological Dosimetry Era: From Skin Erythema Dose to Exposure Rate (1900–1940) . . . . . . . . . 202 Classical Dosimetry Era (1940–1990) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 202 Quantitative Dosimetry Era (1990–Present) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 203 Source Strength Specification Quantities and Units . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 203 Air-Kerma Strength and Absorbed Dose . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 203 Activity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 204 Relationship Between Activity and Exposure Rate . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 204 Obsolete Quantities for Specifying Source Output . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 205 Semiempirical Dose-Calculation Formalisms: . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 206 Isotropic Point Source . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 206 Commonly Used Notation for Semi-empirical Dose Calculation . . . . . . . . . . . . . . . . . . . . . . . . . . 209 Alternative Classical Model Nomenclature . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 209 Relationship of Semi-empirical Dose-Calculation Model to TG-43 notation . . . . . . . . . . . . . . . 210 Application of Semiempirical Model to Practical Sources . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 211 Modeling of Source Anisotropy: The anisotropy factor . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 211 Dose Calculation for Extended Sources: The 1-D Pathlength (Sievert Integral) Model . . . . . . . 212 Clinical Application of Semiempirical Models . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 214 Implementation of Point-Source Models for 192Ir and 137Cs Seed Sources . . . . . . . . . . . . . . . . . . . . 214 Implementation of 1-D Pathlength Models for 137Cs Intracavitary Tubes . . . . . . . . . . . . . . . . . . . . 215 Application of 1-D Pathlength Models to Other Radionuclides . . . . . . . . . . . . . . . . . . . . . . . . . . . . 218 Application of 1-D Pathlength Models to Internally Shielded Applicators . . . . . . . . . . . . . . . . . . . 221 Advanced Analytic Dose-Calculation Algorithms . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 222 Semiempirical Models in Brachytherapy Quality Assurance . . . . . . . . . . . . . . . . . . . . . . . . . . . 223 Example: Manual Calculation of Dwell Time for an HDR Vaginal Cuff Insertion . . . . . . . . . . . . . 224 Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 225 Appendix: Derivation of Exposure Rate and Air-Kerma Rate Constants . . . . . . . . . . . . . . . . . 227 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 229
Introduction Analytical and semiempirical dose-calculation models have a long and distinguished history in brachytherapy. As used in this review, a semiempirical model is a mathematical representation of the doserate distribution about a brachytherapy source characterized by (a) relatively few unknown parameters; (b) dependence of nongeometric model input parameters only on the radionuclide spectrum; and (c) methods of evaluating input parameters that do not require measured dose distributions for sources of the same type and construction as those being modeled. The most common semiempirical dose-computation models assume an idealized point source dose distribution or isotropic dose kernel, and derive dose rates around needle and tube sources by integrating the dose kernel over the extended radioactivity distributions. Such models were used almost exclusively in clinical brachytherapy through the 1980s because of the technical difficulties in performing high-resolution absorbed dose measurements around sources because of steep dose gradients, relatively low dose rates, and low-energy photons emitted by some brachytherapy sources. A form of the point-kernel model was introduced by Rolf Sievert in 1921 (Sievert 1921) which used an
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analytic formula, the Sievert integral, to perform one-dimensional (1-D) integration of the point-source dose kernel over the active length of a radium needle. Such models were quickly adopted by Quimby (1932) and Parker (Meredith 1967; Parker 1938) who incorporated this methodology into the QuimbyMemorial and Manchester implant systems, respectively. Upon the advent of computer-assisted treatment planning, Sievert integral-type models were widely used for 137Cs tubes (Krishnaswamy 1972), radium needles and tubes (Stovall and Shalek 1972), and even 125I seeds (Krishnaswamy 1979). While semiempirical models have largely been displaced in brachytherapy planning by more rigorous Monte Carlo and measurement-based methodologies, they are still useful for calculating dose-rate for 137Cs and 192Ir implants, and for manually verifying computerized dose calculations. Computation of absorbed dose and specification of source strength are tightly coupled together. To appreciate this relationship and to better understand the role of computational models in brachytherapy dose specification, it is useful to describe the history of brachytherapy in terms of three dosimetric eras: biological dosimetry (1900–1940), classical dosimetry (1940–1990), and quantitative dosimetry (1990–present).
Biological Dosimetry Era: from Skin Erythema Dose to Exposure Rate (1900–1940) During this time period, interstitial and intracavitary brachytherapy using radium needle and tubes became established as the first radiotherapeutic technique to successfully treat deep-seated tumors. By 1908, Marie Curie introduced the quantity activity and developed an approach for standardizing 226Ra source strength in terms of the mass of 226Ra contained in a source in 1913. By 1921, 226Ra standards, consisting of known masses of radium salt in glass vials, were available in Paris and Vienna for intercomparison with clinical sources. This allowed implant therapy to be specified in terms of “mg-h” (the product of mass of 226Ra in the implant and treatment duration). In the meantime, development of the free-air chamber [see Wyckoff (1980)] for an interesting historical review) allowed external beam-therapy to be specified in terms of the directly measurable quantity, exposure, by the late 1920s. A major problem was how to specify brachytherapy in terms of a comparable quantity. Some implant systems, e.g., the Manchester system, used Sievert-like dose-calculation algorithms in conjunctions with an estimated 226Ra exposure-rate constant, quantifying treatment delivery in terms of roentgens. However, rigorous physical measurements to confirm the exposure rate constant or otherwise establish the equivalence of 226Ra Roentgens (a megavoltage treatment modality) with the more rigorous external-beam exposure standard (limited to kV x-ray beams) was not possible in this era. The literature of this period documents numerous and often successful efforts to use biological dosimeters to intercompare different kilovoltage therapy beams and transfer instruments through the late 1920s and to intercompare radium treatment with orthovoltage therapy until about 1940. The most common endpoint was threshold erythema dose (TED) developed and promoted by Memorial Hospital (Quimby 1941). One unit of TED was that quantity of radiation needed to produce a detectable erythema in 80% of the subjects so irradiated. An exposure 1 TED of low dose-rate irradiation was biologically equivalent to about 11 Gy.
Classical Dosimetry Era (1940–1990) This era was accompanied by the development of mature classical systems for brachytherapy source implantation and the gradual replacement of 226Ra sources and artificially produced radionuclides such as 192Ir and 137Cs. It began with the successful application of Bragg-Gray cavity theory 22 to the calibration of 226Ra and other high-energy sources in terms of exposure (Attix and Ritz 1957), which allowed brachytherapy quantified using the same system of units and quantities as the external orthovoltage-beam therapy of the day. This development permitted brachytherapy dose delivery to be specified in terms of a rigorously defined quantity, exposure, which could, in principle, be measured by an instrument calibrated
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against a primary standard. In this chapter, we will discuss dose-calculation algorithms that seek to estimate absorbed dose in medium from sources that have been calibrated in terms of an exposure-rate or air-kerma rate standard.
Quantitative Dosimetry Era (1990–Present) The modern or quantitative era of brachytherapy dosimetry began to influence clinical practice in the 1990s and continues to the present. Quantitative dosimetry relies upon either dose-rate measurements around each specific source model or three-dimensional (3-D) radiation transport calculations based upon a validated geometric model of the source. In contrast to the semiempirical dose modeling philosophy, radionuclide-specific model parameters are broadened to include source-specific absorbed estimates derived from either measurements or from Monte Carlo simulations accounting fully for the geometry of the specific source. Quantitative dosimetry methodologies place less emphasis on uniformity of clinical dose-calculation practice in favor of physically accurate dose-rate estimates with well-defined uncertainties. The technical developments leading to clinical acceptance of quantitative dosimetry methods were motivated by concerns that semiempirical dose-calculation algorithms were not valid in the low-energy regime of 125I and 103Pd sources. To clinically utilize dose measurements and Monte Carlo calculations, a fully empirical dose-calculation formalism, the AAPM TG-43 protocol was developed and published in 1995 (Nath et al., 1995) and recently revised in 2004 (Rivard et al., 2004). Both the classical and quantitative dosimetry methods are based on the principle that brachytherapy source strength should be specified in terms of radiation output in free air. The goals of this chapter are to (1) review widely used and historical dose-calculation formalisms, (2) identify their dependence on the source-strength specification quantity used, (3) review the literature documenting the accuracy of semiempirical models, and (4) discuss their application in current clinical practice.
Source Strength Specification Quantities and Units Air-Kerma Strength and Absorbed Dose As discussed in more detail in chapter 11 of this book, photon-emitting brachytherapy sources are specified in terms of air-kerma strength, denoted by SK, in North America. The AAPM (Rivard et al., 2004) currently defines SK as the product of air-kerma rate, K" δ (d ) , in vacuo and due to photons of energy greater than ∂, at distance d, and the square of this distance, d2. SK = K" δ , air (d ) ⋅ d 2 .
(1)
The point of K" δ , air specification is located on the transverse-plane of the source (the plane normal to the longitudinal-axis of the source which bisects the radioactivity distribution). The unit of air kerma strength is 1 µGy·m2·h–1, and is often denoted in the literature by the symbol “U” where 1 U = 1 cGy·cm2·h–1 = 1 µGy·m2·h–1. Air-kerma strength is numerically (but not dimensionally) equal to the quantity reference air-kerma rate, K" ref , a very similar quantity defined by the International Commission on Radiation Units and Measurements (ICRU) (1998) and generally used outside North America. The quantity SK describes source strength in terms of kerma, Kx, which is the total kinetic energy transferred to charged particles by photon interactions with atoms per unit mass of material x. Although kerma can be specified in any medium x, usually air with (x = air), is assumed for radiation metrology and replaces the now obsolete quantity exposure, X. Absorbed dose to air, Dair, and SK are closely related:
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Dair = X ⋅
() W e
= SK ⋅ (1 − g ) ≈ SK ,
(2)
where (W/e) is the average energy imparted to air per ion pair created and is a constant, independent of photon energy: (W/e) = 33.97 eV/ion pair = 33.97 J/C = 0.876 cGy/R (Boutilon and Perroche-Rous 1987). The factor g is the fraction of kinetic energy transferred to the medium that is converted back to radiant energy (photons) by the Bremsstrahlung process. The quantity Kx · (1 – g) is often called “collisional kerma,” and refers to that component of the transferred energy which is ultimately absorbed by the medium via inelastic secondary charged particle collisions. Because g ≤ 0.001 at brachytherapy energies and tissuelike media, this radiative correction is usually ignored, further simplifying equation (2). Equation (2) assumes secondary charged particle equilibrium (CPE) (Attix 1986). CPE is realized when kerma remains relatively constant over the secondary electron range, assuring that the rates of energy absorption and energy transfer are approximately equal, and that kerma closely approximates absorbed dose. Virtually all brachytherapy dose-calculation algorithms and dosimetric analyses assume that CPE exists, and that D ≅ K everywhere. Although generally valid, CPE can be expected to break down in the presence of steep dose gradients very near sources (Roesch 1958), near metal-tissue interfaces (Nath, Yue, and Liu 1999), and within the active elements of thin, bounded detectors (Burlin 1966).
Activity To define the obsolete quantities for describing source output and to understand the classical formulations of semiempirical dose-calculation models, the quantity activity, A, must be introduced. It is defined as the rate of nuclear disintegration or transformation within a radioactive source. Its contemporary unit is the becquerel (1 Bq = 1 disintegration/s). We will freely use the more traditional but obsolete unit, the curie (1 Ci = 3.7 × 1010 disintegrations/s = 3.7 × 1010 Bq). Each disintegration represents the spontaneous transformation of an atom from one nuclear state to another. For most brachytherapy radionuclides, such nuclear state transformations are accompanied by emitted photons in the form of unconverted γ-rays, annihilation photons, characteristic x-rays, and Bremsstrahlung photons. Activity is measured by counting the number of photons, β particles, etc., emitted by an unencapsulated point source of the radionuclide by means of scintillation or proportional counters, from which its activity is inferred (NCRP 1985). For sealed brachytherapy sources, activity refers to that contained inside the encapsulation. As defined in this strict sense, activity is no longer used in brachytherapy dosimetry. However, activity continues to serve as the basis for treatment specification and dosimetry of unsealed radiopharmaceuticals used for diagnosis and therapy, and may play a future role in dosimetry of sealed beta-emitting sources for intravascular brachytherapy. NIST maintains activity standards for a wide variety of radionuclides in aqueous solution (Cavallo et al., 1973; Coursey et al., 1992).
Relationship Between Activity and Exposure Rate The activity of a photon-emitting radionuclide and the exposure rate in free space, X" δ (r ) , at distance r due to photons of energy greater than δ, are related by a fundamental quantity, the exposure rate constant, ( Γδ )x , defined as follows (ICRU 1998):
( Γ δ )x =
X" δ (r ) ⋅ r 2 A
.
(3)
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has units of R·cm2·mCi–1·h–1 and is equal to the exposure rate in R/h at 1 cm from a 1 mCi point source. It describes the rate at which air is ionized due to the emission of photons resulting from radioactive decay. The purpose of the energy cutoff δ in equations (3) and (1) is to eliminate low-energy Bremsstrahlung and characteristic x-rays that are always absorbed within any practical source. Because ( Γδ )x is defined in terms of an isotropic point source and exposure rates corrected for air attenuation and scattering, inverse-square law applies exactly. Thus, ( Γ δ )xis independent of the distance r used in equation (3). ( Γδ )x depends only on the number and energy of the photons emitted per disintegration. Suppose there are N different photons emitted per disintegration with energies E1, E2, . . ., EN in units of MeV. Each time an atom decays, suppose Pi photons of energy Ei are emitted where i = 1, . . ., N. The list {Ei , Pi }N constii =1 tutes the photon spectrum of the radionuclide. If the spectrum is known, then ( Γ δ ) can be calculated by: δ
x
x
(Γ ) δ
N
X
= 193.8 ⋅ ∑ Pi ⋅ Ei ⋅ ( µ en / ρ )iair ,
(4)
i =1
where ( µ en / ρ )iair is the mass energy absorption coefficient (in units of cm2/g) for air at energy Ei. A detailed derivation of this fundamental relationship is given in the appendix of this chapter. As explained in the appendix, the exposure-rate constant has been replaced by the air-kerma rate constant, ( Γ δ )K . As originally defined by the ICRU (1998), is a fundamental property of the unencapsulated radionuclide photon spectrum, applies only to idealized point sources, and neglects many significant properties of real sources such as self-absorption, filtration, and physical distribution.
Obsolete Quantities for Specifying Source Output Numerous quantities and units have been introduced over the decades for specifying strength of brachytherapy sources, many of which make misleading references to physical quantities having no direct connection to the primary standard from which the given calibration was derived. Confusion can be avoided by recognizing that all photon-emitting sealed brachytherapy sources are specified in terms of air-kerma strength, and that obsolete quantities are merely historically convenient means of restating SK. The oldest source-strength quantity, MRa, the mass of 226Ra contained inside a source excluding the nonradioactive core components and radioactive decay products, was introduced by Marie Curie in 1913 (Chavaudra 1995), well before the quantity activity was defined. Indeed, the unit curie was originally defined as the number of disintegrations produced by 1 g of 226Ra. MRa standards were prepared by carefully weighing pure 226Ra samples in an analytical balance. The National Institute of Science and Technology (NIST) MRa standard was introduced by Hönigschmidt in 1934 (Cavallo et al., 1973). To calibrate a user’s source in MRa, its radiation output is compared with that of the NIST radium standard by means of an ion chamber. NIST no longer offers an MRa calibration service. In contrast to the other radionuclides, exposure-rate constant of 226Ra (denoted by the special symbol (Γ δ)Ra,t in (Γ δ)Ra,t this chapter), is tabulated as a function of its effective capsule thickness, t, in mm of platinum. (Γ δ)Ra,t is normalized to the mass of radium contained in the source and has units of R·cm2·mg–1·h–1. Since MRa is limited to 226Ra sources, it is no longer used. The quantity is equivalent mass of radium (Meq), was introduced in the 1950s as a generalization of MRa, applicable to artificial radium substitutes as well as 226Ra. Meq is defined as that mass of 226Ra filtered by 0.5 mm Pt that has the same SK as that of the given source. Describing the strength of 60Co and 137Cs sources in terms of Meq allowed implant and radium needle dosimetry tables, which gave dose per mg-h of 226Ra, to be used without modification for these new sources. Because Meq is simply a statement of SK relative to that of a hypothetical radium needle, the given source being quantified need not contain 226Ra, be encapsulated in Pt, nor have a wall thickness of 0.5 mm. Since Kair = X · (W/e) and (Γ δ)Ra,0.5 = 8.25 R·cm2·mg–1·h–1 for 226Ra filtered by 0.5 mm Pt3, SK and Meq are related by:
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M eq =
(Γ ) δ
SK Ra , 0.5
⋅ (W / e )
=
SK
.
(5)
7.223
Meq continues to be widely used to specify strength of intracavitary and interstitial brachytherapy radium-substitute sources such as 137Cs and 192Ir. Similar to the approach of using Meq, apparent activity, Aapp, is a statement of source output relative to that of a hypothetical unfiltered point source. Aapp is the activity of a hypothetical unfiltered point source that has the same SK as that of the given source. Aapp =
SK
( Γ ) ⋅ (W / e ) δ
(6)
X
Apparent activity in units of mCi continues to be widely used for specifying strength for permanent interstitial implants (e.g., 125I and 103Pd sources). In contrast to Meq, which is based on the universally accepted (Γ δ)Ra,0.5 value of 8.25 R·cm2·mg–1·h–1, no consensus as to (Γ δ)X values for the other radionuclides exists. Often different vendors will assume different values for the same radionuclide. Thus, Aapp is an inherently ambiguous means of describing source strength. For 103Pd and 125I sources, the AAPM recommended (Williamson et al., 1999) that the (Γ δ)X values of 1.476 and 1.45 R·cm2·mg–1·h–1, respectively, be used for specification of Aapp.
Semiempirical Dose-Calculation Formalisms Isotropic Point Source Consider a hypothetical unencapsulated point source embedded in condensed medium, x, with an air· kerma strength of SK, illustrated in Figure 1. Our aim is to evaluate the dose rate, D (r), as a function of · radial distance, r. Because of spherical symmetry, D (r) has no angular dependence. Assuming CPE, equa· tion (2) implies that the dose rate to air in free space is equal to air-kerma strength: Dair (1 cm) = SK. Because
Figure 1. Unencapsulated point source of strength SK immersed in an unbounded water-equivalent medium.
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of the straight-line emission of photons in all directions (see Figure 2), photon intensity or fluence, Φ(r), at any point is proportional to the inverse square of its distance, r: Φ(r ) =
no. incident photons unit area irradiated
=
noo. photons emitted 4π r 2
∝ Dose(r ) ∝ Exposure(r ) ,
(7)
assuming that attenuation and scattering can be neglected. As a result of this purely geometric effect, the absorbed doses in free space, D(r1) and D(r2) at the two distances r1 and r2 are related by:
r = = 2 . D (r2 ) Φ(r2 ) r1
(8)
D" air (r ) = SK r 2 .
(9)
D (r1 )
2
Φ(r1 )
Letting r1 = 1 cm and r2 = r, we obtain
We now consider the effect of the tissue medium, x, on the dose distribution. To correct for difference in energy transfer efficiency between medium and air, we exploit the fundamental relationship between particle fluence and dose derived in the appendix, Dx(r) = Φ(ρ) · E · (µen,x/ρ). This relationship implies that that the dose to arbitrary x, and dose to air are related by: ,
D" x = µ en ρ D" air free space
(
) (µ ρ) ≡ (µ x
en
air
en
/ρ
)
(10)
x air
Figure 2. An isotropic point source of activity, A. To illustrate the derivation of inverse square law, the source is surrounded by vacuum and placed at the center of two concentric spherical surfaces of radii r1 and r2. By definition, an isotropic point source -Has no extension -Radiates photons with equal likelihood in all directions in straight-line paths.
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where the bar indicates that the corresponding quantity is the energy-fluence weighted average over of the photon spectrum. We now consider the effect of the condensed medium surrounding the source. The competition between photon attenuation and buildup of scattered photons can be described by the tissue attenuation and buildup fact, T(r), defined as
T (r ) =
Dx in medium Dx in free space
=
Exposure in mediuum at distance r Exposure in air
. r ce from a point sou
(11)
Sometimes this factor is called the “kerma-to-dose conversion factor,” depending on at which point in the derivation the CPE approximation is invoked. Substituting equations (10) and (11) into (9), and assuming that x = wat (water), we obtain the final result: ( µ / ρ )wat D" wat (r ) = SK ⋅ en 2 air ⋅ T (r ) . r
(12)
For all photon-emitting radionuclides with energies greater than 200 keV, including all radium substiwat tutes, ( µ en / ρ )air has the value 1.11 in water. Figure 3 shows T(r) for several radium-substitute radionuclides as well as for a few low-energy radionuclides. For 226Ra-equivalent radionuclides, T(r) deviates less than 5% from unity for r ≤ 5 cm. Numerous tabulations of T(r) are available in the literature; those of Meisberger and colleagues (Meisberger, Keller, and Shalek 1968), Berger (1968), and Van Kleffens and Star (1979) are among the best known. Most of these data are derived from theoretical photon
Figure 3. Photon attenuation and scatter factors, T(r), for a number of brachytherapy radionuclides. The left panel shows T(r) for a number of radium-equivalent radionuclides, as presented in the classic paper of Meisberger and Keller (1968). The right side shows T(r) for several low-energy radionuclides. Meisberger et al., fit their data to a third-degree polynomial T(r) = A + B · r + C · r2 + D · r3 that is widely used to represent T(r) in treatment planning systems.
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transport calculations. The classical semiempirical model assumes that T(r) is a function only of the radionuclide photon spectrum and that a single data set (e.g., for 192Ir) can be used for all 192Ir sources regardless of their construction. By solving equations (5) and (6) for SK and substituting the results in equation (12), one can derive equations relating the dose rate at distance r to equivalent mass of radium and apparent activity for the same unfiltered point source:
D" wat ( r ) = M eq ⋅ D" wat ( r ) = Aapp ⋅
(Γ ) δ
(Γ ) δ
Ra , 0.5
r X
r
2
⋅ fwat
2
⋅ T (r )
Equivalennt Mass of
226
Ra
(a )
(13)
⋅ fwat
⋅ T (r )
Apparent Activity
(b )
where fwat is the dose-to-exposure conversion factor given by: fwat =
Dwat X
= (W / e) ⋅ ( µ en / ρ )air = 0.876 wat
cGy R
⋅ ( µ en / ρ )air
wat
}
in free space .
(14)
For all 226Ra substitutes, fmed has the value 0.974 cGy·R–1 for water and 0.966 cGy·R–1 for muscle (Johns and Cunningham 1983). · Equations (12) and (13) give Dx (r), for a point source surrounded by water medium (easily generalized to an arbitrary medium) that has been specified in terms of Meq, Aapp, and SK. Assuming that the same exposure-rate constants, (Γ δ)Ra, 0.5 and (Γ δ)X were used to evaluated absorbed dose as were used to convert the measured air-kerma strength to Meq and Aapp via equations (5) and (6), all three equations should give numerically identical dose rates. This demonstrates that (Γδ) is, in fact, a “dummy” constant that plays no physical role in the dosimetry of output-calibrated sealed sources since any arbitrary, but consistently used, value will yield identical dose-rate distributions. These unit conversions may be performed by different individuals. For example, the vendor calibrates 125I sources by intercomparing them with the NIST SK standards. The vendor calculates Aapp from the measured SK by equation (6) using an assumed (Γ δ)X value and records the result on the source’s calibration certificate. The hospital physicist, in calculating dose rates by equation(13), must also use an assumed (Γ δ)X value. If the physicist fails to use the same value as the vendor, significant dose-calculation errors may result. Use of SK for clinical source-strength specification eliminates these dummy constants, thereby eliminating errors due to inconsistent conventional choices. For more details on the relationships between the various source-strength quantities and units and their impact on dose calculation, the reader is referred to the review by Williamson and Nath (1991).
Commonly Used Notation for Semiempirical Dose Calculation Numerous notational systems have been used in the literature to formulate equation(12). Only a few of the more common approaches will be outlined here. Alternative Classical Model Nomenclature One commonly used quantity in classical dose calculation is the energy-absorption build-up factor, B(µr). The build-up factor is defined as B( µ r ) =
total dose in medium at distance r
primary dose in medium m from a point source
,
(15)
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Jeffrey F. Williamson
where µ [cm–1] is the total linear attenuation coefficient of the primary spectrum. The most commonly cited source of buildup factors is the classic article by Berger (1968) which is based upon a semianalytic solution of the Boltzmann transport equation. Meisberger et al. (Meisberger, Keller, and Shalek 1968) were the first to suggest that these theoretical data could be used to approximate the transverse-axis dose distribution around actual interstitial sources. For an ideal point source in water, equation (12) becomes ( µ / ρ )wat D" wat (r ) = SK ⋅ en 2 air ⋅ e− µ r ⋅ B( µ r ) , r
(16)
− µr which implies T (r ) = e ⋅ B( µ r ) . A related quantity, used in many modern analyses (Williamson 1996) is the scatter-to-primary ratio, SPR(r):
SPR(r ) ≡
scatter dose in medium at distance r
primary dose in meddium from a point source
(17)
T (r ) = e− µ r ⋅ (1 + SPR(r ))
A precursor of the radial dose function (RDF), as defined by the AAPM TG-43 report (Nath et al, 1995; Rivard et al., 2004), was introduced by Dale (1983). This function, also called RDF, is denoted by the symbol g– (r) to –distinguish it from the TG-43 quantity, gX(r). Dale defined it as g (r ) =
D (r, π / 2 ) ⋅ r 2
D (1 cm, π / 2 ) ⋅ (1 cm )
2
(18)
and allowed the possibility of using dose measurements or realistic Monte Carlo calculations to evaluate g– (r). As defined, g– (r) = gp(r). Late in the classical era, measured data were often used to evaluate g– (r), especially in the case of 125I seeds (Ling et al., 1983). More often, g– (r) was used to tabulate theoretical point-source data. Using this quantity, equation (12) becomes: wat ( µ / ρ )air ⋅ T (1 cm ) D" wat (r ) = SK ⋅ en ⋅ g (r ) 2 r
(19)
leading to the equivalence g (r ) = T (r ) T (1 cm ) . In the early 125I literature (Krishnaswamy 1978; Ling et al., 1983), the correction T(1 cm) was ignored and dose was calculated by D" wat (r ) = Aapp ⋅
(Γ ) δ
X 2
r
⋅ fwat
⋅ g (r ) .
(20)
Relationship of Semiempirical Dose-Calculation Model to AAPM TG-43 Notation Prior to the first AAPM TG-43 report published in 1995 (Nath et al., 1995), a variety of semiempirical and empirical schemes were in use, some of which are described above. To promote uniformity in dose-calculation practice, and to promote the use of measured or Monte Carlo dose-rate estimations, TG-43 endorsed a common dose-calculation formalism derived from the Interstitial Collaborative Working Group recommendations (discussed in chapter 11) (Anderson et al., 1990). The TG-43 dose-calculation protocol is a
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formalism for representing a single-source 2-D dose-rate distribution about an actual source, in a standardized tabular format. It assumes that a sparse matrix of dose-rate values has been previously measured or calculated by some other methodology (measurement or Monte Carlo). In contrast, the semiempirical dose-calculation formalism is theoretical method (admittedly a somewhat simplistic one) for calculating absorbed dose distributions. Thus in principle, classically calculated dose-rate values can be used to specify the TG-43 dosimetric ratios. For an idealized unfiltered point source, the following equivalences are easy to derive: wat Λ = ( µ en / ρ )air ⋅ T (1 cm )
(21)
gP (r ) = gL (r ) = T (r ) T (1 cm) . Let us now consider the more complex problem of relating semiempirical dose calculations to RDF · and Λ data derived from measurements or Monte Carlo estimates of dose rate, D (r,π /2), about actual sources having an effective active length L>0. If we assume that (i) the line-source geometry function accurately approximates the falloff of dose rate in free space outside the source and that (ii) the actual source spectrum in free space closely matches that assumed by equation (21), one would expect that the following approximations to apply with a high degree of accuracy. wat ( µ en / ρ )air ⋅ T (1 cm ) ⋅ GL (1 cm, π / 2 ) ≈ Λ ≡ D" (1 cm, π / 2 ) SK D" (r, π / 2 ) ⋅ GL (1 cm, π / 2 ) T (r ) T (1 cm ) ≈ gL (r ) ≡ D" (1 cm, π / 2 ) ⋅ GL (r, π / 2 )
(22)
The geometry correction factor, GL(1 cm,π/2), to the semiempirical approximation to the measured dose-rate constant can be quite significant for intracavitary tubes and other linear sources, deviating from unity by as much as 20%.
Application of Semiempirical Model to Practical Sources Modeling of Source Anisotropy: The Anisotropy Factor Despite its simplicity, the semiempirical isotropic point source model, equation (12), accurately predicts the transverse-axis dose-rate distributions of most actual radium-substitute sources. Simply by using an output quantity to calibrate the source, rather than A, the influence of its internal structure (filtration and self-absorption) has been implicitly accounted for. Had A instead of Aapp been used in equation (13)(b), then the expression for (Γ δ)X [equation (4)], would require correction for attenuation and scattering in the N radioactive core and surrounding encapsulation. Any uncertainties in {Ei , Pi }i =1 (which are large for many radionuclides) and filtration corrections would be directly degrade dose calculation accuracy. In addition, fundamental activity measurements are technically difficult for the high-intensity sources used in brachytherapy. For this reason, contained activity does not play a role in photon brachytherapy dosimetry. In contrast, equation (12) infers dose rate from a quantity measured outside the source, which is not influenced significantly influence by knowledge of the unfiltered photon spectrum. The required quantities, ( µ en / ρ )air and T(r), are ratios and are therefore largely insensitive to uncertainties in the assumed med
spectrum.
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Practically all brachytherapy sources are cylindrically symmetric, giving rise to anisotropic dose distributions. Additionally, some sources used in intracavitary brachytherapy have active lengths that are · comparable to typical calculation distances. Thus the dose rate, D (r, θ ), around a brachytherapy source · depends both on distance r, and polar angle, θ D (r, θ ) may deviate significantly from the transverse-axis · dose rate, D (r,π /2), predicted by equation (12) especially near the long axis of the source. In the case of implants consisting of many randomly oriented seeds with active lengths less than the minimum distance of interest, equation (12) will accurately represent the multiple seed dose distribution if an average correction for single-seed dose anisotropy is applied (Williamson 1986). This correction factor, called the anisotropy factor (1-D anisotropy function by the 2004 TG-43 update (Rivard et al., 2004), φan(r), is defined by averaging the dose at each fixed distance r with respect to solid angle, Ω:
φ an (r ) =
Average dose at r Transverse-axis dose at r
∫ =
4π
D" (r, θ ) ⋅ dΩ
4π D" (r, π / 2 )
∫ ⋅=
π
0
D" (r, θ ) ⋅ sinθ ⋅ dθ 2 ⋅ D" (r, π / 2 )
).
(23)
– Often a distance-independent average value of ƒan(r), called the “anisotropy constant,”φ an , is used. Incorporating this average correction into equation (12), leads to: wat S ⋅ ( µ en. / ρ )air D" wat (r ) = K ⋅ T (r ) ⋅ φ an r2
(24)
– Classically, φ an was evaluated by measuring relative photon fluence in air at relatively large distances (30 to 100 cm) using a NaI or GeLi scintillation detector (Ling, Anderson, and Shipley 1979; Ling et al., 1983). According to equation (24), source strength should be increased by a constant fraction ranging from 2% (192Ir seeds) (Nath et al., 1995) to 14% (103Pd seeds) to correct for polar anisotropy effects (Rivard et al., 2004). Lindsay et al. (Lindsay, Battista, and Van Dyk 2001) compared prostate implant 3-D dose distri· butions derived from the isotropic point-source model, D (r), to those derived from the full 2-D · single-source dose-calculation model, D (r, θ ). Based on voxel-by-voxel comparisons, they found that the isotropic point-source model introduced errors exceeding 10% of the D90 (see the chapter on interstitial implant dose specification) in 8% and 33% of the target volume for the Model 6711 125I and Model 200 103 Pd sources. Corbett et al. (2001) found that despite local changes to the dose distribution, inclusion of 2-D anisotropy effects did not alter the dose-volume histogram (DVH): neither the V100 nor the margin between D100 and prostate boundary were significantly altered. For volume implants consisting of parallel arrays of 192Ir seeds, Williamson (1986) reported a similar finding. Dose Calculation for Extended Sources: The 1-D Pathlength (Sievert Integral) Model Dose distributions around larger sources, such as intracavitary tubes and interstitial needles, are calculated by partitioning the extended source into a set of point sources to which corrections for distance, oblique filtration, attenuation, and scattering are separately applied. By summing these point source contributions, the dose at point P can be estimated. This class of algorithms, first described by Rolf Sievert in 1921 (Sievert 1921), is known as the 1-D pathlength model, because filtration and scatter corrections depend only on the distance between each source segment and the point of calculation. The term “Sievert integral,” which refers to a specific elliptic integral to which a 1-D integration of the point-dose kernel can be reduced, is often used to describe this class of algorithms.
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Consider the source of strength SK and contained activity A illustrated in Figure 4. The classical Sievert model approximates the cylindrical active core by a line of radioactivity positioned along its axis. The axial length of the core is called the active length, L. Oblique filtration is modeled by assuming that the capsule reduces dose by exponential attenuation using an effective filtration coefficient, µ′. The dose rate ∆D" ( x, y ) at point (x,y) from the incremental source ∆L located at angle θ is ∆D" ( x, y ) = A ⋅
∆L (Γ δ ) X ⋅ fmed ⋅ ⋅ T ( x / cos θ ) ⋅ e− µ '⋅t / cos θ , 2 L ( x / cos θ )
(25)
where is the exposure rate constant of the unfiltered source material. Since SK = A · (W/e) · (GÎ)X · e–µ¢t, equation (25) becomes: ∆D" ( x, y ) = SK ⋅
∆L L
⋅ e µ '⋅t ⋅
med ( µ en / ρ )air
( x / cos θ )2
⋅ T ( x / cos θ ) ⋅ e− µ '⋅t / cos θ .
(26)
By integrating with respect to Œ, and transforming to polar coordinates, we obtain the Sievert integral: SK ⋅ ( µ en / ρ )air ⋅ e µ ' t med
D" (r, θ ) =
L ⋅ r ⋅ cos θ
∫
θ2
θ1
e− µ ' t ⋅ssec θ ⋅T ( x ⋅ sec θ ) ⋅ dθ
(27)
The extra eº′·t term outside the integral is needed to avoid global “double correction” for filtration. Corrections for attenuation and scatter were first included in the model by Laurence (1935). The Sievert integral analytic solution was generalized to include point-kernel build-up factors by Batho and Young
Figure 4. A typical encapsulated line source, illustrating calculation of dose rate at point P at (x,y) relative to the source center by the Sievert integral method. L and t denote the active length and radial encapsulation thickness, respectively. The distances x and y are referred to as “distance away” and “distance along,” respectively in the literature.
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Jeffrey F. Williamson
(1964), who also devised the numerical solution of (27) used on most early treatment planning computers (Young and Batho 1964).
Clinical Application of Semiempirical Models Implementation of Point-Source Models for 192Ir and 137Cs Seed Sources Since the introduction of computer-based treatment planning, the semiempirical point-source model with average correction for anisotropy, equation (24), has been used for evaluation of interstitial brachytherapy dose distributions for all types of high-energy interstitial seed sources, including 226Rn, 198Au, 60Co, 137 Cs, and 192Ir. Of these, only 192Ir seeds, available as interstitial seeds encapsulated in nylon ribbon for low dose-rate (LDR) brachytherapy and in the form of a cable-driven single-stepping source for high dose-rate (HDR) brachytherapy remain widely used in North America. A few centers continue to use 137 Cs spherical pellets in the Nucletron Selectron remote afterloading system (Grigsby, Williamson, and Perez 1992). The published literature suggests that the semiempirical point-source model is has an overall accuracy similar to that of Monte Carlo simulation for 192Ir and 137Cs seed-like (those with an active length of 1 cm or less) sources for calculation distances up to 5 cm. Most commonly, Monte Carlo simulation has been used to benchmark the semiempirical models in terms of accuracy, often using equations (22). For 137 Cs sources, Williamson (1988, 1998; Williamson and Seminoff 1987) noted that widely T(r) tabulations (Meisberger, Keller, and Shalek 1968; Berger 1968) agreed within 2% of Monte Carlo simulations, both for idealized point sources and transverse-axis dose-rate distributions for a wide range of intracavitary tube sources based on actual source geometry. One exception was a heavily-filtered 137Cs tube with a gold-wire active core (Williamson 1988), for which the large magnitude of oblique filtration resulted in breakdown of the line-source geometry function as a predictor of dose rate in free space. A more recent analysis (Perez-Calatayud 2004) demonstrates for a wide range of intracavitary brachytherapy sources and published Monte Carlo dosimetry studies that (i) geometry-function compensated dose-rate constants and RDFs are essentially independent of source design and (ii) semi-empirical approximations derived from Meisberger’s T(r) data closely approximate Monte Carlo-based dose transverse-axis dose distributions. Similar findings support use of classical semi-empirical models for 192Ir sources. Williamson (1991, 1996) noted that Monte Carlo simulation and semiempirical approximations yielded nearly identical doserate constants and RDFs for LDR 192Ir seeds. Wang and Sloboda (1998a) systematically investigated the dependence of the transverse-axis dose distribution on source geometry and filtration. They concluded that assumptions (i) and (ii) given in the paragraph above are valid within 2% to 3% for a wide range of HDR and LDR sources. However, especially for 192Ir, it is well known that RDFs are sensitive to phantom size (and hence minimum distance from tissue-air interfaces in a clinical implant) especially for distances over 5 cm (Meisberger, Keller, and Shalek 1968; Perez-Calatayud, Granero, and Ballester 2004; Williamson 1967). In summary, for 192Ir and 137Cs sources, the Meisberger T(r) data continue to provide an accurate approximation to transverse-axis dose-rate distribution provided tissue-air boundaries are 7 to 10 cm away from the source. 1-D anisotropy functions and anisotropy constants should be derived from angular doserate distributions at radial distances near 1 cm in condensed medium. Either measurements about the actual source or Monte Carlo simulations based upon an accurate 3-D geometric model of the source are acceptable. Monte Carlo-derived anisotropy functions, and other TG-43 parameters, are available for a variety of HDR (Daskalov, Loffler, and Williamson 1998; Perez-Calatayud et al. 2001; Wang and Sloboda 1998b; Williamson and Li 1995) and LDR (Ballester et al., 2004) sources.
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Implementation of 1-D Pathlength Models for 137Cs Intracavitary Tubes Most recently analyses of the 1-D pathlength model generalize equation (27) to a 3-D spatial integration over spatial volume. This formulation, first proposed by Williamson (19988), will be described here. Consider a source (Figure 5) in which the radionuclide is uniformly distributed #over the inner cylinder (active source), which has an active length, L, and a radius, S. The dose rate, D" (r ) , in cGy/h at location # r is given by:
(
SK ⋅ µ en / ρ # D" (r ) = # F (rc )
)
wat air
3 $ $ $ −2 ⋅ ∫ ρ (r ')( r − r ' ) ⋅ exp − ∑ µ j ⋅ λ ' j ⋅ [1 + SPR(λ ' 3 )] ⋅ dV V' V
j =1
(28)
$ where V (cm3) denotes space enclosed by the active source core and ρ (r ') is the relative radioactivity density (cm–3). The indices j = 1, 3 denote the media composing the active source core, the source encapsulation, and surrounding medium while λ′1, λ′2, and λ′3 denote the corresponding distances traversed by # primary photons passing from dV′ to r. The other symbols are defined as follows: SK = is the air-kerma strength of the source (in units of cGy·cm2·h-1), specifically the product of the air# # kerma rate in free space, K" air (rc ) , at the reference point rc and the square of the distance: # # 2 SK = K" air (rc )⋅ | rc | . # 2 # $ $ $ −2 F (rc ) = rc ⋅ ∫ ρ (r ') ( rc − r ' ) ⋅ exp ( − µ1 ⋅ S − µ 2 ⋅ t ) ⋅ dV ' is the self-absorption/filtration correction V for the reference air-kerma rate specification geometry (source immersed in vacuum). µj = filtration coefficients of the active source and filter media for j = 1 and 2. In all of the calculations described in this report, µj is treated as a thickness-independent parameter. For the case of the surrounding medium (assumed to be liquid water), j =3, µ3 is the average linear attenuation coefficient of water. Published tabulations often specify µj in units of mm–1. $ Usually, ρ (r ') is assumed to be constant, in which case its occurrences in the numerator and denominator of (28) cancel. A major ambiguity in implementing the 1-D pathlength model is specification of required filtration coefficients, µj. The literature describes at least four physical interpretations of the filtration coefficient: (1) The µj describe transmission of γ rays through cylindrical filters of various thicknesses, t. This is the classical experimental interpretation (Shalek and Stovall 1969) proposed by Keyser (1951) and Whyte (1955), who derived µj for 226Ra from γ-ray transmission measurements through cylindrical filters of various thicknesses in a free-air geometry (see Figure 6). (2) The µj are well approximated by the average air-kerma-weighted linear energy absorption coefficient, µen,j of the corresponding material (BIR/IPSM 1993; Krishnaswamy 1972,1978). (3) µ1 and µ2, are simply parameters of best fit, adjusted to optimize the agreement between the 1-D pathlength model predictions and an independent estimate of the dose-rate distribution. This interpretation was first introduced by Diffey (Diffey and Klevenhagen 1975) for optimizing agreement with measured dose-distribution data and by Williamson (1988) for matching Monte Carlo data. (4) The µj describe transmission in a near narrow-beam absorber geometry (Figure 6) and are given approximately by the mean linear attenuation coefficient of the material. This interpretation was introduced by Williamson (1996) to implement a model that combined the 1-D pathlength and scat-
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Jeffrey F. Williamson
Figure 5. Simplified geometry of a cylindrically symmetric encapsulated source, illustrating the 3-D integration elements and ray tracing used to implement the 1-D pathlength algorithm. (Reproduced from Int J Radiat Oncol Biol Phys, vol. 36, “The Sievert integral revisited: Evaluation and extension to low energy brachytherapy sources,” J. F. Williamson, pp. 1239–1250. © 1996, with permission from Elsevier.)
Figure 6. Illustration of measurement geometry needed to realize the cylindrical transmission (Left: Interpretation (1) in text) and the narrow-beam contact absorber geometry (Right, Interpretation (4) in text). (Reproduced from Int J Radiat Oncol Biol Phys, vol. 36, “The Sievert integral revisited: Evaluation and extension to low energy brachytherapy sources,” J. F. Williamson, pp. 1239–1250. © 1996, with permission from Elsevier.)
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ter-separation (Williamson, Li, and Wong 1993) methods. Experimentally, µj describes transmission through a small block of filter material with a length equal to the active length and width equal to the physical diameter of the source. Figure 6 illustrates interpretations (1) and (4), while Figure 7 shows the substantial variation in filtration correction vs. thickness for these different interpretations. In the case of 137Cs γ rays filtered by stainless steel, interpretations (4), (1), and (2), yield µ values of 0.057 mm–1, 0.021 mm–1, and 0.0226 mm–1, respectively (Williamson 1988, 1998). Many commercial implementations of the 1-D pathlength model for interstitial needle or intracavitary tube sources impose various geometric approximation on the user, as illustrated by Figure 8 (right panel). The classical Sievert model was a 1-D line integral that used analytical approximations rather than exact ray tracing through an accurate 3-D geometric model of the source. A common approximation is a ceramic or steel line source embedded in a steel cylinder of diameter 2t. The dosimetric accuracy of the 1-D pathlength model and the optimal choice of filtration coefficients has been documented in a series of published papers by the senior author (Williamson 1988, 1998; Williamson and Seminoff 1987). For the two most commonly used intracavitary tubes (see Figure 8, left panel), 1-D pathlength model dose predictions were tested against more rigorous Monte Carlo-based dose calculations over the range of 0.25 cm to 7 cm away-and-along distances (Williamson 1998). Variants of the 1-D pathlength model tested included the µj ª µen,j approximation (interpretation (2)); treating µj as a fitting parameter (interpretation (3)); 3-D numerical integration over an accurate 3-D model of the source; and 1-D numerical integration over a simplified filtered line-source approximation to the source geometry. The results are shown in Table 1 and graphically in Figure 9 (Amersham, CDCS.J source) and Figure
Figure 7. Transmission, defined as the ratio of air-kerma from 192Ir g-rays in free space with a steel absorber in place to that emitted by an HDR brachytherapy source alone as a function of absorber thickness. Air-kerma is specified 25 cm from the source center on its transverse axis. Cylindrical and contact absorber transmission were calculated by Monte Carlo simulation, along with thickness-dependent air-kerma weighted mean linear attenuation and energy-absorption coefficients. (Reproduced from Int J Radiat Oncol Biol Phys, vol. 36, “The Sievert integral revisited: Evaluation and extension to low energy brachytherapy sources,” J. F. Williamson, pp. 1239–1250. © 1996, with permission from Elsevier.)
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Figure 8. Left: Drawings of the Amersham Model CSCS.J source (1992 version) and the 3M 6500/6D6C design. The active core of the former consists of 9 1.45 mm diameter borosilicate glass beads while the latter consists of zirconium phosphate glass microspheres. The packed array of microspheres was simulated as a solid cylindrical core of reduced density. Both sources are encapsulated in stainless steel. More details are given in Williamson (1998), consisting of a uniform glass cylinder placed symmetrically between source tips, as described in the vendor’s literature. Right: Illustration of three different geometric approximations in modeling the CDCS.J source dose distribution via the 1-D pathlength model.
10 (3M Model 6500 source). Generally, when the full 3-D geometry of the source (or at least a double concentric cylinder approximation) are used and best fit values of µj for steel and ceramic employed, the root mean-square (RMS) error is less than 1% and the maximum error no more than 4%. When the µj ≈ µen,j approximation along with a 3-D geometric model, the RMS error increases to about 3% with maximum errors of about 8% localized near the longitudinal axis of the source (see Figures 9 and 10). When the source geometry is approximated by a line-source on the axis of a cylindrical filter, this approximation dominates dose prediction error, so that regardless of the choice of filtration coefficients, RMS and maximum errors of 3% and 10%, respectively, are observed. In summary, the 3-D generalization of the classical Sievert-like 1-D pathlength dose-calculation algorithm estimates the 2-D dose-rate distribution about 137Cs steel-clad sources with low density cores, provided that the best filtration coefficients reproduced in Table 2 are used. This methodology can be used with confidence in place of Monte Carlo simulation or measurement with this class of sources. The author recommends that new sources be examined by means of both autoradiography and contact transmission radiography to verify the source geometry. Before relying upon a Sievert or 1-D pathlength algorithm clinically, it is recommended that the physicist first reproduce one of the published tables (Williamson 1998) to assess the adequacy of the algorithm implementation and to ensure that the input parameters are properly interpreted.
Application of 1-D Pathlength Models to Other Radionuclides Application of analytic models to predict dose distributions about sources that emit lower-energy photon spectra than 137Cs, is a topic of significant interest. Recent papers have focused on HDR 192Ir sources, of which several models are commercially available. Williamson (1996) evaluated the accuracy of the 1-D pathlength model, equation(28), against Monte Carlo simulation benchmarks for two 192Ir source models, a 169Yb seed, and the Model 6702 125I source for each of the four filtration coefficient evaluation methods
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Figure 9. Spatial distribution of fractional deviations of the 1-D pathlength model calculations from Monte Carlo dose-rate calculations for the Amersham CDCS.J source (1992 version, measured geometry). (a) 3-D geometry, approximating filtration coefficients by men. (b) 3-D geometry, approximating filtration coefficients by best fit values. (c) Ceramic line source model approximating filtration coefficients by men. (d) Ceramic line source model approximating filtration coefficients by best-fit values. (Reprinted from Int J Radiat Oncol Biol Phys, vol. 41. “Monte Carlo-based dose-rate tables for the Amersham CDCS.J and 3M model 6500 137Cs tubes,” J. F. Williamson, pp. 959–70. © 1998, with permission from Elsevier.)
described above. For the Nucletron “Classic” HDR 192Ir source (see Figure 11), the “best fit” and µj ª µen,j methods yielded RMS errors of 3.2% and 6.9%, respectively, and maximum errors as large as 20%. The best-fit filtration coefficients were found to vary significantly with source construction. Similar results for 192 Ir sources were found by other investigators (Baltas 1998; Cho and Muller-Runkel 1997). For 169Yb and 125 I sources, Williamson (1996) found RMS dose-prediction errors of 8% to 17%, and maximum errors as large as 46%, for conventional 1-D pathlength algorithm implementations.
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Jeffrey F. Williamson
Figure 10. Dose-rate as a function of polar angle at a fixed 2 cm distance from the active core center of a 3M Model 6500 source. Data derived from MCPT simulation as well as several implementations of the 1-D pathlength model are shown. (Reprinted from Int J Radiat Oncol Biol Phys, vol. 41. “Monte Carlo-based dose-rate tables for the Amersham CDCS.J and 3M model 6500 137Cs tubes,” J. F. Williamson, pp. 959–70. © 1998, with permission from Elsevier.)
Table 1: Accuracy of various 1D pathlength dose-calculation algorithms, relative to Monte Carlo simulation, for the 3M Model 6500 and Amersham Model CDCS.J intracavitary 137Cs tubes Source
Model
%RMS Error
%Error Range
CDCS.J (1992 Model)
3D (best fit) 3D (µen) Cylindrical core (best fit) Ceramic line (best fit) Ceramic line (µen)
0.7% 2.8% 0.9% 2.7% 3.0%
–1.7% to 4.2% 0.3% to 8.1% –1.4% to 4.0% –16.0% to 5.2% –5.4% to 10.1%
3D (best fit) 3D (µen) Ceramic line (best fit) Ceramic line (µen)
0.7% 2.8% 3.2% 2.7%
–1.8% to 3.6% –1.2% to 7.6% –14.4% to 3.7% –5.2% to 9.2%
3M 6500
Clearly, the classical Sievert/1-D pathlength algorithms fail to deliver dose-prediction accuracy sufficient for clinical needs for radionuclides with mean photon energies below that of 137Cs. The mechanism of these errors is well known (Williamson 1988): the 1-D pathlength algorithm assumes that both primary and scattered-photon dose are perturbed by a factor that depends only on the thickness of source and filtration material traversed by primary photons. This approximation can be valid only if the scatter-toprimary ratio, SPR(r), is independent of polar angle. Even for 137Cs this assumption is false—because
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Table 2: Optimal filtration coefficients for steel-clad ceramic-core 137Cs Tubes Source Material
Linear energy absorption coefficient
Best fit filtration coefficient
Stainless steel - 3M6500 - CDCS-J
0.0226 mm–1 0.0226 mm–1
0.036 mm–1 0.037 mm–1
Source Core - 3M6500 (cylinder: 2.22 g/cc) - CDCS-J (spherical: 2.60 g/cc) - CDCS-J (Cylindrical: 1.73 g/cc)
0.0065 mm–1 0.0076 mm–1 0.0051 mm–1
0.0070 mm–1 (0.032 cm2 g–1) ————— 0.0046 mm–1 (0.027 cm2·g–1)
scattered-photon dose is approximately isotropically distributed about cylindrical sources (Williamson 1988, 1990), the SPR at fixed distance rapidly increases as the polar angle approaches zero. Based on the insight, that scatter dose is approximately isotropically distributed, Williamson (1996) introduced a revised 1-D pathlength model, called the “isotropic scattering model,” which is based upon separating the scattered and primary dose components. D" ( r, θ ) = D" pri (r, θ ) + D" sca (r, π / 2 ) ⋅ C ( r, θ )
(29)
= D" pri (r, θ ) + D" pri (r, π / 2 ) ⋅ C ( r, θ ) ⋅ SPR(r ) · The primary dose rate, D pri (r, θ ), is estimated by the conventional 1-D pathlength model, equation (28), using attenuation coefficients that approximate narrow beam attenuation, interpretation (4) from above. C(r,θ) is an empirical correction that accounts for deviations of the scatter-dose distribution from isotropy, introduced by Karaiskos et al. (2000) to improve the predictive accuracy of the model: Williamson assumed C(r,θ) = 1. As illustrated by Figure 11, the isotropic scattering model improved the predictive accuracy slightly for the HDR 192Ir source, yielding RMS and maximum errors of 2.9% and 8.9%, respectively. An advantage of equation (29) is that its input parameters are independent of the source design and in principle, can be specified without knowledge of the source dose distribution. The isotropic scattering model significantly improved predictive accuracy for lower energy sources as well, producing an impressive 6.6% RMS error (14% maximum error) for the Model 6702 seed. As noted, Karaiskos et al. (2000) improved the isotropic scattering model, for 192Ir sources only, by simplifying evaluation of the narrow-beam filtration coefficients and introducing a source-geometry independent scatter anisotropy correction, C(r,θ) = C(θ) = 1–(3×109) · (θ–90°). With these improvements, these authors obtained an RMS and maximum errors of 1.5% and 6%, respectively, for the Nucletron and VariSource HDR sources as well as LDR seeds and wires (Karaiskos et al., 2000; Pantelis et al., 2002). In summary, classical Sievert-like models should not be applied clinically to 192Ir or lower energy sources. The isotropic scattering model, as formulated by Karaiskos et al. (2000), is an alternative at least for 192Ir sources. However, this model is not available on commercial planning systems and should be used cautiously outside the domain in which it has been validated.
Application of 1-D Pathlength Models to Internally Shielded Applicators Internally shielded colpostats are widely used in intracavitary brachytherapy to spare bladder and rectal tissue anterior and posterior to the vaginal vault. Typically, such applicators contain 2 to 5 mm of lead or
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tungsten-alloy shielding in the medial aspects of the anterior and posterior applicator surfaces and can be produce dose reductions as large as 50% relative to the unshielded source dose distribution (Williamson 1990). Most commercially available planning systems neglect internal applicator shielding, resulting in systematic overestimates of bladder and rectal doses. The 1-D pathlength algorithm was first generalized to accommodate applicator shielding by Meertens and Van der Laarse (1985). The algorithm ray-tracing capability must be extended to the applicator materials. In clinical applications, this means the treatmentplanning system must have the capability of representing the applicator geometry and reconstructing its 3-D dose distributions in patients from planar radiographs or CT images (Lerma and Williamson 2002). Effective attenuation coefficients are then evaluated by optimizing the agreement between model predictions and measurements (or Monte Carlo calculations) about the shielded applicator. Other investigators who have contributed to these developments include Williamson (1990), who developed a scatter separation model for estimating dose around shielded applicators and Weeks (weeks 1998; Weeks and Dennett 1990), who developed a method of extracting optimal filtration coefficient values from Monte Carlo simulations of the applicator geometry. An example of the excellent agreement that can be achieved for 137 Cs-bearing shielded colpostats is shown in Figure 12. Few publications have assessed the accuracy of 1-D pathlength algorithms for predicting dose distributions about shielded applicators for lower energy sources. Lymperopoulou et al. (2004) used Monte Carlo simulation to evaluate the accuracy of the Nucletron/Plato implementation of the 1-D pathlength algorithm for a segmentally shielded HDR 192Ir vaginal/rectal applicator. The 1-D pathlength algorithm yielded acceptable accuracy at the prescription distance, but had dose-prediction errors as large as 15% at larger distances. Based on the high scatter-toprimary ratio of lower energy sources, and the substantial variation of attenuation corrections on shield cross-sectional area and location (Williamson et al., 1993), one would expect accuracy and generality of 1-D pathlength heterogeneity corrections to be quite limited.
Advanced Analytic Dose Calculation Algorithms A number of more sophisticated and general dose-calculation algorithms have been described in the literature. One promising approach is the 2-D scatter-integration algorithm developed by Kirov and Williamson
Figure 12. Dose rate (cGy/h) predicted by Weeks’ implementation (Weeks 1998) of the 1-D pathlength model (broken line) and as estimated by MCNP Monte Carlo calculations (solid line) for a CT-compatible Fletcher-Suit colpostats. The left and right panels illustrate good agreement between the two methods in planes normal to the source axis 2 cm anterior and posterior, respectively, of the applicator center. (Reprinted with permission from Med Phys, vol. 25, “Monte Carlo dose calculations for a new ovoid shield system for carcinoma of the uterine cervix,” K. J. Weeks, pp. 2288–2292. © 1998, with permission of AAPM.)
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(1997). Like the scatter subtraction method (Williamson, Li, and Wong 1993) of which it is a generalization, it is based upon a 2-D SPR table, SPR(r,θ), giving the scatter-to-primary ratio along the axis of an isotropic point-source collimated to a cone of half angle θ. By performing a Clarkson-like 2-D scatter integration, these authors were able to predict the 2-D dose distributions about HDR 192Ir and Model 6702 125I sources with maximum errors of 5%. This is equivalent to Karaiskos’ results (Karaislos et al., 2000) but without having to use an empirical scatter anisotropy correction. In principle, this method could be extended to more complex applicator and tissue heterogeneity problems. A more general approach is the superposition/convolution algorithm (Carlson and Ahnesjo 2000; Tedgren and ahnesjo 2003; Williamson, Baker, and Li 1991). In analogy to its external-beam counterpart, the superposition algorithm defines a 2-D scatter dose array that gives absorbed dose at any point from a colliding primary photon, in a coordinate system centered on the primary collision site and oriented along the primary photon trajectory. This dose-spread array is convolved with the primary TERMA distribution to give an estimate of the total scatter dose. Various methods have been proposed for scaling the scatterdose kernal to account for heterogeneities, to correct for the severe kernel tilt characteristic of brachytherapy, and to manage singularities in the convolution integration domain. Preliminary results suggest superposition is quite accurate but computationally complex. Another and more general approach to brachytherapy dose computation is use of deterministic solutions of the Boltzmann transport equation. Daskalov et al. (2000, 2002) have proposed using discrete ordinates solutions while Zhou et al. have suggested integral solutions of the transport equation (Zhou and Inanc 2003). These solutions approach the complexity, rigor, and computational efficiency of Monte Carlo simulation. Daskalov et al. (2002) have shown that single-processor 2-D discrete ordinates calculations are 1 to 2 orders of magnitude faster than EGS4 calculations of equivalent dimensionality and that single-source 3-D calculations using the PARTISN discrete ordinates code maybe as much as an order of magnitude faster than PTRAN Monte Carlo calculations (Williamson et al., 2001). One advantage of Monte Carlo calculations is that calculations can be cost-effectively parallelized using PC clusters with distributed memory while parallelizing discrete ordinates calculations requires much more expensive shared-memory multiple processor hardware.
Semiempirical Models in Brachytherapy Quality Assurance In addition to supporting computer-assisted treatment planning, semiempirical dose-calculation tools have significant value as a manual calculation aid, either for directly estimating treatment times for certain types of clinical implants or as tools for verifying the accuracy of computerized dose calculations. A particularly useful form of the classical Sievert equation (27) can be obtained by setting encapsulation thickness to zero (t = 0) and tissue attenuation-scatter to unity. Then, equation (27) reduces to a simple closed-form analytic expression: med ∆θ D" ( x, y ) = SK ⋅ ( µ en / ρ )air ⋅ L⋅x
(30)
where ∆θ is the angle, in radians, subtended by the active length, L, with respect to the point of interest (see Figure 13). When the interest point lies on the transverse axis (y = 0), then ∆θ = 2·tan–1 (L/2x), where tan–1 means inverse tangent or arctangent. Angles must be specified in units of radians rather than degrees. This approximation is extremely useful as a manual calculation aid and is highly accurate near the transverse axis of lightly encapsulated 137Cs sources or linear arrays of 192Ir seeds.
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Figure 13. Error in the isotropic point source-model relative to the line source model equation (30) as a function of transverse-axis distance expressed in multiples of active length.
Figure 13 shows that as the distance r = ( x 2 + y 2 ) becomes large in relation to active length, L, equation (30) reduces to the point-source equation (12), with T(r) = 1.0. For distances less than L (~1.5 cm for intracavitary tubes), use of equation (12) will yield errors of at least 10%. For distances greater than 1.5 L (2.0 to 2.5 cm for gynecologic tubes), the point-source approximation is accurate within 5%.
Example: Manual Calculation of Dwell Time for an HDR Vaginal Cuff Insertion Problem: Small (2 cm diameter) Fletcher colpostats are used to deliver 600 cGy to the vaginal apex at depth of 5 mm via a HDR 192Ir insertion as illustrated by Figure 14. Assume that the colpostats centers are separated by 2.5 cm and that the source SK = 2.50 cGy ∑m2 ∑h–1 at the time of treatment. Further assume that three dwell positions are activated uniformly and that the spacing is 5 mm. Calculate the dwell time/position needed to deliver the prescribed dose. We approach this problem in a stepwise fashion: · 1. Estimate the dose-rate, D (1 cm,π/2), at 1 cm distance. Using equation (22) and noting T(1 cm)·GL(1 cm,p/2)ª1.0, we obtain
(31) wat D" (1 cm,π /2) = SK Λ = SK ⋅ ( µ en / ρ )air 2 2. Adapt problem= to (30). ⋅ h -1 ⋅ 2.5equation cGy ⋅ cm
10 4 cm 2 /m 2 2 -1 3600 s/h ⋅ 1.11 = 7.71 cGy ⋅ cm ⋅ s
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Figure 14. Left: Geometry for estimating dwell time needed to deliver specified dose to the vaginal apex using HDR intracavitary brachytherapy Fletcher Colpostats. Right: Method for evaluating the active length, L, the gives a dose distribution equivalent to that of three equally spaced dwell positions.
As illustrated by Figure 14 (right panel) and discussed in more detail elsewhere (Williamson and Brenner 2003), we see that each colpostat loading, consisting of three dwell positions, can be approximated by a line source with active length, L = 1.5 cm. The perpendicular distance, r, between each linear source center and the apex prescription point can be calculated by r = 1.5 2 + (2.5 / 2 )2 = 1.95 cm . 3. Set up equation (30).
D (r, π / 2 ) = T ⋅ SK ⋅ ( µ en / ρ )air ⋅ med
= 6 ⋅ t ⋅ D" (r, π / 2 ) ⋅
∆θ L⋅r
(32)
−1
2 ⋅ tan ( L / 2 r ) L⋅r
where total dwell time T = 6·t. 4. Substitute numerical values into equation (32) and solve for t.
−1
600 cGy = 6 ⋅ t ⋅ 7.71 ⋅
Conclusion
2 ⋅ tan ( 0.75 / 1.95 )
(33)
1.5 ⋅ 1.95 2 ⋅ 0.3672 = 11.62 ⋅ t = t ⋅ 46.3 ⋅ 2.925 ⇒ t = 51.6 s/positiion
As the reader will note, semiempirical dose-calculation methodologies have had a long and venerable history in clinical brachytherapy. For high-energy photon-emitting seeds and linear 137Cs sources, properly implemented semiempirical dose-calculation algorithms have a predictive accuracy comparable to Monte Carlo simulation and produce far more consistent and precise results than direct measurement. Straightforward extensions of these models enable 2-D dose distributions to be accurately evaluated about
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Jeffrey F. Williamson
Ir sources and shielded 137Cs-bearing applicators. Given the growing availability of Monte Carlo simulation, currently as a source of single-source TG-43 dose arrays and in the future as a patient-specific planning tool, what is the future of such algorithms? The author believes that semiempirical algorithms will continue to play an important role in LDR 137Cs-based brachytherapy and a role that deserves more consideration in HDR/LDR 192Ir-based brachytherapy for generating reference-quality single source dose arrays, whether they be packaged in the TG-43 formalism or in terms of classical dosimetry quantities. While Monte Carlo simulation can and should be used for this purpose, developing and publishing such tables is somewhat costly. Neither physicists nor journals have much motivation for pursuing such scientifically unrewarding work. Thus semiempirical dose calculations are cost-effective alternatives for locally generating dose tables using widely available software. Will Monte Carlo-based treatment planning make precalculated single-source arrays obsolete regardless of their origin? Several groups are developing accelerated Monte Carlo codes able to compute 3-D dose distributions in 23 mm3 voxels with computing times of 2 to 20 minutes on single PC or workstation processors. Thus, in principle, the answer is “yes”? However, with rare exceptions, planning system vendors have rarely made state-of-the-art brachytherapy software a priority. With the declining incidence of invasive cervix cancer in the developed world, it is by no means assured that commercially available Monte Carlo-based brachytherapy planning tools will be available for intracavitary brachytherapy in the foreseeable future. One important exception is low-energy seed implants for low-risk prostate cancer. As the number of permanent seed implants is continuing to rise worldwide, development of sophisticated but specialized image-guided intraoperative planning software packages continues. Incorporating Monte Carlo dose-calculation engines into such software packages is a logical step that could dramatically improve the accuracy of clinical dose estimation by rigorously accounting for tissue heterogeneities and interseed attenuation. However, before such dose-calculation tools can be effectively used, a number of problems must be solved, including extraction of seed orientation as well as location from CT images, suppression of streaking and other CT imaging artifacts, development of methods for assessing composition and density of soft-tissue heterogeneities. Thus, the author believes that for some time, classical dose-calculation methodologies will continue to play important roles in brachytherapy, both as planning and QA tools and as sources of reference-quality planning data in specific settings.
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Appendix : Derivation of Exposure Rate and Air-Kerma Rate Constants [Note: This appendix is adapted from Williamson and Brenner (2003) by kind permission of the publishers.]
The derivation of the exposure-rate constant, which relates contained activity of a point source to exposure rate given the photon spectrum, illustrates several fundamental dosimetry concepts. Consider an isotropic point source (Figure A1) of activity A (in mCi) in a vacuum surrounded by a thin spherical shell of radius r and thickness ∆r. The shell is assumed to contain air of density, ρ[g/cm2] where, by convention, the brackets specify the unit of preceding quantity. Further assume that the source emits P photons of energy E (in MeV) per disintegration. We first derive the fundamental relationship between photon fluence [defined by equation (7) in the text] at distance r, Φ(r) and absorbed dose, D(r). We assume that Φ (r) is known. By definition, the linear energy absorption coefficient, µen is given by:
µ en [cm -1 ] ≡
∆Eab Einc
⋅
1 ∆r
(A1)
where Einc is the photon energy incident upon the spherical surface of radius r and ∆Eab is the energy absorbed by the air in the spherical shell of thickness, ∆r. Since the source is isotropic, Einc(r) = 4πr2· Φ (r)·E. Substituting equation (2) and the expression for Einc(r) into the definition of absorbed dose, we obtain D (r ) ≡
∆Eab ∆m
=
Einc (r ) ⋅ µ en ⋅ ∆r 4π ⋅ r 2 ⋅ ∆r ⋅ ρ
= Φ(r ) ⋅ E ⋅ ( µ en ρ )
(A2)
where ∆m is the mass of the shell and (µen/ρ) is the mass energy absorption coefficient [cm2/g] for air medium evaluated at energy E. Note that all information regarding the density, thickness or geometry of the shell cancels out. As an aside, equation (3) describes the fundamental link between dosimetric quantities and more general radiation field descriptions. For radiation fields in which the photon trajectories # # have arbitrary angular and energy distributions as described by the angular fluence, Φ(r , Ω, E ) , equation (3) becomes # D (r ) =
∞
∫
# # # # Φ r , Ω , E ⋅ E ⋅ µ ρ r , E ⋅ d Ω ⋅ dE . ( ) ( ) ( ) en ∫
(A3)
0 4π
Returning to the problem at hand, the steps outlined in Figure A1 are followed to estimate exposure rate at distance r: " r ): Step 1: Calculate the particle flux (fluence rate), Φ( " r ) [ cm -2 ⋅ h -1 ] = A [ mCi] ⋅ P [ photons/dis] ⋅ 3.7 × 10 7 dis/s ⋅ 3600 s/h Φ( mCi 4π r 2
(A4)
" (r ) = 1.05997 × 1010 ⋅ A ⋅ P Φ r2 where r has units of cm. Step 2: Calculate the dose rate, D" air (r ) from the fundamental dose-fluence relationship, equation (A2).
228
Jeffrey F. Williamson -1 -1 " (r ) ⋅ E ⋅ ( µ ρ ) D" air ( r ) [ J ⋅ kg ⋅ h ] = Φ en
= 1.05997 × 10 ⋅ 10
A⋅P r
2
[ cm ⋅ h ] ⋅ E [ Mev] ⋅ ( µ en ρ ) [ cm /g ] 2
-1
× 1.602 × 10 =1.698 ⋅
A⋅P⋅E r
2
-13
2
[ J/MeV] ⋅ 1000 [ g/kg ]
(A5)
⋅ ( µ en ρ )
where “J” denotes the joule, the unit of energy. Step 3: Convert energy absorbed in air to ionization created and apply the definition of exposure. The energy imparted to air medium per unit ionization of like sign created is related by a constant, (W/e), which has a value 33.97 J/C. By definition, one unit of exposure = 1 R = 2.58×10–5 C/kg. Thus the exposure rate at distance r, Xδ (r ) , is: X" δ ( r ) [ R/h ] = 1.698 ⋅
A ⋅ P ⋅ E ⋅ ( µ en / ρ ) air
= 193.8 ⋅
r
2
1R J ⋅ 1C ⋅ kg ⋅ h 33.97 J 2.58 × 10 −4 C/kg
A ⋅ P ⋅ E ⋅ ( µ en / ρ ) air r
(A6)
2
2 Applying the definition of exposure rate constant (Γ δ ) X = X" δ (r ) ⋅ r / A , we finally obtain:
( Γ δ ) X = 193.8 ⋅ P ⋅ E ⋅ ( µ en / ρ )air
(A7)
Figure A1: Calculation of absorbed dose rate to air at distance r about an unencapsulated isotropic point source and air air-kerma rate constant,
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Generalizing this expression to a photon-emitting radionuclide of N different energies E1, E2,..., EN with corresponding emission rates (photons emitted per disintegration), P1, P2,..., PN, we obtain the final result: N
( Γ δ ) X = 193.8 ⋅
∑ P ⋅ E ⋅ (µ i
i
en
/ ρ )air , i
(A8)
i =1
where E has units of MeV and (Γ δ)X has units of R·cm2·mCi-1·h-1. Because the International System (SI) of units has replaced the quantity exposure with the quantity kerma, K, the ICRU has introduced the air-kerma rate constant, (Γ δ)K, to replace the exposure rate constant. Since K" δ (r ) = D" δ (r ) , we obtain from equation (A5)
(Γ δ )K =
K" δ (r ) ⋅ r 2 A
= 1.274 ⋅ 10
−15
N
⋅
∑ P ⋅ E [MeV] ⋅ (µ i
i
en
/ ρ )air , i [ cm 2 /g ]
(A9)
i =1
where (Γδ)K has units of Gy·m2·Bq-1·s-1.
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National Council on Radiation Protection and Measurements (NCRP). NCRP Report No. 5. “A Handbook of Radioactivity Measurements Procedures.” Bethesda, MD: NCRP, 1985. Pantelis, E., D. Baltas, K. Dardoufas, P. Karaiskos, P. Papagiannis, H. Rozaki-Mavrouli, and L. Sakelliou. (2002). “On the dosimetric accuracy of a Sievert integration model in the proximity of 192Ir HDR sources.” Int J Radiat Oncol Biol Phys 53(4):1071–1084. Parker, H. M., “A dosage system for interstitial radium therapy. II. Physical aspects.” Br J Radiol 11:252–266. Perez-Calatayud, J., F. Ballester, M. A. Serrano-Andres, V. Puchades, J.L. Lluch, Y. Limami, and E. Casal. (2001c). “Dosimetry characteristics of the Plus and 12i GammaMed PDR 192Ir sources.” Med Phys 28(12):2576–2585. Perez-Calatayud, J., D. Granero, and F. Ballester (2004). “Phantom size in brachytherapy source dosimetric studies.” Med Phys 31(7):2075–2081. Perez-Calatayud, J., D. Granero, F. Ballester, V. Puchades, and E. Casal. (2004). “Monte Carlo dosimetric characterization of the Cs-137 selectron/LDR source: Evaluation of applicator attenuation and superposition approximation effects.” Med Phys 31(3):493–499. Quimby, E. H. (1932). “The grouping of radium tubes in packs and plaques to produce the desired distribution of radiation.” Am J Roentgenol Radium Ther 27:18–39. Quimby, E. H. (1941). “The specification of dosage in radium therapy.” Am J Roentgenol Radiat Ther 45:1–18. Rivard, M. J., B. M. Coursey, L. A. DeWerd, W. F. Hanson, M. S. Huq, G. S. Ibbott, M. G. Mitch, R. Nath, and J. F. Williamson. “Update of AAPM Task Group No. 43 Report: A revised AAPM protocol for brachytherapy dose calculations.” Med Phys 31:633–674. Also available as AAPM Report No. 84. Roesch, W. C. (1958). “Dose for nonelectronic equilibrium conditions.” Radiat Res 9:399–410. Shalek, R. J., and M. Stovall. “Dosimetry in Implant Therapy” in Radiation Dosimetry, vol III. F. H. Attix and E. Tochlin (eds.). New York: Academic Press, pp. 776–798, 1969. Sievert, R. M. (1921). “Die Intensitatsverteilung der primaren: Strahlung in der Nahe medizinischer Radiumpraparate.” Acta Radiol 1:89–128. Stovall, M., and R. J. Shalek. (1972). “A review of computer techniques for dosimetry of interstitial and intracavitary radiotherapy.” Comput Programs Biomed 1:125. Tedgren, A. K., and A. Ahnesjo. (2003). “Accounting for high Z shields in brachytherapy using collapsed cone superposition for scatter dose calculation.” Med Phys 30(8):2206–2217. Van Kleffens, H. J., and W. M. Sta. (1979. “Application of stereo x-ray photogrammetry (SRM) in the determination of absorbed dose values during intracavitary radiation therapy.” Int J Radiat Oncol Biol Phys 5:557. Wang, R., and R. S. Sloboda. (1998a). “Influence of source geometry and materials on the transverse axis dosimetry of 192Ir brachytherapy sources.” Phys Med Biol 43(1):37–48. Wang, R., and R. S. Sloboda. (1998b). “Monte Carlo dosimetry of the VariSource high dose rate 192Ir source.” Med Phys 25(4):415–423. Weeks, K. J. (1998). “Monte Carlo dose calculations for a new ovoid shield system for carcinoma of the uterine cervix.” Med Phys 25(12):2288–2292. Weeks, K. J., and J. C. Dennett. (1990). “Dose calculation and measurements for a CT-compatible version of the Fletcher applicator.” Int J Radiat Oncol Biol Phys 18:1191–1198. Whyte, G. H. (1955). “Attenuation of g radiation in cylindrical geometry.” Br J Radiol 28:635. Williamson, J. F. (1986). “The accuracy of the line and point dose approximation in Ir-192 dosimetry.” Int J Radiat Oncol Biol Phys 12:409. Williamson, J. F. (1988). “Monte Carlo and analytic calculation of absorbed dose near 137Cs intracavitary sources.” Int J Radiat Oncol Biol Phys 15:227–237. Williamson, J. F. (1990). “Dose calculations about shielded gynecological colpostats.” Int J Radiat Oncol Biol Phys 19:167–178. Williamson, J. F. (1991). “Comparison of measured and calculated dose rates in water near I-125 and Ir-192 seeds.” Med Phys 28:776–786. Williamson, J. F. (1996). “The Sievert integral revisited: Evaluation and extension to low energy brachytherapy sources.” Int J Radiat Oncol Biol Phys 36:1239–1250. Williamson, J. F. (1998). “Monte Carlo-based dose-rate tables for the Amersham CDCS.J and 3M model 6500 137Cs tubes.” Int J Radiat Oncol Biol Phys 41(4):959–70.
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Williamson, J. F., and D. A. Brenner. “Physics and Radiobiology of Brachytherapy” in Principles and Practice of Radiation Oncology, 4rd Edition. C. A. Perez, L. W. Brady, E. Halperin et al. (eds.). Philadelphia: J.B. Lippincott Company, pp. 472–537, 2003. Williamson, J. F., and Z. Li. (1995). “Monte Carlo aided dosimetry of the microselectron pulsed and high dose-rate 192 Ir sources.” Med Phys 22(6):809–820. Williamson, J. F., R. Baker, and Z. Li. (1991). “A convolution algorithm for brachytherapy dose computations in heterogeneous geometries.” Med Phys 18:1256–1265. Williamson, J. F., Z. Li, and J. W. Wong. (1993). “One-dimensional scatter-subtraction method for brachytherapy dose calculation near bounded heterogeneities.” Med Phys 20:233–244. Williamson, J. F., B. M. Coursey, L. A. DeWerd, W. F. Hanson, R. Nath, M. J. Rivard, and G. Ibbott. (1999). “On the use of apparent activity (Aapp) for treatment planning of 125I and 103Pd interstitial brachytherapy sources: Recommendations of the AAPM Radiation Therapy Committee Subcommittee on low-energy brachytherapy source dosimetry.” Med Phys 26:2529–2530. Williamson, J., G. Daskalov, R. Baker, D. Rogers, and I. Kawrakow. (2001). “Accuracy and efficiency comparisons between three-dimensional multigroup discrete ordinates and voxel based Monte Carlo methods for dosimetric modeling of the Model 6702 125I seed.” (Abstract). Med Phys 28(6):1229. Williamson, J. F., and R. Nath. (1991). “Clinical implementation of AAPM Task Group 32 recommendations on brachytherapy source strength specification.” Med Phys 18:439–448. Williamson, J. F., H. Perera, Z. Li, and W. R. Lutz. (1993). “Comparison of calculated and measured heterogeneity correction factors for 125I, 137Cs and 192Ir brachytherapy sources near localized heterogeneities.” Med Phys 20:209–222. Williamson, J. F., and T. Seminoff. (1987). “Template-guided interstitial implants: 137Cs-reusable sources as a substitute for 192Ir.” Radiology 165:265–269. Wyckoff, H. O. “From ‘Quantify of Radiation’ and ‘Dose’ to ‘Exposure’ and ‘Absorbed Dose:’An Historical Review.” Report No. Lauriston Taylor Lecture 4, National Council on Radiation Protection and Measurements, Washington, DC, 1980. Young, M. E. J., and H. F. Batho. (1964). “Dose tables for linear radium sources calculated by an electronic computer.” Br J Radiol 37:38–44. Zhou, C., and F. Inanc. (2003). “Integral-transport-based deterministic brachytherapy dose calculations.” Phys Med Biol 48(1):73–93.
Chapter 15
Quantitative Dosimetry Methods for Brachytherapy Jeffrey F. Williamson, Ph.D.1 and Mark J. Rivard, Ph.D.2 1 Department of Radiation Oncology Virginia Commonwealth University School of Medicine, Richmond, Virginia 2 Department of Radiation Oncology, Tufts University School of Medicine Boston, Massachusetts Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 233 Brief History of Brachytherapy Dosimetry . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 235 226 Ra Dosimetry: From Threshold Erythema Dose to Semiempirical Models . . . . . . . . . . . . . . . . 235 Rise of Modern Experimental Dosimetry Methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 236 Rise of Modern Computational Dosimetry Approaches . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 236 Role of the AAPM in Standardizing Low-Energy Brachytherapy Dosimetry . . . . . . . . . . . . . . . . 237 Experimental Brachytherapy Dosimetry Methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 238 Reference-Quality Dose Measurement System Characteristics . . . . . . . . . . . . . . . . . . . . . . . . . . . 239 Positional Accuracy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 240 Signal-to-Noise Ratio, Dosimeter Readout Precision, and Signal Stability . . . . . . . . . . . . . . . . 240 Energy-Response and Other Artifacts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 243 General Formalism for Absorbed Dose Measurement . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 244 Brachytherapy Dose-Measurement Systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 248 Thermoluminescent Dosimetry (TLD) Techniques in Brachytherapy . . . . . . . . . . . . . . . . . . . . 248 Diode Dosimetry Techniques in Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 254 Plastic Scintillator Dosimetry for Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 256 Radiochromic Film Dosimetry for Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 258 Polymer Gel Dosimetry for Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 261 Computational Dosimetry Methods in Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 264 Monte Carlo-based Brachytherapy Dosimetry: General Considerations . . . . . . . . . . . . . . . . . . . . 265 Technical Aspects of Monte-Carlo Photon-Transport Simulation in Brachytherapy . . . . . . . . . . . 267 Physics of Photon Scattering and Choice of Cross-Section Library . . . . . . . . . . . . . . . . . . . . . 267 Geometric Modeling and Validation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 271 Choice of Estimator and Other Variance Reduction Techniques . . . . . . . . . . . . . . . . . . . . . . . . 273 Application of the Monte Carlo Method to Calculation of Reference-Quality Dose-Rate Distributions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 275 Uncertainty of Monte Carlo Reference Dose-Rate Distributions . . . . . . . . . . . . . . . . . . . . . . . . 280 Additional Recommendations for Use of Monte Carlo Simulation in Preparing Reference-Quality Dose-Rate Distributions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 281 Other Applications of Transport Codes in Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 281 Comparison of Theoretical and Experimental Dose-Rate Distributions for 125I and 103Pd Brachytherapy Sources . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 282 Scope of Seeds Included in Analysis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 282 125 I . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 282 103 Pd . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 283 Ratios of Monte Carlo and Measured Dose-Rate Distributions for 125I . . . . . . . . . . . . . . . . . . . . . 283 Ratios of Monte Carlo and Measured Dose-Rate Distributions for 103Pd . . . . . . . . . . . . . . . . . . . . 284 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 285
Introduction Computation of brachytherapy dose distributions for clinical treatment planning for a specified arrangement of sources is almost universally based upon the principle of source superposition. The dose rate,
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· D(r) , at each point r in the dose-calculation grid is the sum of individual source dose contributions, where each source is defined by its location ri and air-kerma strength SK,i N
D" (r ) = ∑ SK , i ⋅ D0 (r − ri )
(1)
i =1
The dose contribution from each source is a function only of its position ri relative to r and is given · by the single-source dose-rate distribution, D 0(r) which is defined as dose rate per unit SK from a single source with its center positioned at the origin. Prior to the publication of the original TG-43 report (Nath et al., 1995) in 1994, semiempirical or analytic models such as the Sievert integral were used to obtain · D 0(r). Current practice standards dictate a table-based approach. The TG-43 protocol (Nath et al., 1995; Rivrd et al., 2004) is essentially a standardized nomenclature for presenting sparse single-source dose arrays with an associated nonlinear interpolation scheme for estimating normalized dose rate between dose table entries. Clearly, any protocol based upon table lookup requires as input a table of measured or computed dose rates for each type of source. Dosimetry methods useful for estimating reference-quality dose-rate distributions needed to define · D 0(r) and other dosimetric input needed for clinical treatment constitute the main topic of this chapter. Reference-quality dosimetric data is intended for use in clinical treatment planning for sources of the type or model upon which the dose estimates were based. The AAPM recommends (Rivard et al., 2004; Williamson et al., 1998) that reference-quality data be derived from redundant independent dose estimates, be traceable to applicable NIST standards, and undergo a process of peer review before clinical acceptance by the community. This chapter will focus almost exclusively on photon-emitting sources. Although the main emphasis will be on estimation of single-source reference quality dose distributions, estimation of other types of dosimetric data, including direct use of quantitative dosimetry techniques in patient-specific treatment planning, will be briefly considered. Dosimetry is that branch of medical physics that concerns itself with the estimation of absorbed dose by means of experimental or fundamental theoretical techniques. An experimental dosimetry technique consists of a detector, known as a dosimeter, which produces a measurable signal that has a known relationship to absorbed dose at a specified point in the medium in the absence of the dosimeter. Well-defined procedures for dosimeter calibration, irradiation, readout and artifact correction are essential components of any experimental dosimetry method. Theoretical or computational dosimetry techniques must accurately model all physical phenomena that can potentially affect the brachytherapy dose distribution, including emission of radiation via radioactive decay and transport, scattering, and absorption of the emitted radiation by the source itself or surrounding media. For brachytherapy, the goal of most dosimetric techniques is estimation of absorbed dose at geometric points in condensed medium within which the radioactive sealed source or sources are immersed. Useful dosimetry methods must be quantitative: the estimated dose distribution must more or less accurately approximate the actual values of well-defined physical quantities and must be accompanied by a rigorous estimate of uncertainty. The estimated uncertainty specifies an interval about the measured or computed value within which the true absorbed dose may confidently be expected to lie (Taylor and Kuyatt 1994). Widely accepted brachytherapy dosimetry techniques support uncertainties ranging from 3% to 10% over the range of clinically relevant source-to-detector distances (0.5 mm to 10 cm). Dosimetry techniques with indeterminate uncertainties or uncertainties in excess of 10% are generally not acceptable for determination of reference quality dosimetry data (Rivard et al., 2004). Another feature of most quantitative dosimetric methods is that their results are traceable to the applicable NIST standards. For all techniques, this means that absolute dose-rate estimates must be normalized to the appropriate NIST SK standard for the source model under consideration. For experimental techniques using secondary dosimeters, their dose calibrations should be secondarily traceable to the appropriate absorbed dose or kerma
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primary standards. Other measured quantities, e.g., pressure, temperature, charge, mass, time, etc., should also have a clearly defined chain of traceability to the corresponding NIST standards.
Brief History of Brachytherapy Dosimetry 226
Ra Dosimetry: From Threshold Erythema Dose to Semiempirical Models
Experimental techniques for quantifying radiation fields arising from sealed radioactive sources have been used since the discovery of 226Ra in 1898 by Marie Curie. Roentgen is credited with introduction of the gold-leaf electroscope in 1900, which is essentially a primitive ion chamber that functions both as a detector and readout device. Villiard is believed to have introduced the first ionization chamber with electrodes held at constant potential soon thereafter (Villiard 1908). A wide variety of biological and chemical dosimeters were evaluated for use in mapping “dose” distributions about early brachytherapy sources, including bleaching butter and tissue necrosis in rabbits (Quimby 1928) (see also Quimby’s fascinating 1940 review (Quimby 1941)). By plotting the distance to effect (e.g., butter bleaching) as a function of exposure time and source strength, “dose” fall-off as a function of distance could be inferred. Perhaps the most widely used biological endpoint was threshold erythema dose (TED), which is defined as the dose needed to produce a barely detectable skin erythema in 80% of the subjects so irradiated. This technique, developed by Quimby and colleagues at Memorial Hospital, apparently had a reproducibility of about 10% (Dean 1923; Quimby 1928, 1941). The concept of absorbed dose, although first defined by Christen in 1914 (Christen 1914), was itself controversial during the first half of the 20th century. One problem was that no theory existed from which the absorbed dose could be inferred from response of a detector. The development of the free-air chamber (FAC) in the 1920s, from which the quantity exposure could be rigorously inferred, led to the widespread acceptance of exposure in free-space as the primary quantity for quantitatively describing all radiotherapy treatments [see Wyckoff’s historical review (Wyckoff 1980)]. The FAC and the quantity exposure provided a sound foundation for quantifying radiotherapy treatments of the 1930s, which were limited to x-ray spectra of 300 kVp. However, for the much higher energy (about 1.2 MeV on average) gamma rays emitted by 226Ra, exposure could not be measured accurately because of the difficulty in establishing secondary charged particle equilibrium in free-air FACs of manageable size. In her 1941 review of radium dosimetry, Quimby (1941) notes that experimental estimates of the 226Ra exposure-rate constant in the 1920–1940 period varied by a factor of two. For nearly 40 years, the radiological physics community struggled to solve the problem of how to specify brachytherapy treatments in terms of rigorously defined quantities, the measurements of which are traceable to well-defined primary standards. The problem was solved in the late 1930s following the introduction of quantum mechanical treatments of photon scattering (Gray 1929) and electron energy loss (Bethe 1930; Bloch 1933). This made possible the development of a practical cavity theory (Gray 1936; Laurence 1938), which allowed exposure to be rigorously inferred from small ion chambers with condensed matter walls thick enough to establish transient charged-particle equilibrium. By 1938, experimental estimates of (Γ δ)Ra,0.5 had converged to within a few percent of Attix’s definitive 1957 measurement (Attix and Ritz 1957), which yielded 8.25 R⋅cm2⋅mg−1⋅h−1 for radium needles filtered by 0.5 mm of Pt. NIST introduced exposure-based primary standards for 137Cs and 192Ir brachytherapy sources in 1974 and 1980 (Loftus 1970, 1980). To compute dose distributions about implanted 226Ra needles and tubes, physicists employed semiempirical computational dosimetry models, reviewed in more detail elsewhere in this monograph. Quimby (1922) developed a numerical technique for partitioning needles into a linear array of point sources, from which away-and-along exposure tables could be inferred. This was followed by Sievert’s more elegant analytical solution, the Sievert integral method, in 1923. By the late 1930s, the exposure/mg⋅h tables used
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by the Quimby (1952) and Manchester (Parker 1938) interstitial implant systems were based upon computed single-source dose distributions. This approach continued without major modification upon the introduction of artificial radium-substitute radionuclides. Direct measurement of absorbed dose near brachytherapy sources, using thermoluminescent dosimeters (TLDs), radiographic film, solid-state detectors, or small ion chambers was relatively rare even in the research laboratory and had little impact on clinical practice. Since there was no basis for assigning an uncertainty to these calculated dose-rate estimates, this approach does not satisfy our criteria, described above, for quantitative dosimetry. One may speculate that the reluctance to accept measured doses as the basis of dose prescription is due both to the technical difficulties that attend such measurements as well as the empirical character of clinical brachytherapy itself.
Rise of Modern Experimental Dosimetry Methods The classical, semi-empirical approach to brachytherapy dosimetry began to break down with the introduction of 125I interstitial seeds in the early 1970s, as this 30 keV x-ray emitter clearly fell outside the scope of validated analytic models. While 125I dose distributions derived from semiempirical models were published (Krishnaswamy 1978) and widely used, it was recognized (Ling et al., 1983) that internal seed structure could modulate the emitted photon spectrum and have significantly alter the absorbed dose distribution. The introduction of an exposure-based free air chamber (FAC) primary standard in 1985 (Loftus 1984), along with its increasing clinical use, led to a growing realization that more quantitative dosimetry methods were needed. For a more detailed discussion of early of 125I dosimetry, the reader is referred to appendix C of the first TG-43 report (Nath et al., 1995). In 1986, the National Cancer Institute funded a three-year, multi-institutional contract to perform a definitive review of low-energy seed dosimetry. The three institutions, collectively called the Interstitial Collaborative Working Group (ICWG) consisted of Memorial Sloan-Kettering, Yale University, and University of California at San Francisco led by principal investigators Lowell Anderson, Ravinder Nath, and Keith Weaver, respectively (Anderson et al., 1990). Using TLD-100 thermoluminescent chips and powder capsules, embedded in machined solid-water phantoms, the ICWG developed procedures, including TLD dose calibration and energy-response correction, for making quantitative estimates of absolute dose rates in water. Each of the three ICWG investigator groups independently measured (Chiu-Tsao et al., 190; Nath, Meigooni, and Meli 1990; Weaver et al., 1989) transverse-axis dose distributions for the 125 I and 192Ir than available to validate their TLD measurement methodology. This was followed by more complete 2-D dose distributions about 125I, 192Ir, and 103Pd brachytherapy sources then available ChiuTsao et al., 1991; Meigooni, Sabnis, and Nath 1990; Nath et al. 1993). The results showed good agreement among the different measurements and overall, substantial differences between measured and classically computed dose rates for 125I seeds, but good agreement between the classical and experimental approaches for 192Ir. As a result of the ICWG efforts along with many later investigators, TLD dosimetry came to be accepted as the most reliable and best validated experimental approach in brachytherapy and its results are widely used as the basis for clinical dose computation.
Rise of Modern Computational Dosimetry Approaches Independently of the ICWG initiative, other investigators were studying the use of Monte Carlo photontransport techniques as tools for quantitative evaluation of single-source dose distributions. Based on an accurate and detailed mathematical model of the internal structure of the source, photon histories can be generated and then evaluated to assess absorbed dose. 1-D Monte Carlo simulations and other solutions of the transport equation have been used since the 1960s to calculate radial dose distributions arising from isotropic point sources in medium, e.g., the widely
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used point-source buildup factors of Berger (1968) and the tissue-attenuation and scatter-buildup factors of Meisberger (Meisberger, Keller, and Shalek 1968). 1-D Monte Carlo techniques were first applied to 125 I by Dale (1983). However, Monte Carlo techniques were applied to more geometrically complex problems in brachytherapy only relatively recently. One of the earliest 3-D studies (Williamson, Morin, and Khan 1983) used Monte Carlo simulation to assess the accuracy of the Sievert model for platinum-encapsulated 226Ra and 192Ir sources suspended in free space. Burns and Raeside (1988) published one of the first Monte Carlo dosimetry studies based upon a geometrically realistic 125I seed model. Williamson’s 1988 Monte Carlo study (Williamson 1988) was the first to note that semiempirical dose calculation models overestimated absolute dose rates about 125I by 10% to 14% and demonstrated that low-energy titanium characteristic x-rays detected by the NIST primary standard accounted for most of this discrepancy. Comparisons of ICWG TLD measurements and Monte Carlo-based dose distributions (Anderson et al., 1990; Williamson 1991) showed excellent agreement. Williamson and his colleagues published a series of studies comparing Monte Carlo dose-rate predictions to silicon diode (Perera et al., 1994; Williamson, Perera, and Li 1993) and TLD (Das et al., 1996, 1997; Kirov et al., 1995; Valicenti et al., 1995) experimental benchmarks which established the reliability and accuracy of Monte Carlo-based dose predictions across the energy spectrum of brachytherapy in both heterogeneous and homogeneous geometries. Currently, Monte Carlo simulation is a widely used and accepted dosimetry tool.
Role of the AAPM in Standardizing Low-Energy Brachytherapy Dosimetry The AAPM has had a substantial impact on the clinical practice of brachytherapy dose calculation, promoting uniformity of dose-calculation practice that has benefited literally hundreds of thousands of patients. The AAPM’s role in brachytherapy dose computation began with the publication of the TG-43 report (Nath et al., 1995). Its proposed a dose-calculation formalism, derived from the ICWG recommendations (Anderson 1990) was based upon the concept of using measured or Monte Carlo-derived dose distributions, rather than de novo dose calculations using a semiempirical model. In addition, TG43 reviewed the published TLD and Monte Carlo data available for the models 6711 and 6702 125I sources, the model 200 103Pd source, and the stainless steel-clad 192Ir seed. For each of these sources, a consensus dataset was recommended, including dose-rate constants, anisotropy constants and factors, anisotropy functions, and radial dose functions. The 1995 TG-43 recommendations implied that 125I dose rates were 10% to 15% lower than those predicted by then current dose-calculation practices. In a related development, NIST had developed a new primary standard for low-energy seed air-kerma strength based upon the wide-angle free-air chamber (WAFAC) (Seltzer et al., 2003) which suppressed measurement of 4.5 keV Ti characteristic x-rays, which was responsible for much of the discrepancy between TG-43 and semiempirical dose calculations. Because of the potential for error at the user level was high, as both source strength and absorbed dose scales were changing by more than 10%, the AAPM created an Ad Hoc Working Group under the leadership of Dale Kubo to assist the community in adapting to these changes. The Working Group report (Kubo et al., 1998) developed a step-by-step procedure for implementing the new WAFAC standard and the TG-43 formalism and recommended that the 125I monotherapy prescribed dose be revised from 160 Gy to 145 Gy. In the authors’ opinion, the complexity of these changes motivated the community and planning software vendors to rapidly adopt the TG-43 dose-calculation formalism. Clinical brachytherapy practice patterns continued to rapidly evolve, motivating the AAPM to create a permanent Radiation Therapy Committee working group (later subcommittee) on Low-Energy Interstitial Brachytherapy Dosimetry (renamed “Photon-Emitting Brachytherapy Dosimetry Subcommittee,” or PEBD, in 2003), under the leadership of Jeffrey Williamson. The practice-pattern changes included a rapid shift from radical prostatectomy to permanent seed implantation as the dominant modality for treatment of low risk prostate cancer, a proliferation of new low-energy interstitial source products (from
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2 in 1996 to over 20 in 2000), extension of the WAFAC standard to 103Pd sources in 1999, and concerns over the stability of 103Pd dose specification over time. This subcommittee took an active role the introduction of new source strength standards and revised dosimetry practices, coordinating efforts by NIST, source vendors, the clinical community, and Accredited Dosimetry Calibration Laboratory (ADCL) secondary dosimetry laboratories often on a source model-by-model basis. PEBD devised a set of dosimetry guidelines (Williamson et al., 1998), recommending that every routinely used low-energy interstitial source product have a NIST-traceable SK calibration and be accompanied by peer-reviewed published and independent experimental and Monte Carlo dose-rate distributions. These guidelines have become de facto industry standards, accepted by nearly all vendors involved in the market. PEBD also took leadership in implementing the NIST WAFAC standard for the model 200 103Pd source in a coordinated fashion with the vendor including a historical assessment of “administered-to-prescribed” dose ratios that enabled users to duplicate dose delivery practices of the past in the face of multiple revisions to the 103Pd dosimetry parameters and calibration changes (Williamson et al., 2000, 2005). Another PEBD contribution was a major revision of the TG-43 dose-calculation formalism, the preparation of consensus dosimetry parameters for eight source models, and guidelines for the practice of TLD and Monte Carlo dosimetry (Rivard et al., 2004). A supplement to this report, including consensus dosimetry parameters for an additional ~10 source models, is expected to be submitted for AAPM approval in late 2005, with subsequent publication in Medical Physics in 2006. In 2005, the AAPM reorganized the Science Council committee structure, replacing PEBD with the Brachytherapy Subcommittee (BTSC), under the leadership of Zuofeng Li, and with responsibility for guidance in all areas of brachytherapy. Working groups under BTSC supervision extend the system of dosimetry recommendations to all classes of brachytherapy sources including intravascular brachytherapy sources.
Experimental Brachytherapy Dosimetry Methods Experimental measurement is the epistemic foundation of brachytherapy dosimetry as all other quantitative approaches must be directly or indirectly validated against measurement. Dose measurement requires accurately positioned detector that yields a measurable signal with a known relationship to relative or absolute absorbed dose in the medium in the absence of the detector. Detectors commonly used in this application include small TLDs or silicon diode. Emerging detector technologies include radiochromic film, plastic scintillator, and polymer gels. Until about 1980, direct measurement of dose around brachytherapy sources and applicators in support of clinical treatment planning and quality assurance was relatively uncommon even within the research setting, let alone the clinical environment. Even today, clinical quality assurance practice is limited to experimental confirmation of the strength and geometry of purchased sources and applicators (Nath et al., 1997). Historically, this is due not only to the difficulties and labor-intensity of such measurements, but to a consensus view that dose measurement was so difficult and intrinsically inaccurate at small distances from sealed sources that even simplistic theoretical models were more reliable. Indeed, brachytherapy dose measurement does place severe demands on detectors since the dose distributions are characterized by large dose gradients, a large range of dose rates, and relatively low photon energies. A suitable detector must have a wide dynamic range, flat energy response, small size, and high sensitivity. Nearly all detector systems explored to date in brachytherapy are secondary dosimeters, i.e., detectors that require calibration against radiation fields, usually low megavoltage photon beams, in which the dose rate is known at the detector location. The one exception to this generalization is the parallel plate extrapolation chamber, which has been developed into a primary standard of absorbed dose for sealed beta-emitting sources (Soares, Halpern, and Wang 1998).
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Reference-Quality Dose Measurement System Characteristics While there are no hard and fast criteria for what constitutes a suitable dose-measurement system for brachytherapy, systems that provide credible reference-quality dose-rate measurements generally share the following characteristics: (1) Support absolute as well as relative measurement of dose at a geometric point in the 0.5-cm to 10-cm distance range in liquid water or Solid Water-substitute media (2) Support spatial density of measurements of at least 5 mm or 5 degrees in polar angle (3) Support dose measurement uncertainties of no greater than 10%, preferably no more than 5%, in the 0.5-cm to 10-cm distance range. Total uncertainty (Taylor 1994) is assumed to include all known systematic error sources (type B) as well as statistical uncertainties (type A). (4) Support independent dose estimation, i.e., should require minimal knowledge of the dose distribution to correct for detector measurement artifacts. No detector system evaluated to date satisfies all these requirements. All currently used detectors, including organic detectors such as radiochromic film, plastic scintillator, polymer gels, and solid state detectors, such as diodes and TLDs, are subject to artifacts such as volume-averaging, self-attenuation, anisotropy, and variable energy-response. For secondary systems, criterion (1) implies that a dosimeter must have sufficient temporal reproducibility and stability of response that detector responses measured in a brachytherapy irradiation geometry can be compared to those irradiated by a calibrated megavoltage beam. For single element detectors such as ion chamber or diode detectors, intercomparison requires temporal stability of response. Multiple-element detectors such as TLD or film must have sufficiently uniform responses from detector-to-detector that the responses of detectors irradiated by different fields or detectors irradiated in to different doses in the same field can be compared. Criterion (2) addresses the issue of efficiency and labor intensiveness of dose-measurement technologies. The dose-measurement process must not be so labor intensive as to preclude repeated measurement needed to establish measurement precision and dose mapping over a sufficiently dense grid of measurement points, allowing high gradient dose distributions to be meaningfully characterized. For single-source reference-quality dose measurement, the 2004 TG-43 report sets forth general recommendations (Rivard et al., 2004). Approximately 50 to 100 measurement points are needed to characterize the 2-D dose distribution of a typical low energy interstitial seed. Criteria (3) and (4) are the most difficult to satisfy. The minimum uncertainty achievable by a # D" (r)/SK measurement is about 3% (Dempsey et al., 2000), the combined uncertainties of the air-kerma strength calibration (and transfer to the measured sources) and transfer of a megavoltage beam calibration to the detector system. Assuming our target uncertainty is 8%, this leaves a residual uncertainty budget of 7.4% for all other uncertainties associated with the measurement. Criterion (4) implies that detector artifact corrections should not be so large as that their estimations requires detailed knowledge of the dose distribution. For example, volume-averaging corrections over a 1 mm3 TLD detector can be estimated from simple inverse-square law calculations, while using a 0.03 cm3 air-filled volume may require extensive Monte Carlo calculations to derive displacement and gradient corrections. To achieve the target uncertainty, the uncertainty of these corrections, or error added by uncorrected or partially corrected artifacts, must be substantially less than 5%. The following paragraphs review the major types of measurement artifacts and uncertainty sources.
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Positional Accuracy The combination of short source-detector distances and inverse square-law gradients means that brachytherapy dose measurement is exquisitely sensitive to accuracy with which the detector can be positioned at known locations relative to the source. For example, to limit positioning uncertainties to 2% at distances of 2 mm, 5 mm, and 10 mm from a point source, requires positioning accuracies of 20, 50, and 100 µm, respectively. For typical single-element detector mechanical scanning systems, symmetry of orthogonal detector-reading profiles can be used to estimate the center of radioactivity in the detector coordinate system (Ling et al., 1985) as illustrated in Figure 1. This method results in a measurement coordinate system with the origin placed at the center of radioactivity (not physical source center) but does not account for location of the center of the active source volume in the detector housing. The accuracy achievable depends on the extent of hysteresis, deviation from orthogonality, and positioning precision of the detector transport mechanism. With micrometer-quality drives, 50 to 100 µm accuracies can be achieved. Alternatively, for a micrometer-positioned diode detector in a water phantom, Williamson has described a mechanical and optical alignment procedure which achieved an accuracy of 0.03 mm (Williamson, Perera, and Li 1993). For single-element integrating detectors such as TLD chips, solid water-substitute phantoms equipped with machined cavities and recesses for the source and array of TLD detectors is widely used (Anderson et al., 1990; Nath, Meigooni, and Meli 1990). The geometric accuracy achievable is determined by the machining accuracy and cavity size enlargement required for ready removal and insertion of source and detectors (Figure 2). It is estimated that geometric accuracies of 0.1 to 0.2 mm can be achieved by the solid phantom approach (Meigooni et al., 1995; Rivard et al., 2004). 2-D detectors [radiochromic film (Dempsey et al., 2000)] and 3-D detectors [polymer gels (Maryanski et al., 1996)] provide increased opportunity for improved geometric accuracy. Such detectors consist of a continuous sheet or volume of sensitive material. The detector coordinate system is imposed by the digital readout system [optical CT (Gore et al., 1996) or MRI for polymer gels and scanning densitometers (Dempsey et al., 1999) for radiochromic film], which assigns a signal intensity to each voxel or pixel of the detector medium. Digital readout systems often support very high spatial resolutions and geometric precisions (about one-third of the pixel dimension) assuming they are free of geometric distortion. Usually a combination of the orthogonal profile method and mechanical alignment of the detector with respect to the source is used to obtain localization accuracy of the order of 25 to 100 µm. Signal-to-Noise Ratio, Dosimeter Readout Precision, and Signal Stability Achieving a good balance between readout precision and spatial resolution is a challenge faced by all dosimeter systems in brachytherapy. Brachytherapy dose rates range from 50 cGy/s [5 mm from a high dose-rate (HDR) source] to much less than 1 cGy/h (5 cm from a clinical-strength 125I seed). Both small size, to better resolve spatially varying dose distributions and minimize volume-related detector artifacts, and high sensitivity are desired. The inherent sensitivity of a detector system is governed by its quantum efficiency (energy transferred from the radiation field per observed signal quantum) and the collection efficiency. Table 1 summarizes data for several commonly used single-element detectors. Because TLD fluorescent response is a second-order effect, its inherent response is much inferior to other detectors while silicon diode is the most efficient. Despite its poor intrinsic quantum yield, the practical response of TLD per unit volume approaches that of ion chamber, because TLD is solid. By irradiating TLDs long enough, integrated signals with good signal-to-noise ratio characteristics have been obtained over the distance range of 1 to 7.5 cm near low dose-rate (LDR) sources. Another consideration is electronic noise and background, which can dominate readout statistics at low signal intensities. For example, the signal from the ion chamber in Table 1 at a dose rate 0.1 cGy/h is only 1×10−17 amperes, which is significantly less than the leakage current from the best commercially available electrometer. In contrast
15–Quantitative Dosimetry Methods for Brachytherapy
(a)
(b)
(c)
(d)
241
Figure 1. Illustration of source symmetry method for identifying the coordinates, (X0,Y0) denoting the projection of the brachytherapy active source center in the plane (a) of a 2D detector. Panels (a) and (b) illustrate radiochromic film detector readings in an arbitrary detector frame of reference fixed to the upper left hand corner of the film. (X0,Y0) are located by finding the detector coordinates that bisect the distance intervals corresponding to the full-width half maximum of two orthogonal profiles, A and B, near the source center.
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Figure 2. Solid-water phantoms designed by A. S. Meigooni at the Mallinckrodt Institute of Radiology for measurement of transverse-axis (left panel) and polar (right panel) dose-rate distributions. At distances less than 2 cm, small (1 × 1 × 1 mm3) TLD-chips are used. Unused detector slots are filled with solid-water plugs.
Table 1. Absolute and Relative Sensitivity of Dosimeters Commonly Used in Brachytherapy. The relative signal and photon yield estimates assume 1 mm3 detector volume and absorbed dose of 0.1 cGy, respectively. Data from Perera et al. (1992). Detector
TLD-100 Silicon Diode Ion Chamber (air) Plastic Scintillator fiber
Energy (eV) dissipated/ observed quantum
No. quanta emitted
Typical quantum efficiency x geometric collection efficiency
Practical signal relative to ion chamber
8400
1.8 × 106
0.20 × 0.08
0.120
3.6
4.0 × 10
9
1.00 × 1.00
17,000
33.8
2.4 × 10
5
1.00 × 1.00
1.00
100
6.2 × 10
7
0.20 × 0.20
10.3
to the ion chamber, TLD is a practical dosimeter since (a) the TLD detector itself integrates the signal rather than an external readout system, (b) TLD background signal does not increase appreciably with time, and (c) long 6- to 48-hour irradiations are practical as many TLD detectors can be simultaneously irradiated in a single experiment. Similarly, model MD-55-2 radiochromic film, despite requiring nearly 80 Gy to achieve an optical density of 1.0, is practical for LDR brachytherapy dosimetry, because the signal maybe integrated for days (Monroe et al., 2001) or even weeks (Chiu-Tsao et al., 1994) to achieve good signal statistics. Small nonintegrating small diode and plastic scintillation detectors (Fluhs 1996) are useful because they have much higher effective sensitivities than ion chamber. For the purposes of reference-quality dose-rate measurements, the AAPM recommends using detectors that have 1s statistical uncertainties (type A) ≤ 5%. Stability of detector response through time is another problem that must be considered. For example, TLD detectors are subject to fading artifacts if not properly annealed (see below). Radiochromic film response continues to increase after exposure the rapidity and extent of which depends on a complex
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way on temperature (Reinstein and Gluckman 1999a) and dose (Ali et al., 2003, 2005) by careful temporal synchronization of the calibration and brachytherapy exposures (Dempsey et al., 2000). Energy-Response and Other Artifacts Another desirable feature of a brachytherapy dosimeter is a constant response as a function of energy. When this condition is present, sensitivity (net detector reading/dose in medium) is constant at all measurement points in the brachytherapy geometry. Because most dosimeters used in brachytherapy are composed of crystalline or organic media, their effective atomic numbers usually do not exactly match that of water or tissue. Point-to-point variations in detector sensitivity due to corresponding changes in the photon spectrum are known as energy-response artifacts. Energy-response artifacts are illustrated by Figure 3, which shows detector reading, relative to dose in water, as a function of energy for source-detector distances of 1 and 10 cm for several common detectors. Even for organic detectors (Zeff = 5.9) the measured reading/unit dose depends significantly on photon energy. In the low-energy photon range (20 to 30 keV), detector sensitivities tend to be independent of measurement distance, even though detector-response corrections relative to the megavoltage photon calibration maybe as large as 7 for silicon diodes. Thus, any of these detectors maybe used to map 103Pd and 125I dose distributions without detailed knowledge of the photon spectrum at each measurement point. However, for higher-energy sources, e.g., 192Ir at 400 keV, Figure 3 shows that diode response varies by 60%, a finding that has been confirmed experimentally (Williamson et al., 1993). Even when using TLD for 192Ir measurements (Meli, Megooni, and Nath 1988) or diode for 137Cs measurements (Williamson, Perera, and Li 1993), artifacts of approximately 10% to 15% have been observed. Hence for dosimetry of intermediate and high-energy brachytherapy sources, detectors should be matched to the composition of the surrounding medium as closely as possible.
Figure 3. (Left) Theoretical detector response (dose to detector/dose to water) of pure plastic scintillator (PVT), arsenic-doped plastic scintillator and TLD-100 (LiF) detectors at 1 cm and 10 cm distances in water from point sources of various energies. (Right) Same except TLD-100 detector is replaced by a silicon diode detector. The results of each detector type are normalized to unity at 1 cm from the 137Cs source. (Reprinted from Int J Radiat Oncol Biol Phys, vol 23, “Rapid two-dimensional dose measurement in brachytherapy using plastic scintillator sheet: linearity, signal-to-noise ratio, and energy response characteristics.” H. Perera, J. F. Williamson, S. P. Monthofer, W. R. Binns, J. Klarmann, G.L. Fuller, and J. W. Wong, pp. 1059–1069. © 1992, with permission from Elsevier.)
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General Formalism for Absorbed Dose Measurement Consider the general problem of inferring the absorbed dose rate per unit air-kerma strength (SK), # # BTx in the absence of the detector at location r from the detector reading, R(r ) the active [ Dwat (r# ) / SK ]meas center of which is located at the same point in space. As illustrated in Figure 4, the goal is to estimate dose rate to water at a geometric point in the reference liquid-water phantom (30-cm liquid water sphere) recommended by AAPM for specification of single-source dose distributions (Rivard et al., 2004). Clearly, many corrections are needed, including those for averaging of dose over the detector volume, displacement of the medium by the detector, and differences in the composition, size, and shape of the # BTx measurement phantom compared to the reference water sphere. [ D" wat (r ) / SK ]meas is given (Williamson and Meigooni 1995) by: #
#
R(r ) ⋅ g(T ) D" wat (r ) S = S ⋅ F ( R(r# )) ⋅ ε ⋅ E (r# ) K meas K lin λ where # R( r ) = g(T) =
ελ =
# E (r ) =
BTx
(2)
is the net detector reading integrated over the measurement time, T, in hours.
λ 1 − e− λT is the effective irradiation time for an #exposure time T near a source with a decay constant λ (h−1). Applying this correction to R(r ) converts the integrated reading into an hourly rate decayed to the beginning of the exposure.
the detector sensitivity, ε λ = R0 ( D0 )med meas where R0 is net the response of the detector to an arbitrary reference dose (D 0)med from the reference megavoltage photon beam λ used to calibrate the detectors. For detectors that have a linear response to absorbed dose, ελ is independent of the reference dose level. λ
[ R (r# ) / D (r# )] for Brachytherapy source = ε ( R , r# ) ε (R ) [ R / D ] for Calibration Beam λ 0
wat
0
med
BTx
0
λ
0
(3)
This quantity is the energy-response correction, defined as the linearized response of the detector per unit dose in water medium of the brachytherapy source relative to that of the reference beam λ. # # λ BTx # Dwat (r ) and Dmed are the doses necessary to achieve the reference reading R0 and E (r ) = Dmed Dwat (r ) at reading level R0.
Flin ( R ) =
[ R / D ] at a dose level correspoonding to reading R ε λ ( R) = [ R0 / D0 ] at a reference reading R0 λ ε λ ( R0 )
(4)
is the detector linearity correction, which is usually measured using the calibration radiation source. SK =
is the air-kerma strength of the source in units of U where 1 U = 1 mGy·m2·h–1 = 1 cGy·cm2·h–1 decayed to the beginning of the experiment.
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Figure 4. Illustration of the dose-measurement problem in brachytherapy. Given the reading of a volume detector located at (x,y,z) in a potentially non-water, nonspherical phantom, the goal is to estimate the dose rate at per unit SK at a geometric point when the source is placed at the center of liquid-water phantom 30 cm in diameter.
(
)
# # Flin R(r) ⋅ ε λ ⋅ E(r) in the denominator of equation (2) is equal to The product of the three terms # # its sensitivity ε B Tx (r) in the brachytherapy irradiation, i.e., the detector reading R, at position r in the brachytherapy measurement phantom per unit dose to water in the reference liquid water phantom. Of the above corrections, the relative energy response correction is the most difficult to specify accurately. Indeed, for low-energy source dosimetry, its uncertainty dominates the total dose-measurement # uncertainty (Rivard et al., 2004). Corrections for many phenomena are bundled into E(r), including detector artifacts such volume averaging, self attenuation, displacement of surrounding medium as well # as intrinsic energy response. In addition, E(r) accounts for the nonliquid water equivalence of the measurement phantom and while its denominator includes corrections to the megavoltage beam calibration, e.g., use of a solid-water–substitute phantom and any other deviations from the TG-51 protocol (Almond et al., 2005). To avoid systematic errors, the calibration conditions assumed by the denominator of equation (3) must match those of the detector calibration protocol used to measure # ελ. Both direct experimental and theoretical methods have been used to evaluate E(r) for low-energy source TLD dosimetry. If one assumes that the detector response is proportional to dose absorbed by the detector active volume, i.e., Rdet = α ⋅ D det λ
λ
(5)
where α is a constant independent of the photon energy, λ, of the source, then one can show that (Das et al., 1996) # # ∆D det(r)/∆D w at(r) # E(r)= ∆D det /∆D m ed MC
λ
MC
BRx
(6)
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where the delta, i.e., ∆D, refers to dose relative to the inherent normalization of the Monte Carlo code (usually some quantity proportional to Gy/simulated photon history). The terms in the numerator are derived from two parallel Monte Carlo simulations, center of a 30 cm liquid # one with the source at the # ∆D ( r ) water sphere yielding an estimate of dose, at the geometric point r , and a simulation includwat # ing the source and detector geometry in the measurement phantom, yielding an estimate of ∆Dwat (r ), # energy imparted to the detector active volume when centered at r , per unit mass of active material. The denominator requires calculation of the same quantities for the megavoltage photon beam calibration. This theoretical approach rigorously models such effects as volume averaging (displacement correction), displacement of the medium by the detector (replacement correction), and that component of energy response variations due to composition and density mismatches between the detector and the target water medium. Using photon Monte Carlo simulation, Williamson has evaluated relative energy response for TLD and other types of detectors using a modified form of equation (6): # # β ⋅ ∆K det (r ) / ∆K wat (r )]BTx [ # [ E (r )]MCPT = [ β ⋅ ∆K det / ∆K med ]λ
(7)
where absorbed dose has been replaced by collision kerma. The constants β are small corrections of the order of 1% that account for contribution to detector response by secondary electrons arising in the surrounding medium, which can be significant when the minimum linear dimension of the detector is comparable to the range of the secondary electrons. Usually, β can be evaluated with sufficient accuracy using Burlin cavity theory (Attix 1986; Valicenti et al., 1995). # The theoretical method of evaluating E (r ) is valid only if the detector response is inherently linear with respect to energy deposition, a characteristic often referred to as “LET linearity.” As discussed below, such LET linearity cannot be assumed for brachytherapy dosimeters, especially for low photon energies. The direct empirical approach consists of irradiating the detector to a known dose in free space in an x-ray beam (see Figure 5) with a spectrum that matches the brachytherapy source of interest. The criterion of equivalence is usually agreement between the effective energy derived from the beam half-value layer and the kerma-weighted mean energy of the brachytherapy source spectrum. Reference beams must be calibrated with an ion chamber having a NIST-traceable air-kerma calibration factor. During the development of quantitative TLD dosimetry, a number of investigators (Luxton et al., 1990; Meigooni, Meli, and Nath 1988a; Muench et al., 1991; Weaver 1984) used variations of this approach obtaining distance# independent E (r ) values ranging from 1.39 to 1.44 for TLD-100 detectors at 125I energies.
(
FS Measurements illustrated in Figure 5 provide an estimate of the free-air response, R K air
K
FS air
(R
)
hν
, where # is the air-kerma in free space at the detector center and hv is the beam energy. To estimate E (r ) from FS K air
)
hν Meas
Meas
for brachytherapy in-phantom measurements, a number of additional corrections are
required:
(
)
hν
(
)
hν
FS FS K FS D wa R K air air t thy # Meas ⋅ E (r) = λ # BTx ( R Dmed )Meas crepl ⋅ cdisp ( r ) thy
(8)
The quantities bearing the subscript “thy” must be evaluated by Monte Carlo or other theoretical method. The ratio
(
K FS D FS air wat
)
hν
thy
≈ ( µ en (hν ) / ρ )wat air
converts the air-kerma calibration to dose in water.
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Figure 5. Direct experimental measurement of energy-response correction by intercomparing the detector to a calibrated ion chamber in a beam the spectrum of which matches that of the brachytherapy source of interest.
# The replacement and displacement corrections, crepl and cdisp r, account for replacement of absorbing medium by the detector and averaging of dose over the detector volume. They depend on the brachytherapy dose-measurement geometry and are defined as: crepl =
D w at in m edium FS
K w at in em pty cavity occupied by detector
# # 1 # cdisp (r)= D w at (r) atpointr # ∫ D wat (r' )dV ' V (r)V (#r)
(9)
For 30 mg throw-away TLD powder-filled capsules, Weaver (Weaver et al., 1989) estimated crepl values of 0.965 to 0.970 for 125I dosimetry. By averaging 1/r2 over the TLD volume (Meigooni, Meli, and Nath 1988a; Weaver et al., 1989), yield cdisp values of 1.03 and 1.01 for 0.5 cm and 1 cm distances, respectively, for 3×3×0.9 mm3 TLDs and cdisp corrections of 1.06 and 1.02 for the larger TLD capsules at these distances. Using Monte Carlo simulation, Patel and Williamson (Patel et al., 2001) found that # the ratio of the product crepl ⋅ cdisp (r) to that of a perfect point detector to be about 1.05 for 1×1×1 mm3 TLD cubes at distances of 1 cm or greater for 125I seeds in Solid Water. Hence, for low-energy seed dosimetry, these corrections can be significant. The common practice of using published E values that may or may not match the experiment in question should be avoided and is inconsistent with the current TG-43 dosimetry recommendations (Rivard et al., 2004). The 2004 AAPM TG-43 protocol (Rivard et al., 2004) makes many detailed recommendations regarding the experimental determination of TG-43 dose-calculation parameters. While the recommendations are based largely on experience with TLD dosimetry of low-energy interstitial seeds, most are applicable to other detector systems. For measurement of dose-rate constant, AAPM recommends that dose
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measurements be obtained for 8 to 10 different sources, for a total of 15 individual readings at 1 cm on the transverse axis. At least one of the sources should have a directly traceable SK calibration (i.e., have been measured directly by NIST via the WAFAC) and the other experimental sources calibrated against this source by the investigator. For measurement of relative quantities (radial dose and 2-D anisotropy functions), at least three separate experimental runs are recommended using different sources with a total of about 12 readings, including readings in different quadrants assumed to be equivalent by cylindrical or planar symmetry. For most investigational sources, e.g., 241Am and 169Yb, as well the widely used highintensity 192Ir sources, no primary air-kerma strength standards are available. In these cases, it is the investigator’s responsibility to carefully measure SK for these sources using a large-volume ionization chamber that has direct traceability to NIST’s primary air-kerma standards for external beam therapy that bracket the spectrum of the experimental source. A number of investigators (Das et al., 1995; Goetsch et al., 1992) have described procedures for carrying out these measurements. With care, measurements with 2% to 3% uncertainty even for LDR sources are possible.
Brachytherapy Dose-Measurement Systems In the following sections, a number of dose-measurement techniques are briefly reviewed, with an emphasis on their potential for supporting measurement of reference-quality single-source dose-rate measurements. Thermoluminescent Dosimetry (TLD) Techniques in Brachytherapy Currently, TLD dosimetry is the most widely accepted experimental “gold standard” for reference-quality dose measurement in brachytherapy. Dr. John Cameron and his group (Cameron, Suntharalingam, and Kenney 1968) introduced many of the instruments, annealing and handling techniques, and radiation oncology applications that have made TLD a clinically practical dosimetry tool. In 1966, this group published the first TLD dosimetry study in brachytherapy (Ponnunni Kartha, Kenney, and Cameron 1966). As discussed earlier, acceptance of TLD-based brachytherapy dosimetry in the 1990s was due to the efforts of the NIH-supported ICWG (Anderson et al., 1990), the excellent agreement between TLD and Monte Carlo dose-estimation techniques (Williamson 1991), and the emphasis placed by AAPM on quantitative dosimetry in clinical practice standards (Nath et al., 1995; Rivard et al., 2004; Williamson et al., 1998). TLD Handling and Readout Procedures. Although instrumentation and TLD annealing practices vary from investigator to investigator, TLD-100 LiF extruded ribbons subjected to the classical pre-irradiation annealing protocol (annealing TLD detectors at 400 ° C for 1 hour followed by rapid cooling and 24 hours of 80 ° C annealing prior to irradiation) (Zimmerman, Rhyner, and Cameron 1966) is the most common approach. The low-temperature annealing eliminates the short-lived, low-temperature glow-curve peaks so that TLD readings do not depend on the time interval between exposure and reading. Weaver (1984) validated pre-annealed TLD-100 powder. Meigooni has published the most comprehensive study of the influence of annealing and readout procedures on the precision and linearity of TLD-100 detectors (Meigooni et al., 1995). Using a Harshaw 2000 reader, he demonstrated that nitrogen gas purging of the TLD readout chamber dramatically improved the reproducibility of repeated readings in the low dose range and minimized the differences in precision and linearity of TLD response between acute and protracted LDR exposures. In addition, Meigooni showed that the classical pre-irradiation annealing protocol was equivalent to a more time-efficient protocol consisting of 1 hour of 400° annealing before irradiation followed by 10 minutes of 100° annealing just prior to reading (Meigooni et al., 1995). With nitrogen flow and a conventional TLD reader, minimum exposures of 0.5 and 1 cGy are required to maintain precisions of ±5% and ±2%, respectively, for 3 × 3 × 0.9 mm3 detectors. For small (1 × 1 × 1 mm3) TLD chips, the corresponding low-dose limits at are 1 cGy and 15 cGy. Finally,
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Meigooni’s work shows that significant deviation from linearity can occur (Figure 6) for low dose exposures, as well as at high doses, especially for small 1 mm TLD cubes. Thus, linearity of TLD response should always be verified across the entire dose-measurement range and any significant corrections included in dose-rate estimates. Most investigators find it necessary to correct for variation of individual detector sensitivities (Cameron, Suntharalingam, and Kenney 1968; Weller et al., 1983) in order to achieve 2% to 4% standard deviations with 4 to 6 readings. Use of detector-specific sensitivity corrections, Si, requires that individual TLD chips be stored in an aluminum or glass-annealing planchette in individually numbered receptacles. In moving TLDs to and from the irradiation phantom to the planchette, care must be taken not to scramble the detectors with respect to their identifying numbers. The Si factors are measured by exposing the entire batch to an identical dose of megavoltage x-rays. The relative detector sensitivity, Si, is defined as Si = TLi
1 N
N
⋅ ∑ TLi
(10)
i= 1
where TLi denotes the reading of the i-th chip. Several sequential uniform irradiations and readings are required to accurately define a table of Si factors. Background readings should be subtracted from each TLi if significant. Some individual chips may never stabilize and should be discarded. With reference to # equation (2), the net mean reading at position r becomes n # # R(r)= 1/n∑ i= 1 (TLi(r)− TLbkgd )/Si
(11)
Linearity of the TLD detectors should never be assumed (Meigooni et al., 1995), especially when measuring doses on the order of a few cGy. Near the low-dose limit, the measured TL may not vary linearly with dose because of surface effects arising from nonradiation induced thermoluminescence or uncertainty in dark-current (zero-dose) TL measurements. Linearity corrections should be carefully
Figure 6. TLD-100 responses (left axis) and linearity corrections (right axis) for graded 4 MV x-ray doses for large (Left panel) and small (right panel) detectors. Nitrogen gas purging was used along with the standard 24 hour 80∞ pre-irradiation annealing protocol. (Reprinted from Med Phys, vol 22, “Instrumentation and dosimeter-size artifacts in quantitative thermoluminescence dosimetry of low-dose fields.” A. S. Meigooni, A. S., V. Mishra, H. Panth, and J. F. Williamson, pp. 555–561. © 1995, with permission from AAPM.)
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measured by irradiating small groups (n = 4–6 chips) of detectors to graded doses of megavoltage xrays over the brachytherapy dose range (generally 1 to 1000 cGy). A plastic phantom (polystyrene or solid-water), which lends itself to accurate determination of absorbed in medium, should be used. During the reading session, the background or “dark current” reading, TLbkgd, should be carefully measured and subtracted from every reading. The first author’s approach is to take TLbkgd as the reading with no detector positioned in the reader. At each dose level k, an average reading
TL
k
= 1/ n
∑ (TL
i
− TLbkgd ) / Si
can be calculated where i is the i-th detector in k-th dose group which is exposed to a dose Dk to the phantom material. From this data, the response ε k = TL k / Dk can be calculated and plotted as a function of TL. The linearity correction, Flin(TL), can be defined as Flin (TLnet ) =
[TL
net
/ D ] at a dose D corresponding too TLnet
[TL
0
/ D0 ] at a reference reading TL0
=
ε (TLnet ) ε (TL0 )
(12)
where TL0 is the net reading that results from an exposure in the linear region of detector response (D0 = 40–100 cGy) in which the correction is normalized to unity. The example in Figure 6 shows that this correction can approach 10% in the brachytherapy dose range. Relative Energy Response of TLD. According to the 2004 TG-43 protocol (Rivard et al., 2004), the relative energy response correction, E(r), is the largest single source of type B (systematic) uncertainty for TLD dosimetry in the low-energy range. For 125I and 103Pd dosimetry, most investigators treat E(r) as a distance-independent constant with a value near 1.4, although when volume-averaging and solidto-liquid water corrections are included, E(r) varies significantly with distance (Patel et al., 2001) (Figure 7, lower left). However, for higher-energy sources, in which the scatter-to-primary ratio rapidly increases with distance, variations in E(r) with distance as large as 10% to 15% have been reported (Meigooni, Meli, and Nath 1988b) (Figure 7, upper right). The standard value of 1.4 represents an average of reported TLD-100 responses to low-energy x-ray beams measured in air, illustrated in Figure 7 (upper left). Both Monte Carlo (Patel et al., 2001) and experimental (Meigooni, Meli, and Nath 1988a) data demonstrate that E(r) values inferred from in-air measurements depend on TLD size. Since freeair measurements relate TLD reading to dose in a void left by removing the chip, a replacement correction (2%-5%), is needed to correct for the phantom material displaced by the detector. Accurate measurement of E(r) is difficult because: (i) photons from the low-energy tail of the Bremsstrahlung spectrum bias the measurements to an unknown extent, (ii) the limited precision of TLD readout, and (iii) the relatively large uncertainty of ion-chamber dosimetry in this energy range. Because of the uncertainties associated with E(r) measurements, some authors have proposed calculating E(r) directly by Monte Carlo simulation (Patel et al., 2001; Vlicenti et al., 1995; Williamson and Meigooni 1995). Although volume-averaging, displacement, and detector self-attenuation corrections can be easily included, the method assumes that TLD response is proportional to energy imparted to the detector (intrinsic linearity), an assumption which has been questioned for some TLD phosphors and annealing and glow-curve analysis techniques (da Rosa and Natte 1988; Tochilin, Goldstein, and Lyman 1968). For the widely used TLD-100 chips, used TLD-100 chips, and Cameron annealing and readout techniques, the evidence for intrinsic linearity is controversial. Das et al. (1986) compared the Monte Carlo and the experimental free-air x-ray beam approaches for four beam qualities ranging down to 19 KeV effective energy. Their measured relative responses (1.42 to 1.50) were in good agreement (relative to stated 4% experimental precision) with measurements reported by other investigators and their own Monte Carlo calculations (Figure 7, lower left). However, a more recent and comprehensive paper by Davis et al. (2003) concludes the opposite, that the measured TLD-100 energy response correction is underestimated by Monte Carlo calculations by 10% to 5% in the 24 to 47 keV energy range. Their
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Figure 7. Several plots illustrating properties of the relative energy-response correction. Upper Left: Plot of early (1980-1992) energy-response data (Hartman, Schlegel, and Scharfenberg 1983; Meigooni, Meli, and Nath 1988a; Muench et al., 1991; Weaver 1984) for TLD-100 detectors using calibrated x-ray beams as function of effective energy. Upper Right: Relative energy-response as a function of transverse-axis distance for 192Ir and 169 Yb sources calculated by the first author’s Monte Carlo code. Lower Left: TLD-100 detector relative energy response corrections calculated by the code PTRAN_CCG as a function of transverse-axis distance from a Symmetra 125I source for two compositions of Solid Water™ and two detector sizes (Patel et al., 2001). Lower Right: Comparison of measured to calculated energy-response corrections as a function of effective x-ray beam energy for two relatively recent studies (Das et al., 1996; Davis et al., 2003).
E(r) values, measured for nine beam qualities, have a stated uncertainty of 0.6%. Their data, presented as normalized TL/Kair ratios, suggest that relative energy response corrections for 103Pd–125I spectra are in the range of 1.58 to 1.61, about 10% higher than previously reported values. Clearly, further research is needed to resolve the discrepancy between published E(r) values, to identify the appropriate role for transport calculations in TLD dosimetry, and to reduce the large uncertainty associated with relative energy-response corrections (Rivard et al., 2004). Using a 10% larger energy response correction would
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cause measure dose-rate constants to fall by 10%. This would improve the agreement between Monte Carlo and TLD for some sources (Hedtjarn, Carlsson, and Williamson 2000) but would clearly worsen agreement for others, e.g., 103Pd (Monroe and Williamson 2002). In utilizing measured or Monte Carlo E(r) estimates published by others, the recommendations of the 2004 TG-43 report should be heeded. TLD experimentalists should confirm that the measurement methodology assumed by a published E(r) ratio matches their experimental technique with regard to TLD detector type and size, annealing and readout technique, and megavoltage beam calibration technique. The latter requires accounting for differences in calibration phantom material and dose-specification media used by the experimentalist and assumed by the selected E(r) estimate. The experimentalist should confirm that appropriate volume averaging, displacement, and self-absorption corrections regardless of whether they are included in E(r) or applied separately. Choice of Phantom Material for Brachytherapy Dose Measurements. As noted earlier, one of the most significant sources of measurement error is uncertainty in the relative position of source and detector. Since TLD measurement commonly is based upon sparse sampling of the dose distribution using a relatively small number of independently positioned detectors, the symmetry method (Figure 1) cannot be used to identify the source center. The ICWG investigators controlled positioning artifacts by using plastic phantoms with precisely-machined detector cavities whose location relative to the source can be accurately specified by using digitally-controlled milling machines (see Figure 2). The phantom material selected by the ICWG, which remains widely used to this day, was a commercial water substitute, Solid Water™, manufactured by Gammex-RMI (Middleton, WI). It is a mixture of epoxy resin and calcium carbonate, the proportions of which are designed to reproduce megavoltage beam depth doses in water and the appearance of liquid water on CT imaging. Because early studies (Meigooni, Meli, and Nath 1988a) were interpreted as demonstrating that Solid Water and liquid water were equivalent for low-energy 125 I dose distributions, it was not immediately appreciated that this level of calcium additive overcompensated in the 30 keV range. Using Monte Carlo simulation, Williamson (1991) first showed that use of Solid Water measurement medium underestimated the dose rates in liquid water by factors ranging from 4% at 1 cm to 25% at 10 cm. The AAPM now recommends that all low-energy brachytherapy measurements be corrected for nonwater equivalence of the measurement medium (Rivard et al., 2004). Recently, the measured calcium concentration of Solid Water was found to deviate from the vendor’s specification by as much as 30% (1.6% Ca by weight compared to the vendor’s specification of 2.3%) (Patel et al., 2001), the consequences of which are illustrated by Figure 7 (lower left). Another measurement of Solid Water composition by the same author found that the calcium concentration was lower than the specification by 10% (Chiu-Tsao et al., 2003) which raises concerns over the constancy of its composition. Therefore, when Solid Water or other proprietary water substitutes are used in experimental dosimetry, the AAPM recommends (Rivard et al., 2004) that their atomic composition be measured and these results used via Monte Carlo simulation to derive distance-dependent phantom-to-liquid water corrections. The present authors recommend using high-purity commercial plastics such as polystyrene or polymethylmethacrylate, which have more uniform, reproducible, and better-characterized compositions, for future low-energy, photon-emitting brachytherapy dosimetry studies. However, the plastic-to-water conversion coefficients, which are much larger (Meigooni, Meli, and Nath 1988a) than corresponding Solid Water corrections, need to be carefully evaluated. For 192Ir and higher-energy source dosimetry, absorbed dose is independent of phantom composition (Meli, Meigooni, nd Nath 1988), allowing use of commercial plastics such as PMMA without significant corrections. Summary: Uncertainty of TLD and Role in Clinical Dosimetry. Relatively few of the brachytherapy experimental and computational investigations published prior to 1999 included a rigorous uncertainty analysis. Based upon the approach developed in several more recent studies (Gearheart et al., 2000; Monroe et al., 2001; Monroe and Williamson 2002; Nath and Yue 2000), the AAPM (Rivard et al., 2001) now recommends that all publications claiming to provide reference-quality brachytherapy
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dosimetry data include an uncertainty analysis adhering to the approach outlined in NIST Technical Note 1297 (Taylor and Kuyatt 1994). This report recommends using the Law of Propagation of Uncertainty (LPU) to estimate the uncertainty of a quantity y, that has a functional dependence on measured or estimated quantities x1,…xN, as follows: y = f ( x1 ,..., x N ) N −1 N ∂f 2 ∂f ∂f σ = ∑ σ 2 σ x ,x + ∑ ∑ x i = 1 j = i + 1 ∂xi ∂x j i = 1 ∂xi 2
N
2 y
i
i
(13) j
where σxi,xj (assumed zero here) represents the covariance of the two variables. By defining the relative uncertainty propagation factor as %
∂Y
∂x
≡
x ∂Y
Y ∂x
, equation (13) becomes
∂f %σ y = ∑ % %σ x ∂xi i =1 N
2
(
i
)
2
(14)
For the detector-to-dose conversion equation (2), in which the quantities xi are multiplied or divided #
only, equation (14) can be further simplified by noting that % ∂ [ D" wat (r ) / SK ]meas ∂xi = 1 BTx
. The cont-
ributing uncertainties, %σxi are classified by the NIST report as either type A or type B. The former are statistical uncertainties describing the reproducibility of repeated measurements. NIST recommends taking %σxi to be the %standard error about the mean (%standard deviation divided by square root of number of independent observations) for type A uncertainties. Type B uncertainties are defined as “other than type A,” including estimated systematic uncertainties. If a quantity xi is believed with confidence to lie in the interval [a1,a2], a common approach is to assume that xi is uniformly distributed over this
(
interval yielding σ x = ( a2 − a1 ) / 2 3 i
)
(Taylor and Kuyatt 1994). By following these recom-
mendations, the final %σy will approximate the half-width of the 68% confidence interval expressed as a percentage of the mean. Table 2 represents the authors’ estimate of TLD and Monte Carlo (discussed elsewhere in this chapter). The 4% estimate of TLD readout precision is that proposed by Gearheart (Gearheart et al., 2000) and Nath (Nath and Yue 2000), while the 5% uncertainty in E(d) is recommended by AAPM (Rivard et al., 2004). The medium (specific to RMI Solid Water) conversion uncertainty estimate is based upon Figure 7 (lower left) and the assumption that calcium content by weight is a uniform distribution140 bounded by [1.6%, 2.3%]. The final result, that TLD measurement uncertainty varies from 8% to 10%, depending on distance, agrees well with the uncertainty analyses of Gearheart, Nath, and the AAPM ((Gearheart et al., 2000; Nath and Yue 2000; Rivard et al., 2004). In summary, TLD dosimetry, using TLD-100 LiF chips or extruded ribbons, has emerged as the detector offering the best compromise between small size, sensitivity, flat energy response, and ease of accurate positioning: it is currently accepted as the experimental “gold standard” for measurement of absolute dose rates in brachytherapy. Its disadvantages are: labor intensiveness, relatively poor spatial resolution because of practical limits on the number of measurement points, and relatively large total uncertainty dominated by limited readout precision, uncertainties in available energy-response estimates, and what has turned out to be an unfortunate choice of measurement medium. While Monte Carlo simulation has,
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Jeffrey F. Williamson and Mark J. Rivard Table 2: Uncertainties for 125I Transverse-Axis TLD and Monte Carlo Dose Estimation · → TLD uncertainties in measurement of Dwat (r )/SK for 125I in Solid Water
Component
1 cm distance
5 cm distance
%σxi
Type
%σxi
Type
Repetitive TLD measurements
4%
A
4%
A
TLD calibration (including Linac calibration)
3%
A+B
3%
A+B
Solid-to-liquid water conversion
2%
B
6%
B
Seed and TLD positioning errors (∆d = 100 µm)
2%
B
1%
B
Energy-response correction
5%
B
5%
B
ADCL SK measurement + transfer
2%
B
2%
B
Total combined uncertainty
7.9% 9.5% · → 125 Uncertainties for Monte Carlo estimates Dwat (r )/SK for I in liquid water
Statistics
0.3%
1.0%
Photo ionization cross-sections (∆σPE = 2.3%)
1.5%
4.5%
Seed geometry
2.0%
2.0%
Source energy spectrum
0.1%
0.3%
Total combined uncertainty
2.5%
5.0%
in principle, a much smaller uncertainty, its results depend on the accuracy of the input data. Monte Carlo can anticipate such phenomena as fail of the vendor or NIST calibration procedures to adhere to the definition of air-kerma strength, the presence of contaminant radionuclides on the source, or limited capacity to reconstruct the detailed internal structure of radioactive seeds. TLD measurement, on the other hand, is sensitive to all these effects, provided their dosimetric consequences are large enough to be detected as significant differences. For the reasons discussed above, the AAPM continues to recommend that all marketed low-energy brachytherapy products be subjected to both experimental (TLD or in some cases diode) characterization and Monte Carlo dosimetric characterization (Rivard et al. 2004;Williamson et al., 1998). In practice, Monte Carlo estimates of relative dosimetry parameters (radial dose function and anisotropy functions), verified by measurements within experimental uncertainty, are recommended for clinical use (Rivard et al. 2004) while an average of experimental and Monte Carlo dose-rate constants is recommended. The AAPM report (Rivard et al. 2004) assigns an average uncertainty of 4.8% to the averaged, or “consensus” dose-rate constant. Using its recommended consensus datasets, the AAPM report estimates that the total uncertainties of TG-43 dose calculations at 0.1 cm, 1 cm, and 5 cm are 6.7%, 5.7%, and 7.3%, respectively. This includes estimated uncertainties of the dose-rate constant and radial dose functions and an additional 3% uncertainty associated with transfer of the NIST-traceable SK values by the vendor’s calibration procedure. Diode Dosimetry Techniques in Brachytherapy Diode dosimetry was one of the first quantitative measurement techniques applied to low-energy seed dosimetry. One of the earliest applications by Ling et al. (1983) was measurement of the radial dose function for the newly introduced model 6711 125I seed in 1983. Anisotropy functions derived from diode
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255
measurements followed soon thereafter (Ling et al. 1985; Schell et al., 1987). These data were widely used in clinical practice until the late 1990s, when the community accepted the TG-43 approach. Comparisons of diode measurements to TLD measurements and Monte Carlo calculations (Ling et al., 1985; Williamson 1988, 1991; Williamson and Quintero 1988) revealed good agreement. The 2004 TG-43 report (Rivard et al., 2004) endorses use of diode measurements for the experimental determination of relative (radial dose and anisotropy functions) but not for absolute dose measurements (dose-rate constant). Diode dosimetry has also been applied to higher energy brachytherapy, most notably in the mapping of dose distributions about 137Cs-bearing shielded gynecological applicators (Mohan et al., 1985; Yorke et al., 1987). The most comprehensive set of diode measurements is that of Williamson et al. (Perera et al., 1994; Williamson, Perera, and Li 1993) who used diode measurements to benchmark Monte Carlo calculations for a wide range of brachytherapy photon spectra and geometries. As noted earlier, silicon diodes are among the most sensitive dosimeters used in brachytherapy, providing good signal-to-noise ratios at long distances from low strength brachytherapy sources. Active detector volumes are often very small, providing excellent spatial resolution. Finally, diode dosimeters are designed for use in liquid water medium, avoiding large corrections for non-water measurement media. Diode detectors are most often used with scanning water phantoms designed for external beam dosimetry. The P-type silicon diode, designed by Rikner (Rikner 1985; Rikner and Grusell 1987) for electron-beam measurements and marketed by Scanditronix as the “electron field diode,” or EFD, is the most frequently used detector. P-type diodes such as the EFD must be used cautiously. Because of their low impedance, most integrating electrometers used for ion chamber dosimetry are not appropriate. Most investigators have used the dedicated Scanditronix diode electrometer and readout system. Williamson (Williamson, Perera, and Li 1993) used a Keithley 602 electrometer in the fast current mode to readout EFD signals, obtaining a final reading by timer-driven integration of the strip chart output. Diode readings are very sensitive to small potential differences across the electrometer input. Williamson found that potential differences greater than 0.05 µV were associated with large sensitivity to temperature changes, detector nonlinearity, and unpredictable background currents. By using a constancy-check source, e.g., 90Sr ion chamber check source, a session-to-session reproducibility of 1% can be achieved. The mechanical scanning system must be very carefully calibrated so that the diode position relative to the source center can be estimated accurately. The source symmetry method (Figure 1) is widely used for this purpose. Williamson and colleagues (Williamson, Perera, and Li 1993) have described a mechanical and optical alignment technique that assures a positioning accuracy of 0.02 to 0.03 mm. Diode detectors do respond linearly the instantaneous doses characteristic of linac x-ray beams. Hence a 60Co teletherapy unit is preferable as a calibration source. While most investigators account for the distance-dependence of diode response due to spectral changes for high-energy sources, such as 137Cs, the effect is usually assumed to be negligible for 125I. Li (Li, Williamson, and Perera 1993) evaluated this hypothesis by means of Monte Carlo simulation, which included detailed geometric modeling of EFD internal structure and estimating energy deposition in its 60 µm thick active volume. The ratio of Ddet to silicon kerma for a point detector was found to be independent of distance from a model 6702 source within 1.5%, indicating that variation in the energy or anguMC
lar distribution of photons in water did not significantly affect results. In addition, the Rm eas D det
ratio was found to be constant within 3% over the 1 to 10 cm distance range, although the results show a definite trend towards lower responses at larger distances. In the same study described earlier in our discussion of TLD energy response, Das investigated the dependence of EFD diode response on photon energy (Das et al., 1996). For x-ray and γ-ray sources ranging from 19 keV to 60Co, these investigators found that the fundamental energy response, α = R/Ddet [see equation (5)], was constant within a 3% range.
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Jeffrey F. Williamson and Mark J. Rivard
The most comprehensive study of diodes in brachytherapy is a series of papers published by the first author’s group (Das et al., 1996, 1997a; Kirov et al., 1995; Li, Williamson, and Perera 1993; Perera, Williamson, and Li 1994; Williamson, Li, and Perera 1993). The major goal of these works was not to characterize brachytherapy dose distributions, but to systematically assess the accuracy of Monte Carlo photon-transport simulation as a clinical dosimetry tool through systematic experimental benchmarking. As illustrated by Figure 8 (lower left and right panels), for relative measurements (Perera, Williamson, and Li 1994; Williamson, Perera, and Li 1993), excellent agreement was found. The average discrepancy between Monte Carlo simulation and EFD measurements ranged from 0.8% to 3% for 125 169 I, Yb, 192Ir, and 137Cs sources both in homogeneous water and in the presence of cylindrical air, titanium, aluminum, steel, and lead heterogeneities of various diameters and thicknesses. Figure 8 (lower left panel), addresses the accuracy of absolute EFD measurements in brachytherapy. For four different radionuclides, the inherent energy response, αBTx(d) is plotted as a function of transverse-axis distance, d, in water:
α BTx (d ) =
( R( d )
( ∆D
det
SK )meas BTx
(d ) ∆SK )MC
BTx
(15)
For different sources and distances, αBTx(d) is an energy-independent constant within ±2.5%. The average of the αBTx(d) agrees with Das’ measurements (Das et al., 1996) of intrinsic energy response within 3%. Thus, available data suggests that absolute dose-rate diode measurements with uncertainties of about 5% are possible, using Monte Carlo estimates of the relative energy response correction. In summary, diode dosimetry has a well-validated role in relative dose measurements about lowenergy brachytherapy sources. Although not widely used for absolute dose-rate measurements in this energy range, further investigation is warranted, as diode dosimetry may have significantly lower measurement uncertainties than those reported for TLD dosimetry. Because of the relatively large variation in detector sensitivity with measurement position, silicon diode is probably not a good choice for higher energy sources. Plastic Scintillator Dosimetry for Brachytherapy Plastic scintillator (PS) consists of a solid polystyrene or polyvinyl toluene base, along with one or more wave-shifting dyes. It converts about 3% of the energy absorbed from ionizing radiation to visible light through a complex multi-step process. PS detectors are approximately tissue equivalent and are much sensitive per unit detector volume than either TLD or ion chamber (see Table 1). In brachytherapy, singleelement PS probes has been intensively investigated by the Essen University group (Bambynek et al. 2000; Fluhs et al., 1996). These probes consist of a small PS detector mounted on the tips of a nonscintillating PMMA-clad light-conducting fiber which transports scintillation photons to a photomultiplier tube. Their system for measuring β-emitting eye plaque or intravascular source dose distributions consists of a 0.4 mm thick by 1 mm diameter BC-400 PS detector (Bambynek et al., 2000). To subtract optical background, which is dominated by Cerenkov radiation above the electron energy threshold of 170 keV, a second coaxial optical fiber that is optically isolated from the detector is readout by a second photomultiplier tube. Excellent linearity and reproducibility of 1.5% to 2% has been reported by several authors (Mourtada et al., 2003; Soares et al., 2001) using this system. Soares et al. (2001) and Mourtada et al, (2003) have calibrated single-element detectors against NIST-traceable sealed β-source standards and have compared absolute dose rates measured by PS to Monte Carlo calculations and other dose-rate measurements about a variety of therapeutic sealed β sources. Their intercomparison assigned an uncertainty of 10% to the PS system, and found that PS results agreed reasonably well with the other
257
Transverse axis dose distance2 (cGy·cm2·h–1)
15–Quantitative Dosimetry Methods for Brachytherapy
Figure 8. Comparison of brachytherapy diode measurements in water to Monte Carlo simulation for radionuclides ranging from 125I to 137Cs. Top: Relative transverse-axis dose rates as measured (data points) and as calculated by Monte Carlo simulation (lines) (Perera et al., 1994; Williamson, Perera, and Li 1993). Note that doses averaged over the active detector volume are plotted. Lower Left: Ratios of measured diode readings in water to dose in the active detector volume calculated by Monte Carlo simulation as a function of transverse axis distance for four different radionuclides (Das et al. 1996). The broken line indicating constant sensitivity is the average absolute detector response measured by Das using calibrated x-ray beams. (Reprinted from Phys Med Biol, vol 41, “Accuracy of Monte Carlo photon transport simulation in characterizing brachytherapy dosimeter energy response artifacts.” R. K. Das, Z. Li, H. Perera, and J. F. Williamson, pp. 995–1006. © 1996, with permission from IOP Publishing, Bristol.) Lower Right: Measured and calculated heterogeneity correction factors (HCF = dose in presence of heterogeneity/dose at same point in homogeneous geometry) for a 169Yb source perturbed by lead disks of various shapes and thicknesses. Replotted from data in Perera et al. (1994).
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Jeffrey F. Williamson and Mark J. Rivard
measurements. In interpreting such intercomparisons, one must bear in mind that primary dose-rate standards for sealed β-sources utilizing extrapolation chambers themselves have rather large uncertainties of 6% to 10% (Soares, Halpern, and Wang 1998). The first author’s group has proposed using thin PS sheets as 2-D planar dosimeters (Perera et al., 1992) and organic scintillation cocktails to support 3-D dosimetry (Kirov et al., 2000). While the brachytherapy source irradiates the organic scintillator, low-noise CCD optical cameras are used to integrate the 2-D optical yield distribution (repeated at multiple viewing angles for 3-D dosimetry), from which the corresponding dose distribution is inferred. While this work is highly investigational and many problems remain to be solved, 2-D and 3-D organic scintillation offer the potential of on-line, efficient, high-resolution measurement of brachytherapy dose distributions. An important issue for photon-emitting brachytherapy PS dosimetry, is the linear energy transfer (LET) linearity, i.e., constancy of α = R/Ddet over the range of secondary electron energies encountered in brachytherapy. Reduction of PS sensitivity to high-LET radiations has been widely observed and is thought to be caused by ionization quenching (Birks 1970), in which alternative pathways compete with radiative de-excitation (fluorescence) are created by densely ionizing charged particle tracks. Whether ionization quenching or other forms of LET nonlinearity affect PS response to electrons under 100 keV is controversial [see Williamson et al. (1999) for a recent review]. Using methods similar to the previously discussed TLD and diode energy-response investigation (Das, Perera, and Williamson 1996), Williamson et al. (1999), measured the intrinsic response, α = R/Ddet of BC-400 and two experimental PS mixtures to photon sources ranging from 19 keV to 1.25 keV. After correcting for the radiological properties of PS, they found that PS response to 20 and 100 keV photons was 30% and 10% smaller, respectively, than the response to than the response to 192Ir photons. Thus, PS detectors should be used cautiously for measurement of absolute dose-rates about low-energy sources. In summary, plastic scintillator holds promise for brachytherapy dosimetry, both in its technologically well-developed single-element detector format and in the more investigational multidimensional format. However, more quantitative investigation of the accuracy and uncertainty of single-detector systems for photon-source dosimetry is needed before it can be accepted for reference-quality dose-rate or relative dose measurement. Radiochromic Film Dosimetry for Brachytherapy Radiochromic film (RCF) is emerging as a promising secondary dosimeter for 2-D brachytherapy dosimetry because of its high precision, excellent spatial resolution and approximate tissue equivalence. The RCF active sensor consists of diacetylene monomer crystals suspended in a water based gelatin, which in turn is sandwiched between protective Mylar™ sheets (Niroomand-Rad et al. 1998). Ionizing radiation initiates polymerization of the monomer, resulting in complex changes in the optical absorption spectrum, which are approximately proportional to absorbed dose. The signal is readout by means of a scanning densitometer, most commonly using a He-Ne monochromatic (633 nm) light source or redfiltered broadband light source with a mean wavelength of 680 nm. More recently, document scanners (Alva et al., 2002; Mack et al., 2003) has been explored for quantifying film response. The most widely investigated commercial film type, MD-55-2, has a thickness of 240 µm and two 12-µm thick active layers (Klassen, van der Zwan, and Cygler 1997). It is a rather insensitive detector, requiring a dose of nearly 80 Gy to produce a net optical density (OD) change of 1.0 at 633 nm. Recently, a nearly ten-fold more sensitive RCF version has been developed (Lynch et al., 2004), which reduces the useful dose range to under 4 Gy. Its application to quantitative brachytherapy dosimetry remains to be investigated. Most investigations of RCF in photon-emitting brachytherapy were based upon the double active layer RCF model MD-55-2. One of the earliest investigations (Chiu-Tsao et al., 2004) used a predecessor of MD-55-2 to map dose distributions from a 57 µ Gy⋅m2⋅h−1 model 6702 source. An exposure time of two weeks was used. Good qualitative agreement between relative dose rates measured by RCF and TLD
15–Quantitative Dosimetry Methods for Brachytherapy
259
and calculated by Monte Carlo simulation was observed. NIST (Soares 1991) has developed a procedure for using radiochromic film for mapping the relative dose distributions near sealed beta-emitting therapeutic sources. Piessens et al. (2002) found that RCF absolute dose-rate measurements agreed with plastic scintillator probe and TLD measurements within 3.5% at distances of 1 to 3 mm. This study accounted for RCF nonuniformity using the double exposure technique. In a comparison of measured absolute dose rates about an 90Y intravascular. RCF dosimetry exhibits a number of artifacts that must be meticulously controlled if accurate results are to be obtained. Zhu and colleagues at Washington University (Zhu et al., 1997) found that RCF strips exhibited 8% to 15% systematic variations in response for two different scanning wavelengths apparently due to non-uniform application of the sensitive emulsion on the film. These authors devised the so-called “double-exposure” non-uniformity correction technique. Prior to exposing a film to the unknown radiation source, the film is first exposed to a uniform dose of 20 Gy and scanned by a laser densitometer to obtain a pixel-by-pixel relative sensitivity map. The 2-D OD image obtained via scanning following the second exposure is then registered to the flat-field image. The flat-field image is subtracted from the second exposure image and the difference image corrected pixel-by-pixel by means of the relative sensitivity map. Zhu demonstrated that double-exposure densitometry eliminates structural non-uniformity, producing a constant response characterized by 2% to 5% random fluctuations. Despite the insensitivity of MD-55-2 film, the response above 3 Gy was relatively precise, with repeated readings exhibiting a standard deviation of 2% at a spatial resolution of 0.25 mm. Dempsey et al. (1999) characterized another source of artifact arising from the interaction of the film with the widely used 633 nm scanning laser densitometry systems, specifically the Molecular Dynamics Personal Densitometer. Nonreproducible interference fringes, due interference of light reflected from layer boundaries within the film and the glass scanning bed, as large as 7% were observed. These were managed by replacing the scanning bed by a defusing glass plate. Even more disturbing were severe light scattering/temporal system response artifacts. As illustrated by Figure 9, large OD gradients and discontinuities cause OD in dark regions to be underestimated by as much 40% and OD in high-transmission regions adjacent to darker regions to be substantially overestimated. For inverse square-law OD variations characteristic of brachytherapy, Dempsey showed that the system response artifact introduces dose
Figure 9. (a) Unprocessed Molecular Dynamics scans through uniform OD rectangular steps of varying widths, demonstrating OD underestimates as large as 30%. (b) The same scans after filtering and deconvolution (Dempsey et al., 1999).
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underestimates of about 10%. These artifacts were successfully eliminated by a combination of Wiener filtering and deconvolution in light-transmission space, based upon the measured point-spread function of the system. Since potentially any high-speed densitometry system and film (including silver halide) type combination could give rise to such artifacts, bar pattern scanning as recommended by Dempsey should be a routine quality assurance test for any quantitative application of film dosimetry. Once ionizing radiation exposure initiates polymerization, this chemical transformation persists upon removal of the radiation source, resulting in progressive increase of the OD as a function of post-exposure time and temperature (McLaughlin et al., 1991). The so-called “post-exposure OD growth effect” can persist for many months. The OD growth rate is initially quite fast (as large as 20% in 12 hours for 30 Gy exposures) varies systematically with post-exposure temperature history (Reinstein and Gluckman 1999b). In addition, both the growth rate and magnitude of growth vary systematically with dose (from 100% for small doses to 20% for large doses) (Ali et al., 2003; Dempsey et al., 2000). Published (Ali et al., 2005) and unpublished (Le and Williamson) data demonstrates that fractionated and protracted low dose-rate exposures, especially at low doses, result in moderately elevated (2% to 5%) OD responses, relative to single-fraction exposures, that take several hundred hours to dissipate. Thus the potential for significant dose-measurement errors is large if the acute calibration exposures, needed to measure the dose-response curve, and the brachytherapy exposures are not temporally synchronized. Despite these technical problems, the potential of radiochromic film to significantly reduce the uncertainty of brachytherapy dose measurements is high. After correcting for RCF nonuniformity and densitometer artifacts, as described above, the reproducibility (95% confidence interval) of uniformly irradiated MD-55-2 doubly-exposed film was 2% for a 0.4 × 0.4 mm2 pixel at 2 Gy and 1.5% % for a 0.8 × 0.8 mm2 pixel at 1 Gy (Dempsey et al., 2000). At doses greater than 10 Gy, even single 100 µm pixel readings have a precision better than 1%. Dempsey and colleagues (Dempsey et al., 2000) at Washington University benchmarked RCF absolute dose measurements against Monte Carlo calculations for an HDR 192Ir source. The HDR and 6 MV calibration exposures were temporally synchronized, minimizing post-exposure OD growth artifacts. Double-exposure densitometry (Zhu et al., 1997) and densitometer system response corrections (Dempsey et al., 1999) were employed. Between doses of 10 and 150 Gy, 95% of the measured doses (100 µm resolution) agreed with the Monte Carlo benchmark within 4%. Even at low doses of 1 to 5 Gy away from the source cable, the Monte Carlo-RCF was 1% to 2% (2σ). The authors estimated that the total uncertainty of the measured absolute dose rates was 3.6% to 4.2% (coverage factor of two) for single-pixel readings at doses above 10 Gy. This is the lowest uncertainty reported for a brachytherapy dose measurement using a secondary dosimeter. Two major problems must be solved before RCF dosimetry can be applied to determination of reference-quality dose-rate distributions for low-energy brachytherapy. Accurate relative energy-response measurements for low energy x-rays are needed to confirm their independence from LET established. In an investigation patterned after the Das et al. TLD-diode study (Das, Perera, and Williamson 1996), Bohm et al. (Bohm, Pearson, and Das 2001) found that [OD/Drcf] was a constant, within ±5%, independent of photon energy in the 30 keV to 1.25 MeV energy range. Although MCNP photon-electron simulations were used to calculate dose to the active film layer, the precision of the experiment was estimated to be about 7%. Monroe et al. (2001) compared doses measured by RCF dosimetry to Monte Carlo doses for a balloon applicator filled with organically bound 125I solution. They found excellent agreement between the two techniques for high doses, indirectly confirming the energy-response correction derived from Monte Carlo calculations. While there is no theoretical basis for questioning the linearity of RCF response with respect to energy deposition, more accurate measurements and more detailed Monte Carlo calculations are needed. Finally, as described above, the problem of post-exposure growth, which gives rise to discrepancies between acutely exposed calibration films and films subjected to protracted LDR exposures, must be solved. The solutions advanced to date include (a) modeling of postexposure growth artifacts (Ali et al., 2005), (b) annealing of RCF to accelerate convergence of
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post-exposure OD growth (Reinstein, Gluckman, and Meek 1998), and (c) selection of densitometry wavelengths that minimize growth artifacts (Meigooni and Williamson, unpublished data). As noted above, the new high-sensitivity film may address some of these issues. While some problems remain to be solved, RCF dosimetry is the best developed of the emerging dosimetry systems and has well-documented performance advantages compared to TLD. Polymer Gel Dosimetry for Brachytherapy Since the 1950s, liquid gelatin materials have been used as tissue equivalent media for experimental dosimetry (Frigerio 1962; White 1978). However, consideration of gels as 3-D radiation detectors did not occur until the 1980s (Gore, Kang, and Schulz 1984). In this first decade of use, a variety of radiotherapy applications were considered (Schulz et al., 1990; Maryanski et al., 1993). Readout techniques also developed with some groups using nuclear magnetic resonance (NMR) imaging while others pursuing optical imaging techniques (Gore et al., 1996; Oldham et al., 2003; Oldham and Kim 2004) due to concerns of widespread MRI scanner unavailability (Oldham et al., 2001). Furthermore, gel dosimeters may be divided into three classes of detectors: Fricke gels (MacDougall et al., 2002; De Deene, De Wagter, and Baldock 2003); polymer gels (Maryanski et al., 1994); and superheated drop detectors (Lamba et al., 1998; d’Errico et al., 2002). This section will review application of polymer gels to HDR and LDR brachytherapy sources using both readout techniques. Free radicals are produced during irradiation of polymer gel dosimeters, with average distances of approximately 0.2 µm and 0.02 µm at doses of 0.1 and 100 Gy, respectively. In less than 1 ns, these radicals then form cross-linked micropolymers which both increase the gel opacity and cause local changes in the polymer structure which subsequently disturb the NMR relaxation rates. Thus, optical and NMR imaging techniques are suitable to perform detector readout to glean the distribution of radiation damage. While there are a variety of polymer gel dosimeter mixtures available such as the various recipes of BANG (Maryanski et al., 1996; Maryanski 2005), VIPAR (Baras et al., 2002; Kipouros et al., 2003), PABIG (Fong et al., 2001), and MAGIC (Pantelis et al., 2004), the mechanism of damage and micropolymerization are all similar since they all strive for tissue-equivalence over the photon energy range (10 keV to 10 MeV) used in radiation therapy. Among conventional radiation detectors used in the field of radiation therapy, polymer gels have the unique capability to acquire and render 3-D radiation dose distributions. With modern polymer gels, spatial resolution is generally limited to 2 mm in all directions by the readout technique and the need to balance signal intensity with diminishing detector sensitivity. Optical readout may be performed using a variety of light sources such as He-Ne laser at 633 nm or a laser diode at 670 nm. However, maximum opacity peaks at approximately 450 nm. Optical computed tomography (optical CT) reconstruction techniques may be implemented as is commonly performed for diagnostic radiology CT (Xu, Wuu, and Marayanski 2004). A photo of an optical CT scanner is presented in Figure 10. Conventional MRI units (~1.5 T) can be used to readout the nuclear magnetic transverse relaxation rate, T2 signal, with 1 < TR < 8 s and 20 < TE < 1,600 ms. Depending on the desired spatial resolution and signal-to-noise ratio, high-quality results using MRI readout can be typically obtained in less than one hour. As most applied research in this field occurs in an academic setting, availability of MRI for detectors was not a limitation and the majority of studies performed to date used MRI instead of optical means for detector readout. The suitability of polymer gels as radiation dosimeters for brachytherapy is evidenced by considering detector response as a function of various parameters. As long as the irradiation is kept to less than one hour, polymer gel detectors typically exhibit negligible dose rate dependence over the range used in radiation therapy—including the high dose rates used in HDR 192Ir brachytherapy. Furthermore, detector response as a function of post-irradiation time to readout does not significantly as this time is kept less than 30 days (Farajollah et al., 1999; Xu, Wuu, and Marayanski 2004). With an offset to account
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Figure 10. The OCTOPUS™ optical CT scanner for polymer gel readout. (Photograph courtesy of M. Maryanski.)
for intrinsic opacity, polymer gel detector response as a function of delivered dose is quite linear. The explicit range depends on detector size and readout technique, but linearity ±3% over a range of 1 to 20 Gy may be expected. Slightly larger variations in flatness may be expected for photons with energy ranging from 10 keV to 10 MeV (Pantelis et al., 2004). Because of polymer thermodynamics, detector response is somewhat dependent on irradiation temperature. For a dose of 7 Gy and temperature change from 23 °C to 25 °C, polymer gel detector efficiency increases by approximately 7% when readout using MRI (Spevacek et al., 2001). The majority of studies performed using polymer gels as dosimeters for brachytherapy have been to compare dose distributions of HDR 192Ir sources using other experimental) or Monte Carlo techniques to benchmark the gel detector. McJury and colleagues (McJury et al., 1999) compared measured dose distributions using the BANG gel with calculated dose distributions using an HDR 192Ir brachytherapy treatment planning system (PLATO v.11.3 from Nucletron BV, Veenendaal, The Netherlands) based on Monte Carlo measurements by Williamson and Li (Williamson and Li 1995) and measurements by Mishra and colleagues (Mishra, Waterman, and Suntharalingam 1998). The HDR 192Ir source was centrally positioned in a 500 cm3 Pyrex flask (see Figure 11) to deliver 6 Gy at 1.0 cm over 111 seconds as determined using the treatment-planning system (TPS). Separate vials containing polymer gel were prepared for absolute dose calibration using a linac 6 MV photon beam. MRI readout occurred 24 hours after irradiation using TR = 2 s, TE = 0.05 s, 2 mm slice thickness, 1 × 1 mm2 pixel size, and lasted 25 minutes. Over a radial range of to 50 mm, the average difference between the measured and TPS results was 0.17 Gy with a standard deviation of 0.13 Gy. At a distance of 1.0 cm, a dose of 6.5±0.2 Gy was measured using the polymer gel detector. This was considered good agreement since a positional uncertainty of 0.5 pixel corresponded to an absorbed dose error of ±0.8 Gy. Papagiannis and colleagues (Papagiannis et al., 2001) performed a similar study using the same type of HDR 192Ir brachytherapy source, planning system, and calibration beam. Even with subtle differences (TR = 4 s, TE = 0.04 s, readout 48 hours post-irradiation, 0.43 × 0.43 mm2 pixel size, 3-mm slice thickness, etc.), similar results were obtained. At distances less than 0.5 cm, differences in Monte Carlo and poly gel doses ranged from 10% to 40%. At larger distances, substantially better agreement was observed.
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Figure 11. Polymer gel dosimeter flask for measurement of HDR 192Ir source dose distribution. (Reprinted from McJury et al., “Experimental 3D dosimetry around a high-dose-rate clinical 192Ir source using a polyamide gel (PAG) dosimeter,” Phys Med Biol, vol 44. Copyright 1999, with permission from IOP Publishing)
However, there was concern that a practical and lucid technique for performing an uncertainty analysis was not available for polymer gel dosimeters (Oldham et al., 1998; Baldock, Murray, and Kron 1999). This concern was later addressed (MacDougall et al., 2002; De Deene, De Wagter, and Baldock 2003). De Deene and colleagues (De Deene, Reynaert, and De Wagter 2001) also assessed the accuracy of using polymer gel dosimeters in close proximity to an HDR 192Ir source. The methodological details were similar to the previous two studies, but emphasis was placed on the impact of oxygen and its role on inhibiting radiation damage. They observed 2% lower dose at 1.0 cm that estimated using Monte Carlo methods (Williamson and Li 1995) with a root mean square (RMS) uncertainty of 5.6%. In retrospect, positioning of the source within 30 mm of the flask inner wall might have significantly minimized the scatter and resultant absorbed dose at distances greater than 10 mm (Perez-Calatayud, Granero, and Ballester 2004). For gels containing oxygen concentrations ranging from 0 to 0.8 mg O2 per liter of gel, a linear increase in the threshold dose was observed with a slope of 28.2 Gy⋅L⋅mg–1. Thus, the threshold dose may be kept under 0.2 Gy as long as the oxygen concentration remains less than 0.007 mg/L. Additional studies have been performed to obtain polymer gel detector response for other brachytherapy sources. Farajollah and colleagues (Farajollah et al., 1999) measured the dose rate from 36 Nucletron Selectron LDR 137Cs brachytherapy sources using TLDs and BANG gel with MRI readout. At 2 cm, the average gel dose rate was 10% larger than calculated. Dose rates measured using TLDs from two different institutions were 3% low and 2% high. Considering the source calibration uncertainty of ±5% and other uncertainties such as applicator tube attenuation, both the gel and TLD results were in good agreement with the calculated dose rate. Fragoso and colleagues (Fragoso et al., 2004) also measured dose rate in the vicinity of Selectron LDR 137Cs sources, but this time EGSnrc was used to complement gel measurements with Monte Carlo calculations to account for applicator tube attenuation. Monte Carlo voxels were 1 × 1 × 1 mm3, and 54 M photon histories were calculated. A 1.5 T MRI with TR = 2 s and 50 < TE < 800 ms was used for gel readout. At distances beyond 1 cm and doses less than
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10 Gy, agreement between Monte Carlo calculations and polymer gel measured doses were typically within 5%, which was well within the experimental uncertainties. Hasson (1998) also measured dose distributions near 192Ir and 137Cs sources, and obtained agreement within 3% for dose distributions when compared to radiochromic film and Monte Carlo methods. Farajollah performed a similar study (Farajollah 1999) and obtained similar results. Ibbott and colleagues (Ibbott et al., 1999) measured the dose rate of an LDR 125I source towards determining brachytherapy formalism parameters, but did not give quantitative information as to the agreement of results using other techniques. Hafeli and colleagues (Hafeli et al., 2000) compared polymer gel dose measurements of a 188W/188Re intravascular brachytherapy source (IVBT, Nath et al., 1999) to an EGS4-based Monte Carlo calculation and measurements using TLD rods and radiochromic film from 0.5 to 5.0 mm. Normalized to the calculated dose rate at 2.0 mm, the, gel, TLD, and film dose rates were 1.01, 0.89, and 1.20, respectively. Volume averaging and positioning uncertainties subtended the largest components of the overall experimental uncertainties. Chan and colleagues (Chan et al., 2001) measured dose distributions in the vicinity of a 15.5 mm diameter LDR 106Ru eye plaque. Calibrations used four 22 mm diameter vials and irradiation with 6 MV to doses of 0.00, 1.00, 2.00, and 3.00 Gy. MRI readout was performed using TR = 2 s, TE = 20ms or 100 ms, 3 mm slice thickness, and 0.31 mm in-plane pixel size. Though 3-D dose distribution data were obtained, only central axis dose data were presented, and results were not compared to established techniques. In summary, use of polymer gel dosimeters as relative instruments having impressive 3-D capabilities is well established for a variety of radiotherapy applications, including the diverse scope of brachytherapy. However, additional research is needed to validate polymer gel dosimeters as an absolute detector for the demanding application of brachytherapy dosimetry.
Computational Dosimetry Methods in Brachytherapy As noted in the introduction, computational dosimetry methods have come to play an important role in the estimation of reference-quality dose-rate distributions used to implement TG-43 dose calculations in clinical practice. Indeed, AAPM guidelines (Williamson et al., 1998) specify that every low-energy interstitial seed product used in routine clinical practice should be supported by both measurements and dose-rate distributions computed by means of Monte Carlo simulation. Eligible measured or computed reference-quality dose distributions must meet the standards of the peer-reviewed publication process. Most of the AAPM-endorsed consensus TG-43 datasets (Rivard et al., 2004) were based upon purely Monte Carlo estimates of radial dose and anisotropy functions. Computational dosimetry methods fall into one of two broad categories: semiempirical dose-calculation algorithms and numerical solutions of the Boltzmann transport equation (BTE). Semiempirical algorithms are reviewed in chapter 14 of this volume and are not discussed further in this chapter. These algorithms are based upon heuristic approximations that support very efficient dose calculations for treatment planning. While very useful within limited domains, the accuracy of such algorithms must be established by benchmarking them against more rigorous computational or empirical dosimetry methodologies. Rigorous numerical BTE solutions collectively form one such class of methodologies and is the topic of this section. In principle, the BTE gives a complete mathematical description of the motion of particles through a host medium when the particle density is small relative to that of the host medium. When applied to the interaction of ionizing radiation with matter, the stationary (time independent) BTE takes the following form (Duderstadt and Martin 1979; Williamson 1988): $ # # ˆ ⋅ ∇Φ(r$, Ω, E ) + µ ( r$, E ) Φ(r$, Ω, E ) Ω # # # $ # $ # = ∫∫ Φ(r , Ω ', E ') ⋅ µ ( E ', Ω ' → E , Ω | r ) ⋅ dΩ ' dE ' + S (r , Ω, E )
(16)
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$ # # Φ(r,Ω,E ) is the angular flux (particles per unit where angle and energy at location r, # area,#solid $ # µ r,E is the linear attenuation coefficient, µ(E ',Ω '→ E ,Ω |r) is the differential cross # # section E ' , Ω' t o E, Ω giving the number of particles scattering from energy and angular trajectory per unit $ # # fluence, and S(r,Ω,E ) is the source # term, giving the number of particles arising at r per unit volume with direction and energy E and Ω . Equation (16) is essentially a balance equation. Its four terms, proceeding from left to right describe passage of particles along straight-line trajectories without collision, losses of particles due to collisions, gains of particles due to inscattering, and gains due to primary particle# production. of the problem geometry (i.e., specifying cross sections # # Given a description $ µ(E ',Ω '→ E ,Ω |r) and µ r,E as a function of spatial location), the primary source term and other boundary conditions, equation (16) can be solved for angular fluence, giving a complete characterization of the radiation distribution that accounts for particle transport, scattering, and absorption. Unfortunately, equation (16) is too complex to be solved accurately by analytic methods in any but the simplest of 1- and 2-D geometries. General deterministic numerical solutions, e.g., discrete ordinates, have been developed (Lewis and Miller 1993) mainly for neutron transport applications. As deterministic BTE solutions are increasingly used for solving photon transport problems in brachytherapy and other radiotherapy applications, they will be briefly discussed in this chapter. However, the most widely used numerical methodology used in brachytherapy dosimetry is stochastic or Monte Carlo solution of the BTE. In this approach, (16) is converted to an integral equation, which is solved by summing the quantity of interest over a randomly generated set of integrand values, called particle histories. Even in the presence of a complex 3D geometry, Monte Carlo will give statistical uncertain but unbiased estimates of that converge to the “exact” BTE solution in the limit of large number of histories.
(
)
(
)
Monte Carlo-based Brachytherapy Dosimetry: General Considerations The basic principles of Monte Carlo simulation have been extensively described elsewhere (Jenkins, Nelson, and Rindi 1988; Williamson 1988) and will be only briefly reviewed here. A 3D geometric description of all sources and attenuating media in system is required. A discrete-event Monte Carlo simulation involves randomly constructing a set of photon tracks or “histories” (see Figure 12(b)), each consisting of a series of successive collisions or interactions of the photon with matter connected by straight-line flight paths. The point of origin of each history is randomly selected from the radioactive distribution (usually assumed to be uniformly distributed in some volume or surface). Its trajectory in space is randomly selected from an isotropic distribution and its energy from the photon emission spectrum of the radionuclide. By ray tracing along the trajectory to find mean free path as a function of distance, the site of the first collision can be selected from an exponential distribution. The type of collision is randomly selected from the relative frequencies of photoelectric, coherent scattering, incoherent scattering, and other competing interaction types (see Figure 12(a)). The photon direction and trajectory of the scattered photon emerging from the collision are sampled from the normalized differential cross section. This process of randomly simulating serial collisions continues until the photon history is terminated by absorption, escape from the system, or Russian roulette. This process is repeated for 100,000 to many millions of histories depending on the variance reduction techniques employed. During random construction of each collision, the history is scored, i.e., the actual or estimated energy deposition to each simulated “detector” is computed and tallied, a process called “estimation.” The simplest estimation process is the so-called analog estimator, illustrated by Figure 12(b), in which energy is deposited in a detector volume if and only if a collision occurs in that spatial volume. Both the mean dose and variance about the dose are tallied. By central limit theorem, the confidence interval, describing the statistical uncertainty of the sample mean dose, is inversely proportional to the square root of the number of histories. In principle, it is straightforward to include transport of secondary electrons in the simulation, which is clearly necessary to solve problems in the megavoltage photon domain. However, including electron
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Figure 12. (a) Principle mechanisms of photon scattering and energy deposition in the 10 keV to 10 meV energy range, illustrating the various photon scattering processes and secondary photon production processes that a Monte Carlo photon transport code must simulate. From Williamson (1988d) with permission. (b) 2-D illustration of a photon history in a plane bisecting a low-energy seed.
transport into the simulation reduces the efficiency of the calculation by two orders of magnitude, limiting the simulation to relatively simple 1-D and 2-D geometries. In contrast, Monte-Carlo photon-transport (MCPT) problems of great complexity can be readily solved on PCs or workstations within a few hours. MCPT approximates absorbed dose in medium by collision kerma, which is a valid only when secondary charged-particle equilibrium (CPE) obtains. Based on theoretical considerations, Roesch first hypothesized that inverse-square law gradients could induce a state of electron disequilibrium giving rise to a buildup effect at short distances from brachytherapy sources wherein collision kerma could significantly overestimate absorbed dose (Roesch 1958). Several recent Monte Carlo studies (Luxton and Jozsef 1999; Nath, Yue, and Liu 1999; Wang and Li 2002) have confirmed this hypothesis. For example, Luxton (Luxton and Jozsef 1999) compared buildup factors calculated by EGS4 with those calculated by Berger (1968). For 600 and 800 keV sources, the dose to kerma (D/K) ratio is about 0.5 at 0.5 mm distance and effects of 2% persist out to 1.5 and 2 mm, respectively. Such effects occur at lower energies (Ling and Yorke 1989). Because D/K deviations scale with secondary electron range, extremely thin scoring shells at very small distances are needed to resolve them. Wang (Wang and Li 2002) recently used EGSnrc to investigate secondary particle disequilibrium effects for two commercial HDR 192Ir sources, including the effects of encapsulation and β-ray transport as well as secondary electron transport. In the 0.5 to 1 mm distance range, the combined effects of two electron transport components was as large as 13% and 29% for the current Nucletron and old VariSource source models, respectively. The D/K ratio was 1.07 at 1 mm and essentially unity at distances ≥ 2 mm. When secondary electron transport alone was simulated D/K ranged from 1.03 to 1.08 in the 0.5 to 1.0 mm range. Thus simulation of both photons and secondary charged particles is important for clinical applications, e.g., intravascular brachytherapy, in which near-zone doses are important. In addition, the older orthovoltage and 60Co dosimetry literature (Dutreix and Bernard 1966; Gray 1940; Paterson 1942) documents that significant deviations from D/K = 1 occur at media interfaces, e.g., metal applicator-tissue interfaces. In addition to calculating clinically meaningful doses to soft-tissue cavities in bone, or mucosa in contact with metal applicators, secondaryelectron disequilibrium effects need to be considered when simulating the response of thin detectors. Aside from these important exceptions, we can conclude that secondary-electron transport and CPE-failure effects are relatively unimportant for interstitial and intracavitary brachytherapy, and that neglecting these effects, as does MCPT simulation, leads to less than 1% errors in most clinical settings.
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Technical Aspects of Monte-Carlo Photon-Transport Simulation in Brachytherapy Accurate and reliable MCPT-based dosimetry requires attention to many technical details such as use of appropriate physical models for photon interactions and cross-section library; accuracy of the underlying geometric model; and use of appropriate estimators and other variance-reduction techniques. Each of these issues will be reviewed in the following sections. Physics of Photon Scattering and Choice of Cross-Section Library As noted above, a MCPT code requires total cross sections for each photon collision process, tabulated as a function of incident photon energy and medium. In addition, mass-energy absorption coefficients maybe needed to convert energy fluence into estimates of absorbed dose in the charged-particle equilibrium approximation. Many public-domain codes (e.g., ITS and MORSE) neglect the influence of electron binding on photon scattering, approximating coherent scattering by uncollided primary photons (by subtracting µcoh from µtot) and incoherent scattering by the Klein-Nishina differential cross section, which assumes that the target electrons are free and at rest. Electron binding corrections are usually made by multiplying the Thompson and Klein-Nishina differential cross sections by atomic form factors (AFFs) and incoherent scattering factors (ISFs), respectively, giving rise to the incoherent and coherent scattering cross sections. The tabulations of AFFs and ISFs used by up-to-date cross section libraries are those of Hubbell, et al. (Hubbell and Øverbø 1979; Hubbell et al., 1975). Williamson (Williamson, Deibel, and Morin 1984) suggests that the free-electron approximation Klein-Nishina (KN) is of questionable accuracy below energies of 100 keV. The binding of target atomic electrons to nuclei significantly modifies the angular distribution of scattered photons, particularly in high-atomic number media. Although binding effects have less influence on energy deposition, the free-electron approximation was found to over estimate energy fluence transmitted through thick barriers by 2% to 15% in the low-energy range. A more clinically relevant comparison (Figure 13), shows that the choice of collisional physics
Figure 13: Ratio of transverse-axis dose for the model 6711 125I seed computed by various photon interaction models, to dose computed based upon incoherent scattering using the Hubbell’s incoherent scattering functions and coherent scattering in the atomic form-factor approximation. ISF only = incoherent scattering with coherent scattering omitted; KN only = free-electron Klein-Nishina scattering; KN + CS = electron Klein-Nishina scattering with coherent scattering; IA + CS and BIA + CS = coherent scattering with two approximations to Compton Doppler broadening. (unpublished data: C. Costescu and J. Williamson).
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model modifies model 6711 transverse-axis dose distribution by less than 2% at distances less than 4 cm. An interesting conclusion is that KN alone is much more accurate than KN with coherent scattering, the model used with EGS4 with KEK extensions. For 20 keV point sources and for doses near the longitudinal axis of 125I seeds, Costescu’s and Williamson’s unpublished data shows much larger 5% to 8% variations with model choice. Incorporating the more complex Doppler broadening model of incoherent scattering (broadening of the discrete Compton scattering energy at each angle into a continuous spectrum due to the orbital electron momentum distribution) influences brachytherapy dose distributions by less than 2%. However, mass-energy absorption coefficients used to convert energy fluence into dose must be consistent with the physical model used to transport photons to avoid systematic errors. For example, if KN only is used to transport photons, both the µ and µen/ρ tables must be recalculated for the KN model. Below 100 keV, both the authors and the AAPM recommend that coherent scattering be explicitly simulated and that the Compton free-electron distribution be corrected for orbital-electron binding effects via ISF factors. If for reasons of efficiency, a less accurate model, e.g. KN-only, must be used, this approach should be carefully benchmarked against the standard photon-collision model for the domain of interest. For liquid-water medium, biological tissues, and certain plastics, coherent scattering, in contrast to other interaction mechanisms, is strongly influenced by molecular bonds and inter-molecular forces. For water, these extra-atomic forces increase the value of the coherent-scattering cross section by 20% to 40% in the 10 keV to 40 keV energy region compared to the predictions of the mixture-rule model, which accounts only for atomic binding forces. For liquid-water, Morin’s (1982) AFF tabulation is still the most comprehensive tabulation. An important secondary scattering mechanism in brachytherapy is characteristic x-ray production following ejection of an orbital electron. While any collision process that ejects an orbital electron can initiate the atomic relaxation process and the resultant cascade of characteristic x-rays and Coster-Konig or Auger electrons, photoelectric collisions in medium- and high-atomic number materials dominates. For example, approximately 25% of the photon fluence emitted by the model 6711 125I seed is due to emission of characteristic x rays within the silver wire upon which the radioactive material is adsorbed (Ling et al., 1983; Williamson 2002). The significant contribution of low-energy (4.5 keV) characteristic x-rays emitted by the titanium encapsulation of 125I seeds to the free-air exposure rate has been experimentally observed by Kubo (1985) and theoretically demonstrated by Williamson (1988b) to significantly influence the 1985 NIST iodine seed calibration standard (Loftus 1984). Using a simplified atomic relaxation process and tabulated vacancy probabilities, transition probabilities, and fluorescent yields of Plechaty (Plechaty, Cullen, and Howerton 1978), Williamson has shown good agreement between his Monte Carlo calculations and spectroscopic measurements by NIST for Ag characteristic x-rays (Williamson 2002) and by Kubo (1985) for Ti x-rays. Recently, the more up-to-date and complete Lawrence Livermore Laboratory EADL data base (Perkins and Cullen 1991) has become available. EADL tabulates all radiative and nonradiative transitions for each subshell of each element from Z=1 to Z=100. Future low-energy brachytherapy simulations should utilize EADL-compatible data. Simulation of atomic relaxation in the EGSnrc code is now based upon EADL data (Kawrakow and Rogers 2003). Accurate MCPT-aided brachytherapy dosimetry would not be possible without the evolution of accurate photon cross-section libraries made over the last 30 years (Hubbell 1999). The success of these efforts owes much to John Hubbell of NIST who has systematically compiled and critically evaluated both measured and theoretically calculated cross-sections to provide evermore accurate and comprehensive cross-section libraries. During this evolution, photo ionization cross sections for light elements in the 20 to 40 keV energy range have changed significantly. Several investigators (Bohm, DeLuca, and DeWerd 2003; Demarco, Wallace, and Boedeker 2002; Williamson 1991) have demonstrated that 103Pd and 125I brachytherapy dose distributions are quite sensitive to choice of cross-section library and that
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use of obsolete cross sections can introduce errors of 5% to 10% or more. Hence we will briefly review the evolution of cross-section libraries and their use in popular public-domain codes. Virtually all modern libraries use very similar scattering cross sections, based upon the Hubbell’s nonrelativistic incoherent scattering and relativistic atomic form factors (Hubbell and Øverbø 1979; Hubbell et al., 1975) as described above. The major differences from one library to another reside in the origin of photo ionization cross sections. In describing the different libraries, we will use the Radiation Safety Information Computational Center (RSICC) Data Library Code (DLC) nomenclature for distinguishing the different libraries (see Table 3). RSICC, located at Oak Ridge National Laboratory, packages and distributes transport codes and data libraries for nominal fees. With regard to treatment of photo ionization, three broad groups of libraries are currently used in public-domain codes. The Storm and Israel (1970) photoelectric cross sections, DLC-15, are based upon the theoretical calculations of Rakavy and Ron (1967) using screened atom models of Schmickley and Pratt (1967). These cross sections continue to be used by EGS4 and EGSnrc, despite up to 4% deviations in the low-energy brachytherapy range relative to currently-recommended libraries (Figure 14). The second family of libraries, represented by DLC-7F, is based upon the semi-empirical photo-ionization cross sections of McMasters et.al. (1969), which were derived from least-squares curve fitting to an extensive collection of theoretical and measured data (Ca 1968) weighted according to their assessed accuracy. The experimental cross-section data had generally been obtained by subtracting the theoretically calculated scattering cross sections from the measured total cross sections. These photoeffect cross sections underestimate absorption by nearly 10% in the 20 to 30 keV energy range in water (Bohm, DeLuca, and DeWerd 2003; Demarco, Wallace, and Boedeker 2002; Williamson 1991) relative to more current libraries. Despite these large errors, these cross sections have been used in the MCNP cross-section libraries DLC-200 and DLC-189 (MCNP libraries MCPLIB01, MCPLIB02, and MCPLIB03) since 1982. The third, and currently recommended library family introduced in 1983, uses the theoretical photo-effect cross sections of Scofield (1973) based upon a relativistic Hartree-Slater model renormalized to the Hartree-Fock model for low atomic-number
Table 3. Major Characteristics of Photon Cross-Section Libraries Distributed by RSICC DLC designation
Year of origin
Data sources
Client Codes
DLC-15
1971
Storm and Israel. Older theoretical treatments of photo-ionization and scattering
EGS4, EGSnrc
DLC-7F
1975
Current scattering and semi-empirical MCNP 4b and 4c McMasters photo ionization cross-sections
DLC-189 DLC-200
1997 2000
RSICC MNCP libraries. DLC-7F cross sections. Includes MCPLIB 01 and 02
DLC-99
1983
HUGO library. Hubbell form factors, normalized Scofield photo-ionization cross sections
DLC-146
1989
Same as DLC-99 except photo-ionization normalization dropped
DLC-136 DLC-174
1989 1997
NIST PHOTX and XCOM databases. Identical to DLC-146
DLC-179
1997
EPDL97 library including EADL atomic relaxation data. Identical to DLC-146, except subshell cross sections included
MCNP 4b and 4c PTRAN prior to 2000
PTRAN after 2000
MCNP library MCPLIB04 (LANL website) since 2002
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Figure 14. (Left) Ratio of photo-ionization cross sections from various libraries to current NIST XCOM database for water as a function of photon energy. (Reprinted from Phys Med Biol, vol 47, “An analysis of MCNP cross-sections and tally methods for low-energy photon emitters,” J. J. Demarco, R. E. Wallace, and K. Boedeker, pp.1321–1332. © 2002, with permission from IOP Publishing, Bristol.) (Right) Ratio dose of dose rate on transverse-axis of a Model 2000 103Pd seed as calculated by MCNP/DLC-179 to MCNP/DLC-200. (Reprinted from Med Phys, vol 30, “Brachytherapy dosimetry of 125I and 103Pd sources using an updated cross section library for the MCNP Monte Carlo transport code,” T. D. Bohm, P. M. DeLuca Jr., and L. A. DeWerd, pp. 701–711. © 2003, with permission from AAPM.)
elements (Das et al., 1997b). Extensive comparisons (Saloman and Hubbell 1986, 1987) of Scofield’s calculations with the NIST experimental cross-section data base, including many recent “direct” measurements of photo-ionization, demonstrated that this theoretical approach was more accurate than competing approaches. The initial library, DLC-99 was used in the first author’s code, PTRAN_CCG through 2000. Further intercomparisons revealed that unnormalized Scofield calculations further improved the match between tabulated and measured cross sections Hubbell (1999). Hence, the current generation of crosslibraries (DLC-146, 136, and 189), introduced in 1989, incorporate this improvement. Below 1 MeV in water, dropping the normalization increases the photoelectric cross section by 2.7%. Essentially, all current libraries, including the NIST PHOTX/XCOM and EPDL97 tabulations all use the same theoretical approaches to calculating scattering and photo-ionization cross sections. The LANL website indicates that an MCNP EPDL97-equivalent library, MCPLIB04 has been available since 2002. Figure 14 (left) compares the three families of photoelectric cross sections to the current NIST XCOM cross sections. Mass-energy absorption coefficients, compatible with post-DLC-146 libraries, have been published by Seltzer (1993) and can be downloaded from the NIST website1. Several papers demonstrate that use of obsolete cross section libraries significantly compromises the accuracy of 125I and 103Pd TG-43 dosimetric parameters. Bohm et al. (Bohm, DeLuca, and DeWerd 2003) demonstrated that use of default MCNP cross sections (DLC-200), overestimates dose by 4% at 1 cm to 20% at 5 cm. For analog dose estimation, DeMarco et al. (Demarco, Wallace, and Boedeker 2002) show similar results (4% to 8% errors in dose-rate constant estimation in 20 to 40 keV energy range) when default MNCP cross sections are used and somewhat smaller (2.5%) errors when default EGS (DLC-15) cross sections are used. DeMarco also demonstrates that large and unpredictable dose-estimation errors arise when mass-energy absorption coefficients based upon DLC-146 are used to convert particle fluence scores into dose for MCNP simulations using DLC 15 or 200 for constructing histories. Hedtjarn et al. (Hedtjarn, Carlsson, and Williamson 2000) shows that the smaller differences between 1
http://physics.nist.gov/PhysRefData/XrayMassCoef/cover.html.
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the DLC 99 and 146 libraries introduce dose-calculation errors ranging from 1% at 1 cm to 3.5% at 5 cm for an 125I seed. The updated TG-43 report recommends (Rivard et al., 2004) using DLC146/XCOM/EPDL97-equivalent cross sections for reference-quality dosimetry and will not accept as candidate data sets Monte Carlo simulations based upon default MCNP cross sections. Geometric Modeling and Validation Accurate MCPT dose calculation requires a flexible and general system of geometric modeling. Brachytherapy sources and applicators have complex internal designs often leading to highly anisotropic dose distributions. In addition, dose distributions are sensitive to the shape and size of the surrounding scattering medium. Calculation of absolute dose rates may require simulation of the experimental geometry used to standardize air-kerma strength for the source. Finally, comparison of MCPT predictions with measured results may require simulation of detector response. Simulation of detectors such as commercially available silicon diodes poses a difficult modeling challenge, since these detectors are small and may have intricate internal structures. The approach of choice is surface- (Briesmeister 1997) or volume (Li and Williamson 1992)-based combinatorial modeling. This approach, sometimes called complex combinatorial geometry or CCG, defines complex spatial regions as set-theoretic intersections, unions and differences of elementary volumes such as cuboids, ellipsoids, elliptic cylinders, cones and half-planes. The modeling code must allow the composition, location, and dimensions of each geometric region to be independently specified, must allow complex structures to be nested inside of one another, and must support point classification, line-segment classification, and ray tracing. Most of the public-domain codes support CCG modeling packages. Exceptions are EGS4 and EGSnrc (Kawrakow and Rogers 2003) which effectively limit users to modeling relatively simple cylindrically symmetric and or 3-D Cartesian slab geometries supported by the user packages DOSrz and DOSxyz. Nevertheless, a number of high- and lowenergy sources can be adequately modeled by the DOSrz package. Of the codes widely used in brachytherapy, MCNP has the most general and sophisticated geometric modeling capability. Because low-energy seed dose distributions are so sensitive to the assumed internal structure of seed, it is essential to validate through measurements seed and, by extension, applicator and detector geometries. Both authors have found vendor-supplied mechanical drawings of sources to be misleading and sometimes inaccurate. Tools for verifying seed geometry include contact transmission radiographs, micrometer measurements, microscopic examination of unassembled seed components, and pin-hole autoradiography (Kirov et al. 1995b). Some investigators (Hedtjarn, Carlsson, and Williamson 2002; Rivard 2001) have used milling machines to longitudinally “slice” open seeds along their symmetry axis. Figure 15 illustrates three investigations in which such measurements significantly influenced the geometric model adopted. In Figure 15(a), transmission contact radiographs indicated that the inverted cup ends were spherical in shape rather than cylindrical as represented in the vendor’s literature. Examination of unassembled end-weld components in a dissection microscope enabled the authors to devise a detailed model (Monroe and Williamson 2002). In the case of the DRAXIMAGE LS-1 seed, Figure 15(b), such radiographs documented that the gold localization rod and the two radioactive spheres were mobile, assuming different configurations depending on the seed orientation. Because the thickness of hemispherical end caps (see Figure 15(c)) is thinner (0.05 mm) than the cylindrical sections of the capsule (0.10 mm), the motion of the radioactive beads significantly modulated the seed output. For clinical use, the first author (Williamson 2002) recommended a weighted average of dose distributions corresponding to different geometric configurations. This strategy was first proposed by Rivard (2001), who noted (Figure 15(d)) that the radioactive and nonradioactive spherical beads in the North American Scientific 125 I seed (Model MED3631-A/M) assumed a staggered array configuration that deviated from cylindrical symmetry. Impact of this effect using this principle on non-uniformly loaded sources was examined further (Rivard, Kirk, and Leal 2005).
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(a)
(b)
(c)
(d)
Figure 15. (a) Contact radiograph and final geometric model of the Theragenics Model 200 103Pd seed (Monroe and Williamson 2002); (b) contact transmission radiograph and final geometric model of the DRAXIMAGE LS-1 125I seed with its axis vertical illustrating mechanical shift of the internal structures (Williamson 2000); and (c) 200X magnified electron microscopy image of the end of a Ag rod from a Model 6711 125I seed illustrating that it illustrating previously unappreciated beveled rather than right circular cylinder shape. (d) Geometry of the model 3631-A/M seed when aligned vertically, illustrating the staggered configuration of the internal spherical elements and a cut-away photograph of the seed (in horizontal orientation) that has been mechanically sliced open (Rivard 2001).
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Choice of Estimator and Other Variance Reduction Techniques An estimator is a computational device for extracting a statistical estimate of the quantity of interest from a simulated photon trajectory. Since MCPT simulations assume that energy transferred from photon fields to matter is absorbed locally, the quantity of interest in brachytherapy is collision kerma. Kerma rates calculated by MCPT codes are generally expressed in units of cGy/simulated primary photon or history. Using the photon spectrum assumed by the simulation, it is straightforward to convert kerma/emitted primary photon into kerma rate per unit contained activity, represented by ∆K = K" /A with
cont
M C PT
units of cGy⋅mCi ⋅h . Thus the source strength specification quantity intrinsic to the MCPT method is contained activity. In the sections below, the process of normalizing MCPT ∆K estimates to clinically relevant source-strength specification quantities will be described. The choice of estimator and method of implementation depends on whether dose-at-a-geometric point or dose averaged over an active detector volume is desired. Simulation efficiency (statistical precision achieved for a given CPU time), the proximity of media boundaries to the point or volume of interest, and the spatial resolution desired are also important considerations. Each of the sections below will review the estimators of importance to brachytherapy and their principal applications. The simplest device for improving spatial resolution and computational efficiency, especially at short distances, is to deterministically calculate the primary-photon collision kerma component rather than to stochastically score randomly constructed primary-photon trajectories. This reduces the variance of the primary dose component to zero. The calculation consists of straightforward 3-D numerical integration # # − ∑ µ ⋅% # −2 ⋅ (µen /ρ )⋅ r , over the distribution of radioactivity, ρ(r), in terms of of the operator, ρ(r)⋅ e –1
−1
mCi/cm3. By limiting Monte-Carlo simulation to scoring of scattered-photon trajectories, the variance of the total-dose estimates near sources can be reduced by a factor of two. Numerous other variance-reduction techniques have been developed and implemented which can greatly reduce the CPU time required to achieve the desired statistical precision for complex problems. Common strategies (Jenkins, elson, and Rindi 1988; Lewis and Miller 1993; Williamson 1988c) include Russian Roulette and splitting, forced scattering, and biased angle sampling. Each of these techniques uses biased sampling distributions in conjunction with correction factors designed to eliminate any bias in the sample mean of the quantity of interest. These sampling distributions and photon-weight correction factors are cleverly designed so as to force the computer to preferentially focus CPU effort on simulated random events that contribute significantly to the quantity of interest rather than those that do not. Spherically Symmetric Tracklength Estimator. This estimator is applicable only to spherically symmetric 1-D problems in which the collision kerma rate depends only on the radial distance, r, from the origin. The most commonly studied problem is that of an isotropic point source positioned at the origin in unbounded homogeneous medium. ∆K(r) is estimated by tallying those photon trajectories that cross an imaginary spherical detector surface of radius r centered on the source. A complete technical description is given by Williamson (1987). Major advantages of this approach are simplicity, computational efficiency, and that it gives an unbiased estimate of the dose rate at a geometric point. Freedom from volume-averaging artifacts follows from the fact that an infinitely thin scoring surface rather than a finite-thickness scoring volume is used to simulate a point detector. Many of the early MCPT studies in brachytherapy, e.g., Dale (1983), are based upon the spherically-symmetric tracklength estimator. Analogue and Tracklength Estimators. Most MCPT applications utilize so-called analog estimators to estimate absorbed dose or collision kerma. This approach consists of tallying energy transfers from those photon collisions that occur in an detector volume, V, centered about the point of interest. The resultant estimated dose is simply the quotient of energy transferred to the detector volume by its # mass. Thus for collisions j such rj ∈V, the score is (Wj−1Ej−1 – WjEj)/ρV where Wj is the weight of the
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photon. In MCNP, analogue scoring is called the F4 tally. In heterogeneous geometries, care must be taken to ensure that the detector contains homogeneous medium, i.e., that its boundaries do not intersect media interfaces. To use analog estimation to evaluate dose at a point, the detector volume is centered on the point of interest and dose at the detector center approximated by dose averaged over its volume. This approach must be used cautiously in brachytherapy. If dose gradients are large relative to the detector dimensions, spatial resolution will be lost due to volume averaging. If the detector volume is constrained to be small by nearby media interfaces or the need for high spatial resolution, efficiency may be unacceptably low since only a small fraction of the simulated collisions will occur within the detector. The tracklength estimator, which, like the analogue estimator, also gives an unbiased estimate to the energy absorbed by the detector volume, V, greatly increases simulation efficiency by analytically estimating the energy transferred to the detector volume by all photon flight paths that cross the detector (Williamson 1987). Where ∆% is the distance traveled in V by a photon of energy Ej and weight Wj, the score is then W j E j ⋅ ∆% ⋅ ( µ en / ρ ) V . MCNP calls this estimator the *F8 tally. An even more efficient variant of tracklength scoring, expected-value tracklength estimation (Williamson 1987), scores all photons whose projected trajectories intersect the detector. In simple single spherical detector geometries, use of tracklength scoring can improve efficiency by factors of 7–10. In CT voxel detector grids, containing many thousands of detectors, expected-value tracklength scoring increases efficiency by factors of 20–50 relative to analogue scoring if an efficient voxel tray tracing algorithm (Siddon 1985) is available. If only scattered-photon trajectories are stochastically scored, volume averaging bias (average dose in V/dose at geometric center of V) for spherical detectors will be limited to 1% if the detector radius is limited to 60% of the source-detector distance or 15% of the primary photon mean-free path, whichever is smaller (Williamson 1987). Because of its greater efficiency, tracklength scoring should always be used in preference to analogue scoring in photon-transport simulations. The most important application of the volume detector estimators is simulation of detector response in brachytherapy dosimetry. In this situation, one wants to model detector artifacts such volume averaging and medium displacement. Thus an analogue or tracklength estimator is required. To simulate the response of a TLD dosimeter placed at a small distance d from a brachytherapy source, a LiF-filled rectangular prism detector having the same # dimensions as the actual TLD should be used. This MCPT simulation will yield the quantity ∆K det (r ) ; the energy transferred to a TLD-size cube#of LiF per unit # mass of LiF centered at r per contained mCi·h. By using analog estimation, the ∆K det (r ) estimate will include the effects of volume averaging, directional anisotropy, and self-attenuation. To correlate this theoretical endpoint with the measurable response of the TLD, the fundamental assumption, # # TLlin (r ) ∝ ∆Ddet (r ) , (see equation (5) must be made. Using Burlin cavity theory, as described above, to # # convert ∆K det (r ) into ∆Ddet (r ) , the relative energy response [equation (7)] and other quantities needed to convert detector readings into dose rates can be evaluated by Monte Carlo simulation. Analytic Point-Dose Estimators. In evaluating reference dose-rate distributions for brachytherapy sources, in contrast to simulating detector response, the quantity of interest is dose at a geometric point. In this application, the volume averaging characteristic of tracklength or analogue estimation presents a source of artifact no different than that faced by the experimentalist. Analytic estimators seek to overcome this conflict between spatial resolution and computational efficiency by forcing every simulated collision to contribute to the quantity of interest at the desired geometric point. This is done by analytically estimating the probable energy transfer contribution made by each simulated photon collision to each of the specified target points. Although analytic point-dose estimators are very complex to implement and greatly increase the computing time required/simulated collision, overall efficiency is often dramatically improved by increasing the number of simulated events that contribute to the sample mean. The simplest analytic point-dose scoring technique, next flight estimation (Kalos 1963), has been used successfully by the author and his colleagues for many years in computational brachytherapy dosimetry
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(Williamson 1987). For each simulated collision, next flight calculates the probability that the scattered photon emerging from the collision intersects the point detector and transfers energy to the medium at its location. This requires correction for loss of scattered photon flux emerging from the collision site by inverse square-law and attenuation through the intervening media separating the collision and the detector. Unfortunately the (collision-to-detector distance)−2 singularity in the estimator destroys the statistical stability of the dose estimate, since a single collision occurring close to the detector may make a huge but random dose contribution. Although the history-averaged estimate is bounded and unbiased, the second moment of the distribution does not exist, the variance of the estimate cannot be computed, and the simulation converges slower than central limit theorem (Kalos 1963). This problem can be solved by replacing the e−µr r−2 term in the estimator by its average value within a small critical sphere surrounding the point detector (Williamson 1987). This modified “bounded” next-flight estimator solves the problem of accurately calculating kerma at a point in a dilute medium such as air found in calibration geometries. In addition, bounded next flight works well in condensed media so long as the point of interest remains 1 to 2 mm from any media interface and the critical averaging sphere encloses a single medium. It significantly increases efficiency for low-energy sources and is the estimator of choice for 125I dosimetry. Williamson (1987) has given a rules of thumb for the critical radius, R: R should be set at 30% of the source-to-detector distance or 1 mean-free path (whichever is smallest) to ensure bias of less than 1% when primary dose is computed analytically. Although bounded next flight scoring is very powerful, it has several limitations. To eliminate the 1/r2 singularity, the detector point must be enclosed by a sphere of homogeneous medium. If this sphere is made too small, the variance will not converge (will not decrease with in proportion to number of simulated histories). Thus the standard deviation calculated over the number of histories simulated cannot be used to predict statistical precision of the mean dose rate, and the dose rate estimate will not be reliable. Problems in which media interfaces severely constrain the averaging radius, e.g., dose near an applicator-tissue interface and dose to a small detector, cannot be solved with bounded next flight estimation. A more powerful generalization of next flight estimation, once-more collided-flux estimation (OMCFE) (Kalli and Cashwell 1977), was modified and validated for use in brachytherapy MCPT simulations by Li et al. (Li, Williamson, and Perera 1993). This method avoids the need for averaging over a homogeneous region, allowing the point detector to be placed arbitrarily close to a media boundary. Whenever a photon collision occurs near the point of interest and results in an excessively high next-flight score, the position and trajectory of this collision are randomly resampled from biased probability distributions and next-flight estimation applied to the resampled collision site. The biased probability distributions are cleverly chosen so that the resultant weight correction factor cancels out the 1/r2 singularity. The result is an unbiased estimate of kerma at a point and a 30-fold improvement in efficiency compared to analog estimation applied to typical detector sizes. An important application of OMCFE is calculation of ∆Kdet for small TLD or diode detectors at large (>20 mm) source-detector distances where the efficiencies of analog and tracklength estimation are poor, volume averaging artifacts are small, and dose to the detector volume is well approximated by dose at its geometric center.
Application of the Monte Carlo Method to Calculation of Reference-Quality Dose-Rate Distributions # To calculate clinically relevant absolute dose rates, D" wat (r ), in water or other condensed media, the native # MCPT normalization of ∆K wat (r ) , dose per simulated collision (or equivalently dose per contained mCi-h), must be changed to air-kerma strength. This is achieved by running two simulations: one simulation with the seed model located at the center of a 30 cm liquid-water phantom for the purpose of calculating the kerma-rate distribution, ∆K wat (r, θ ), and a second simulation in which the seed model is placed in a vacuum or large air sphere for the purpose of calculating the air kerma-strength per mCi-h
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(or history), ∆SK. Given these two quantities, appropriately normalized absolute dose rates can be estimated:
[ D"
wat
(r, θ ) SK ]MCPT = BRx
∆K wat (r, θ ) ∆SK
(17)
assuming that ∆Dwat(r,θ) ≈ ∆Kwat(r,θ) is a sufficiently accurate approximation. Calculation of the various TG-43 dose ratios is straightforward:
Λ=
∆K wat (r = 1 cm, θ = π / 2 )
gL (r ) =
∆SK
(18)
∆K wat (r, π / 2 ) ⋅ GL (1 cm, π / 2 ) ∆K wat (1 cm, π / 2 ) ⋅ GL (r, π / 2 )
To calculate the distribution, ∆Kwat(r,θ), the senior author (JFW) typically calculates polar angle dose distributions at distances r = 0.25, 0.5, 0.75, 1.0, 1.5, 2.0, 3.0, 4.0, 5.0, and 7.0 cm using an angular mesh with 1° increments near the longitudinal and transverse axes to a maximum of 5° increments at intermediate angles. For transverse-axis dose distributions, a fine grid (0.1 to 0.2 cm increments) is used in the 0.1 to 1 cm distance range, and 0.5 to 2.5 cm steps used for larger distances. The recent TG-43 report specifies minimum spatial resolutions and ranges for Monte Carlo calculations (Rivard et al., 2004). Typically, the bounded next flight estimator is used at distances greater than 0.5 cm and the OMCFE at shorter distances to avoid intersection of the scoring spheres with the seed structure. Simulation of 500,000 to 20,000,000 histories, depending on distance and radionuclide, is sufficient to guarantee statistical precisions better than 0.5% near the source and 1.5%-2% far from the source. Generally, the author always calculates primary kerma rate deterministically. There are two approaches to estimating the effective air-kerma strength constant, ∆SK, for a source: the point-extrapolation method and the WAFAC method. In the more widely used “point-extrapolation” approach (Hedtjarn, Carlsson, and Williamson 2000; Williamson 2000), the seed model is placed at the center of 2-to 5-m diameter sphere of dry air. Then, using the bounded next flight estimator, the transverse-axis dose-rate distribution, ∆Kair(d), is calculated over the 1.0 to 150 cm at distances range. To ensure that these results are not contaminated by low energy characteristic x-rays from Ti, Ti characteristic x-ray production should be suppressed or a 5 keV energy cutoff employed by the estimator. Then, the quantity (∆SK)extr is evaluated by fitting the ∆Kair(d) data to the following equation:
∆ K air (d ) ⋅ d 2 = ( ∆SK )extr ⋅ (1 + α d + β d 2 ) ⋅ e− µ d
(19)
where the fitting parameter, µ, describes primary photon attenuation in air, while the best-fit parameters, α and β describe the buildup of scattered photons. This idealized geometry excludes all scattering surfaces and approximates the air-kerma detector by a geometric point, as specified by the definition of air-kerma strength. Figure 16(b) illustrates the fit of equation (19) to Monte Carlo data for the three seed geometries described by Figure 16(a). In the case of the Model 200 seed, there is a dramatic breakdown of inverse-square law that persists out to 20 cm distance. At long distances, oblique filtration through the thin palladium metal layer reduces the contribution of radioactivity distributed in the four circular end surfaces of the two carbon pellets. At shorter distances, primary photon trajectories are less oblique
15–Quantitative Dosimetry Methods for Brachytherapy
(a)
(b)
(c)
(d)
277
Figure 16. (a) Assumed geometry of the model 200 103Pd seed, the model 6702 125I seed, and the model 6711 125 I seed. The two graphite pellets inside the model 200 source are thought to be coated with palladium metal 2 µm in thickness for current accelerator-produced product. The model 6711 seed contains a solid Ag right circular cylinder, which is coated with radioiodine; (b) Fit of equation (19) to the transverse-axis air-kerma distribution; (c) the polar air-kerma profile in air at 30 cm distance; (d) the 2-D anisotropy function at 1 cm distance in liquid water.
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and traverse thinner layers of Pd, causing air-kerma to increase faster than inverse square law (Monroe and Williamson 2002; Williamson 2000). Screening of radioactivity as illustrated by Figure 16, can occur whenever radioactivity is dispersed within or on the surface of a high-density core with sharp corners or edges. When this phenomenon occurs, the polar dose profile may contain discontinuities or become extremely sensitive to alignment of the interior seed components. As illustrated by Figure 16(c), radioactivity screening introduces significant polar anisotropy (“overshoot”) near the transverse axes of the models 6711 and 200 seeds. This invalidates use of the point-extrapolation method since the WAFAC chamber integrates over a 16° cone centered on the transverse plane. As shown by Monroe and Williamson (Monroe and Williamson 2002; Williamson 2000), use of (∆SK)extr introduces a 15% variation into Λ with respect to Pd metal layer thickness and worsens the agreement with Λ measurements. Theoretically, the WAFAC appears not accurately realize air-kerma strength for sources with transverse-plane anisotropy, since SK is defined in terms of kerma to a geometric point. On the other hand, had source strength been specified in terms of point-kerma measurements, the calibration uncertainty would have been much larger. In cases where radioactivity screening introduces real or apparent anisotropy near the transverse plane, the AAPM (Rivard et al., 2004) recommends using the WAFAC method of ∆SK estimation, in which more of the geometric detail of the NIST primary standard is explicitly modeled in the simulation. As described in more detail in papers from Williamson’s group (Monroe and Williamson 2002; Williamson 2000, 2002), the WAFAC collimator, shielding barriers, and surrounding walls, floor and ceiling are added to source-air sphere geometry. As shown in Figure 17, the WAFAC chamber is operated in two active
Figure 17. Side view if the Wide Angle Free-Air Chamber (WAFAC) illustrating use of the short 11 mm and long 153 mm cylindrical electrodes. The chamber is cylindrically symmetric about the axis bisecting the seed center, which is placed on a rotating holder during measurements to average over any equatorial anisotropy. The primary collimator located just upstream of the Al filter is not illustrated.
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volume configurations, denoted by the subscripts “153” and “11,” which correspond to the length of the active ionization volume in mm as defined by the separation of its two circular electrodes. The subtraction of the charge measured by the 11 mm long chamber, from that measured by the full-length chamber geometry is intended to reduce the effects of wall scattering and electric field non-uniformity (Seltzer et al., 2003). Using an expected value track-length estimator, simulation of 25 million histories typically yields ab estimates of ∆ Et , the energy absorption rate within the WAFAC collection volume of length t, with statistical precisions of 0.3%. In this case, both primary and scattered photons must be scored stochastically, in order to include volume averaging and other affects. Then (∆SK)WFC can be evaluated by
( ∆SK )WFC =
( ∆ E ab − ∆ E ab )d 2 153
11
ρair (V153 − V11 )
kinv ⋅ katt = ( ∆ K air ⋅ d 2 )WFC ⋅ kinv ⋅ katt
(20)
where ρair is the density of air and kinv and katt are inverse square-law and attenuation corrections defined by Seltzer (Seltzer et al., 2003). Because Ti x-ray production is suppressed, no correction of the Al filter used by NIST is needed. Similarly, all of the other correction factors required by NIST’s WAFAC measurement protocol are unity by virtue of the Monte Carlo simulation design. The factor kinv is needed because the aperture is a flat plane rather than a spherical element. Thus, the transverse-axis fluence at distance d is underestimated by the fluence averaged over the WAFAC aperture area. By assuming an isotropic point source, the photon fluence can be analytically averaged over the aperture area yielding (Daskalov and Williamson (2001) −1
kinv
R 2 π R 2 2π %d % R2 . 2 = 2 ⋅ ∫ 2 = R d ⋅ ln 1 + 2 2 d d 0 d + %
(21)
The factor katt in (20) corrects for attenuation between the source and the aperture and within the collecting volume. It is evaluated by exploiting the fact that point-extrapolation and WAFAC ∆SK simulations yield identical results for any source with an isotropic fluence distribution over the WAFAC collection volume. Both WAFAC and point-extrapolation simulations are performed for a point source of 103Pd or 125I, yielding values of (∆SK)extr and (∆Kair⋅d2)WFC. Since, (∆SK)extr = (∆SK)WFC, katt is given by
katt =
( ∆SK )extr
for a poinnt source
(22)
kinv ⋅ ( ∆ K ⋅ d 2 )WFC Equation (22) yields katt values of 1.023 (Monroe and Williamson 2002) and 1.013 (Williamson 2002) for 103Pd and 125I, respectively, which are very close to the values of 1.028 and 1.013 published by NIST (Seltzer et al., 2003). Using the WAFAC method of SK normalization, excellent results have been achieved by the Williamson and his group. The Model 200 103Pd seed dose-rate constant is nearly independent of palladium metal layer thickness and is close to the theoretical unfiltered point-source Λ (Monroe and Williamson 2002; Williamson 2000). The final calculated Λ for the current light-seed geometry, (0.691) is very close to measured value (0.68) measured by Nath et al. (2000). For the model 6711 seed, the
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WAFAC method lowers the computed Λ by 3.7% to 0.925 compared to 0.959 for the point-extrapolation method (Williamson 2002). Uncertainty of Monte Carlo Reference Dose-Rate Distributions Relatively little work on estimating the total uncertainty of Monte Carlo results has been published. Based upon partial uncertainty analyses in Williamson’s papers, the updated TG-43 report (Rivard et al., 2004) endeavored to make a more comprehensive uncertainty analysis for 125I seed transverse-plane dose calculations using NIST’s uncertainty estimation protocol (Taylor and Kuyatt 1994). The reader is referred to section Summary: Uncertainty of TLD and Role in Clinical Dosimetry of this chapter for an introduction to the NIST Report 1297 (Taylor and Kuyatt 1994) methodology. For Monte Carlo estimates of quantity, Y (= Λ or g(r), etc.), TG-43 considered the total per cent uncertainty, %σY, to consist of three sources: type B uncertainty due to uncertainty of the underlying cross sections, %σYµ; type B uncertainties arising from uncertainty of the seed geometric model, %σYgeo; and the type A statistical uncertainty, %σYs inherent to the Monte Carlo technique. Applying equation (13), one obtains:
∂Y ∂Y %σ Y2 | geo + %σ Y2 | s = % %σ µ2 + % ∂µ ∂geo 2
%σ Y = %σ
2 Y |µ
+ %σ
2 Y | geo
+ %σ
2 Y |s
where the relative uncertainty propagation factor is defined as %∂Y /∂x ≡
2
(23)
x ∂Y
. The variable x denotes Y ∂x either the cross-section value, µ, or geometric dimension, geo, of interest. The uncertainties estimated here are standard uncertainties, having a coverage factor of unity, approximating a 68% level of confidence. The influence of cross-section uncertainty was derived from Hedtjarn’s Monte Carlo data (Hedtjarn, Carlsson, and Williamson 2000) which gives Monte Carlo estimates of Λ and g(r) calculated for two different cross-section libraries, DLC-99 (circa 1983) and DLC-146 (1995). The photoelectric cross sections of the two libraries differ by about 2% between 1 to 40 keV, corresponding to a 1.1% change in µ for the mean photon energy emitted by 125I. Using these data to numerically estimate the derivative in equation (17), one obtains %∂Λ/∂µ = 0.68. Assuming that %σΛ = 2% (Hubbell 1999), then uncertainty in Λ due only to cross-section uncertainty, %σΛµ, is 1.4%. Estimation of geometric uncertainty, %σΛG, is complex and poorly understood. Each source design is characterized by numerous and unique geometric parameters, most of which have unknown and potentially correlated probability distributions. However, a few papers in the literature report parametric studies, in which the sensitivity of dosimetric parameters to specified sources of geometric variability is documented. For example, Williamson (2002) has shown that range of distances between the two radioactive spherical pellets of the DraxImage 125I source (3.50 to 3.77 mm) corresponds to a 5% variation in calculated dose-rate constant. Rivard (2001) and Williamson (2000) published similar finding for the NASI model MED3631-A/M 125I and the model 200 103Pd sources, respectively. If this phenomenon is modeled by a Type B rectangular distribution bounded by the minimum and maximum Λ values, the standard uncertainty is given by:
%σ Λ | geo = 100
Λ max − Λ min
(24)
2Λ 3
For the DraxImage source, (equation (24) yields a %σΛ|geo = 1.4%. The TG-43 report took 2% to be a reasonable and conservative estimate of %σΛ|geo. The reported statistical precision of Monte Carlo Λ estimates ranges from 0.5% for Williamson’s recent studies to 3% for Rivard’s MED3631 A/M study.
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Thus for a typical Williamson study, one obtains a %σΛ of 2.5%. Using the %σΛ|s reported by each investigator for the sources included in TG-43, %σΛ varied from 2.5% to 3.7% for the eight seeds considered in that report. The methods above can be extended to other calculation distances. As distance increases from 2 to 5 cm, progressively increases from 0.2% to 4.6%, respectively. As summarized in Table 2, the contributions from geometric uncertainty are reduced but those due to statistical and spectral uncertainties increase. The final conclusion, summarized in Table 2, is that with Williamson-level statistical uncertainties, 1-σ uncertainties of MCPT simulation range from 2.5% at 1 cm to about 5% at 5 cm. % σthe D" (r )|µ However, much more systematic study of Monte Carlo uncertainties is indicated. Additional Recommendations for Use of Monte Carlo Simulation in Preparing Reference-Quality Dose-Rate Distributions The updated AAPM TG-43 protocol (Rivard et al., 2004) makes numerous recommendations regarding use of photon-transport simulations for low-energy brachytherapy dosimetry. Recommendations discussed above in detail include use of XCOM-compatible cross-section libraries, characteristics of an adequate collisional physics model, and the need to explicitly model seed internal geometry. The report also makes recommendations on maximum statistical uncertainties and volume-averaging artifacts, the need for a formal uncertainty analysis, and documentation to be included in peer-reviewed publications. Regarding choice of codes, AAPM recommends three particular codes (EGS, MCNP, and Williamson’s PTRAN code) that have been widely used in brachytherapy and benchmarked against experimental measurements or each other. Other Monte Carlo codes and other types of transport equation solutions (multigroup codes, discrete ordinates codes, and integral transport methods) not previously used in brachytherapy dosimetry, “should be more rigorously tested and documented in the peer-reviewed literature before proposing to use their results clinically.” Perhaps the single most important recommendation relevant to Monte Carlo is best quoted directly “However, regardless of the transport code chosen and its pedigree, all investigators should assure themselves that they are able to reproduce previously published dose distributions for at least one widely used brachytherapy source model. This exercise should be repeated whenever new features of the code are explored, upon installing a new code version, or as part of orienting a new user.”
Other Applications of Transport Codes in Brachytherapy This chapter has emphasized application of Monte Carlo simulation to estimation of single- low-energy source dose-rate distributions for use in treatment planning. However, radiation transport codes have been used much more broadly in brachytherapy. We will only briefly touch upon some of these uses. • Application of Monte Carlo simulation to estimation of single-source dose distributions for 192Ir (Williamson 1995) and 137Cs (Fragoso et al., 2004; Williamson 1998)sources • Validation and parameter estimation for semi-empirical dose-calculation models (Tedgren and Ahnesjo 2003; Williamson 1996) • Characterization of shielded applicator corrections for treatment planning (Lymperopoulou et al., 2004; Weeks 1998; Williamson 1988a) • Use condensed history Monte Carlo codes in intravascular brachytherapy dosimetry • Monte Carlo dose-calculation engines for patient-specific treatment planning A number of investigators have applied Monte Carlo simulation to clinically realistic multiple-source implants with the goals of exploring the clinical impact of interseed attenuation and tissue heterogeneities
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on interstitial implants (Christensen et al., 2001; DeMarco et al., 1999; Lerma and Williamson 2002; Markman et al., 2001). To date, such codes generally require at least several days of computing time and are not easy to interface to planning images and dose-display and -analysis software found in commercial planning systems. However, several groups are developing accelerated Monte Carlo codes especially designed to facilitate efficient and more accurate planning in brachytherapy. Thus, integration of transport codes into clinical planning and delivery of brachytherapy is an important research area. Monte Carlo simulation is not the only kind of Boltzmann transport equation solver that has been investigated for use in brachytherapy dosimetry and treatment planning. Deterministic transport codes that have been successfully used in brachytherapy include discrete ordinates codes (Daskalov et al., 2000, 2002) and integral transport methods (Zhou and Inanc 2003).
Comparison of Theoretical and Experimental Dose-Rate Distributions for 125I and 103Pd Brachytherapy Sources While it is commonplace to compare LDR low-energy photon-emitting brachytherapy dosimetry parameters, such as g(r) and F(r,θ) as used in treatment-planning systems, comparisons of the dose-rate distributions of brachytherapy sources in not readily available. In practice, the dose-falloff is dominated by the solid-angle (c.f., the geometry function), which masks the subtle differences between sources. Additionally, there are a multitude of brachytherapy source models available having marked different construction. As a consequence, variations in brachytherapy dosimetry parameters between source models are expected to be larger than parameters obtained by different investigators for the same source. Furthermore, not all investigators used the same active length towards determining the geometry function or other dependent parameters. Therefore, we analyzed the ratios of dose-rate distributions obtained using theoretical and experimental techniques for a variety of brachytherapy sources.
Scope of Seeds Included in Analysis Specifically, the ratios of the theoretically calculated dose-rate distribution were divided by the experimentally determined dose-rate distribution on the transverse plane. A total of 19 seeds were included in this analysis, divided into 14 125I seeds and 5 103Pd seeds. The scope of seeds included in this comparison were as follows: 125
I Amersham model 6702 Amersham model 6711 Amersham model 6733 Best model 2301 Draximage model LS-1 IBt model 1251L Imagyn model IS-12501 IsoAid model IAI-125A Mentor model SL-125/SH-125 NASI model MED3631-A/M Nucletron model 130.002 SourceTech Medical model STM1251 Theragenics model 125.SO6
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103
Pd Best model 2335 Draximage model Pd-1 IBt model 1031L NASI model MED3633 Theragenics model 200
While some of these seeds are no longer available for clinical use, their data were still included in this analysis toward increasing the scope of this comparison to glean subtle differences between theoretical and experimental dosimetry techniques. Candidate datasets encompassing Monte Carlo and measured datasets were considered and compared in a manner similar to that described in the 2004 update to the AAPM TG-43 Report (Rivard et al., 2004). There were 38 candidate datasets for the 125I seeds with 25 ratios, and 14 candidate datasets for the 103Pd seeds with 10 ratios. Many of the references for these 52 datasets can be found in either appendix A or the references section in Rivard et al., (2004). In addition to sorting by radionuclide type (e.g., 125I or 103Pd), agreement between calculated and measured dose-rate distributions was assessed as a function of distance, r, and based on whether or not the radionuclide-backing material was either highor low-Z.
Ratios of Monte Carlo and Measured Dose-Rate Distributions for 125I The 25 ratios for 125I are shown in Figure 18a. Due to the related uncertainties of the data and the tight clustering on this linear-linear graph, no attempt as been made to discern particular ratios since the project focus is a general assessment of dose-rate ratios for all 125I seeds. The ratios typically fall within ±20% for all distances. From the observed trends, the following hypotheses may be made: (i) at close distances such as r = 0.5 cm, detector positioning uncertainties become pronounced for some datasets; (ii) differences between calculated and measured dose rates are minimized 1 ≤ r ≤ 2 cm, possibly due to maximally increasing the ratio of detector signal strength and positioning uncertainties; and (iii) differences generally increase in a monotonic fashion, possibly due to uncertainties in the phantom material correction factors for measurements or possibly due to uncertainties in Monte Carlo cross-section libraries which become more pronounced at larger distances. Of these 25 ratios, 13 ratios were for 8 125I sources having low-Z radionuclide-backing materials such as graphite or resin, and 12 ratios were for the remaining 6 125I sources having high-Z backing materials such as silver or tungsten. These dose-rate ratios are shown in Figures 18b and Figures 18c, respectively. While the calculated-to-measured dose-rate ratios for seeds having high-Z backing materials seeds are somewhat more clustered for 2 ≤ r ≤ 5 cm, the differences are not definitive and it is evident that correlation of the ratios is not significantly dependent on the radionuclide-backing material. As the lines are all stacked, the data in Figure 18a may be plotted in a different manner as in Figure 18d. Here, the frequency distributions for the 25 125I ratios are presented at r = 1 cm and r = 5 cm with 4% intervals (e.g., 100% ± 2%). As expected from Figure 18a, it is now clearly apparent that the r = 1 cm data has a tighter distribution than the r = 5 cm data.
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Figure 18a. Twenty-five ratios of calculated-tomeasured dose-rate distributions on the transverse plane for 14 125I seeds and 38 candidate datasets.
Figure 18b. Ratios of calculated-to-measured dose-rate distributions on the transverse plane for 125 I seeds having low-Z backing materials.
Figure 18c. Ratios of calculated-to-measured dose-rate distributions on the transverse plane for 125 I seeds having high-Z backing materials.
Figure 18d. Frequency distribution of the number of calculated-to-measured 125I dose-rate distribution ratios at r = 1 cm and r = 5 cm.
Ratios of Monte Carlo and Measured Dose-Rate Distributions for 103Pd Figure 19a depicts the 10 ratios for 103Pd. Just as for 125I, the best agreement between calculated and measured results was at r = 1 cm. However for 103Pd, the disparity as a function of distance was even more pronounced than for 125I. None of the 103Pd sources had high-Z backing materials, therefore, a comparative analysis as performed in Figures 18b and 18c was not possible. However, the frequency distribution presented in Figure 18d may be compared to that of Figure 19b. Again, the same trend of improved correlation at smaller distances was observed.
Conclusions In the past decade, quantitative dosimetry methods have continued to improve and become more widely used as the basis for clinical treatment planning. Low-energy photon-emitting brachytherapy planning now depends almost exclusively on Monte Carlo-based or measured single-source dose distributions. It appears that quantitative dosimetry techniques are rapidly overtaking more traditional model-based dose
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Figure 19a. Ten ratios of calculated-to-measured dose-rate distributions on the transverse plane for five 103Pd seeds and 14 candidate datasets.
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Figure 19b. Frequency distribution of the number of calculated-to-measured 103Pd dose-rate distribution ratios at r = 1 cm and r = 5 cm.
calculation in the higher energy domain as well. While no major breakthroughs have occurred in the last decade, the uncertainty of Monte Carlo and TLD dose-estimation has become better established: 3% to 5% for Monte Carlo and 7% to 10% for TLD. As quantitative dosimetry techniques have become more widely disseminated and applied to more source models, broader intercomparisons of theoretical and measured dose distributions are possible, further increasing confidence in the basic dosimetric infrastructure of brachytherapy. In the next decade, we can anticipate the following improvement: (a) Improved dose measurement techniques with uncertainties near the theoretical limit of 3% to 4%; and (b) development of more accurate dose-calculation engines based upon Monte Carlo simulation and other transport equation solvers for patient-specific treatment planning.
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Chapter 16
The TG-43 Brachytherapy Dose Calculation Formalism Mark J. Rivard Department of Radiation Oncology, Tufts University School of Medicine Boston, Massachusetts Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 296 Clinical Rationale for Accurate Dosimetry . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 297 Task Group # 43 Dosimetry Formalism . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 298 General 2-D Formalism . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 299 Air-Kerma Strength . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 300 Dose-Rate Constant . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 301 Geometry Function . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 301 Radial Dose Function . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 302 2-D Anisotropy Function . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 303 General 1-D Formalism . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 303 Consensus Datasets for Clinical Implementation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 304 Source Geometry Variations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 306 General Discussion of TG-43 Dosimetry Parameters . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 306 Air-Kerma Strength . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 306 Dose-Rate Constant . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 307 Radial Dose Function . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 307 2-D Anisotrophy Function . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 307 Uncertainty Analysis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 308 Recommended Dosimetry Parameter Acquisition Methodology . . . . . . . . . . . . . . . . . . . . . . . . 308 General Recommendations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 309 Preparation of Dosimetry Parameters . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 309 Air-Kerma Strength Standard . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 309 Dose-Rate Constant . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 309 Radial Dose Function . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 310 2-D Anisotrophy Function . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 311 1-D Anisotrophy Function . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 311 Reference Data and Conditions for Brachytherapy Dosimetry . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 311 Radionuclide Data . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 311 Reference Media . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 311 Methodological Recommendations for Experimental Dosimetry . . . . . . . . . . . . . . . . . . . . . . . . . . 312 Detector Choice . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 312 Medium and Energy Response Characterization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 313 Specification of Measurement Methodology . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 314 Methodological Recommendations for Monte Carlo-based Dosimetry . . . . . . . . . . . . . . . . . . . . . . 314 Specification of the Monte Carlo Calculation Methodology . . . . . . . . . . . . . . . . . . . . . . . . . . . . 315 Good Practice for Monte Carlo Calculations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 316 Publication of Dosimetry Results . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 316 Clinical Implementation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 317 Dose-Calculation Formalism . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 318 Acceptance Testing and Commissioning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 319 Source Calibrations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 319 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 321
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Introduction In 2004, the American Association of Physicists in Medicine (AAPM) published an update (TG-43U1) by Rivard et al., (Rivard et al., 2004a,b) to the Task Group No. 43 brachytherapy dosimetry protocol (TG43) by Nath et al., (1995). In this joint AAPM/American Brachytherapy Society (ABS) Summer School text on brachytherapy physics, a synopsis of this updated protocol (AAPM TG-43U1) is included for completeness and internal reference. While the reader is referred to the original report, this chapter should not be construed as AAPM guidance or recommendations since some of the descriptions and sections are somewhat modified from the published AAPM report/protocol. Specifically, this chapter focuses on the formalism and background of the protocol, and leaves the reader to reference the 2004 update (TG-43U1) for seed-specific data. The original TG-43 report introduced a new brachytherapy dose calculation formalism in 1995 that was largely based on findings of the Interstitial Collaborative Working Group (ICWG, 1990). Previous calculation formalisms were based upon apparent activity, equivalent mass of radium, exposure-rate constants, and tissue-attenuation coefficients. These older formalisms did not account for source-to-source differences in encapsulation or internal construction. With the exception of radium, the exposure-rate constants and other input parameters to these algorithms depended only on the radionuclide (Williamson 1997). In contrast, TG-43 employed dose-rate constants and other dosimetric parameters that depended on the specific source design, and are directly measured or calculated for each source design. Additionally, TG-43 presented consensus dosimetry data, in terms of the recommended formalism, for the three low-energy photon emitting source models then available (Theragenics Corporation model 200 103Pd source and Amersham Health models 6702 and 6711 125I sources). These data were based upon a critical review of ICWG measured dose-rate distributions using LiF TLD as well as other measurements and Monte Carlo calculations available in the literature. Overall, the TG-43 protocol has resulted in significant improvements in the standardization of both dose-calculation methodologies as well as dose-rate distributions used for clinical implementation of brachytherapy. For example, the differences between the previously used dose-rate distributions and those recommended by TG-43 were as large as 17% for some sources. These changes have been exhaustively reviewed by the physics community and are generally accepted. Many treatment planning software vendors have implemented the TG-43 formalism and the recommended dosimetry parameters in their systems. LiF TLD dose measurements and Monte Carlo dose calculations have largely replaced the semi-empirical dose-calculation models of the past. Since the publication of the TG-43 protocol a decade ago, significant advances have taken place in the field of permanent source implantation and brachytherapy dosimetry, necessitating an update of this protocol for the following reasons: (a) To eliminate minor inconsistencies and omissions in the original TG-43 formalism and its implementation (Meigooni and Nath 2000; Kline 1996; Fung 1998). (b) To incorporate subsequent AAPM recommendations, addressing requirements for acquisition of dosimetry data as well as clinical implementation (Williamson et al., 2000). These recommendations, e.g., elimination of Aapp and description of minimum standards for dosimetric characterization of low-energy brachytherapy sources (Williamson et al., 1998, 1999b) needed to be consolidated in one convenient document to complement the joint AAPM/RPC Brachytherapy Source Registry (RPC 2005), (c) To critically reassess published brachytherapy dosimetry data for the 125I and 103Pd source models introduced both prior and subsequent to publication of the TG-43 protocol in 1995, and to recommend consensus datasets where appropriate.
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(d) To develop guidelines for the determination of reference-quality dose distributions by both experimental and Monte Carlo methods and to promote consistency in derivation of parameters used in TG-43 formalism. Approximately 20 new low-energy interstitial brachytherapy source models have been introduced to the market since publication of TG-43 in 1995. These commercial developments can be attributed to the rising popularity of permanent prostate brachytherapy. Some of these new brachytherapy sources were introduced into clinical practice without thorough scientific evaluation of the necessary dosimetric parameters. The AAPM addressed this issue in 1998, recommending that at least one experimental and one Monte Carlo determination of the TG-43 dosimetry parameters be published in the peer-reviewed literature before using new low-energy photon-emitting sources (those with average photon energies less than 50 keV) in routine clinical practice (Williamson et al., 1998). Thus, many source models are supported by multiple dosimetry datasets based upon a variety of basic dosimetry techniques. This confronts the clinical physicist with the problem of critically evaluating and selecting an appropriate dataset for clinical use. To address this problem, this AAPM protocol presented a critical review of dosimetry data for eight 125I and 103Pd source models which satisfied the aforementioned criteria as of July 15, 2001, including the three low-energy source models included in the original TG-43 protocol. A further update of this protocol is anticipated for 2006 in Medical Physics to provide consensus, single source dose distributions and dosimetry parameters for low-energy photon sources. The present protocol (TG-43U1) recommends a single, consensus dataset for each source model from which the full 2-D dose-rate distribution can be reconstructed. Beta- or neutron emitting sources such as 90Sr, 32P, or 252Cf are not considered in this protocol. Finally, methodological guidelines are presented for physicist-investigators aiming to obtain dosimetry parameters for brachytherapy sources using calculative methods or experimental techniques.
Clinical Rationale for Accurate Dosimetry While low-energy, photon-emitting brachytherapy sources have been used to treat cancers involving a variety of anatomical sites, including eye plaque therapy for choroidal melanoma and permanent lung implants (Ling et al., 1983; Lee et al., 2003), their most frequent indication today is for the treatment of organconfined prostate cancer (Yu et al., 1999). Prostate cancer is the most prevalent cancer in men in the United States with approximately 220,000 new cases incident per year and an annual death rate of about 29,000 (Mettlin et al., 1998). While approximately 60% of new cases are confined to the organ at time of diagnosis, only about 2.2% of these new cases were treated with brachytherapy in 1995. Since that time, the percentage has increased to about 30% of all eligible patients receiving implants in current practice. This increase was largely due to improvements in diagnosis and case selection facilitated by introduction of the prostate specific antigen (PSA) screening test, and to improved ultrasound-guided delivery techniques. In the United States, the pioneering work was performed by a group of investigators located in New York (Zelefsky and Whitmore 1997) and Seattle (Ragde et al., 1998) based on work conducted in Denmark (Holm et al., 1983). The ABS critically reviewed the current published data and published consensus guidelines in 2001 to direct appropriate selection of patients. Prostate brachytherapy is appropriate as monotherapy for men with low risk favorable disease (T1-2a, Gleason score ≤6 and PSA < 10). While a prospective, randomized trial will likely never be performed due to strong patient preferences, long term survival rates published thus far appear comparable between prostate brachytherapy and radical prostatectomy for men with organ-confined disease. The most widely used technique utilizes transrectal ultrasound (TRUS) guided implantation of either 125I or 103Pd brachytherapy sources using a templateguided needle delivery system to avoid open surgery required by the retropubic approach (Whitmore, Hilaris, and Grabstald 1972; Peschel, King, and Roberts 1998). Due to its higher dose rate, 103Pd is often prescribed for higher grade tumors with 125I reserved for their low grade counterparts, despite a lack of clinical evidence supporting this practice.
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The ABS recommends that CT-based post-implant dosimetry be performed on every patient undergoing a permanent prostate seed implant. Multiple studies have shown that clinical outcomes in prostate brachytherapy, both for the retropubic approach and the TRUS-guided technique, correlate with dose coverage parameters. Similarly, there is evidence that toxicity correlates with urethral and rectal doses. The extensive clinical experience of Memorial Sloan-Kettering Institute (1078 patients with retropubic approach surgery) from 1970–1987 was reviewed by Zelefsky and Whitmore (Zelefsky and Whitmore 1997). Multivariate-analysis revealed a D90 implant dose of 140 Gy to be an independent predictor of recurrence-free local control at 5, 10, and 15 years (p=0.001). D90 is defined as the dose delivered to 90% of the prostate volume as outlined using post-implant CT images. Similarly, a review of 110 implants at Yale University by Nath et al. (1998) using the retropubic implant approach from 1976 to 1986 reported a correlation (p=0.02) of recurrence-free local control after 10 years with V100; V100 is defined as the percentage of target volume receiving 100% of the prescribed dose. Two recent retrospective studies of the TRUS technique demonstrate that the clinical outcome depends on dose delivered and prostate volume coverage. Stock et al. (1998) reported on an experience of 134 prostate cancer patients implanted with 125I and not treated with teletherapy or hormonal therapy. They assessed rates of freedom from biochemical failure as a function of the D90 dose. A significant increase in freedom from biochemical failure (92% vs. 68% after 4 years) was observed (p=0.02) for patients (n=69) where D90 > 140 Gy. Potters et al. (2001) recently reviewed the impact of various dosimetry parameters on biochemical control for their experience of 719 patients treated with permanent prostate brachytherapy. Many of these patients also received teletherapy (28%) or hormone therapy (35%). Furthermore, 84% of the implants used 103Pd with the remainder using 125I. Their results indicated that patient age, radionuclide selection, and use of teletherapy did not significantly affect biochemical relapse-free survival (PSA-RFS). The only dose-specification index that was predictive of PSA-RFS was D90. Like the other two studies, studies by Stock et al., and Potters et al., were based on pre-TG-43 prescription doses of 160 Gy, and both indicated a steep dependence of clinical outcome with dose in the range of 100 to 160 Gy. For example, Stock et al., reported freedom from biochemical failure rates of 53%, 82%, 80%, 95%, and 89% for patients receiving D90 ± arcsin(L/2r) for right-cylinder sources coated with low-energy photon emitters due to screening of photons by the active element at angles towards the transverse-plane. Active length, L, used to evaluate GL(r,θ) in equation (4) shall be the same L used to extract gL(r) and F(r,θ) from dose distributions via equations (6) and (8), respectively. Otherwise, significant errors in dosimetry results at small distances may arise. For example, at r = 0.5 cm, a change in L from 3 to 5 mm results in a 5% change in GL(r,θ0).
General 1-D Formalism While a 1-D isotropic point-source calculation [equation (9)] only provides an approximation the true 2D dose distribution, it simplifies source localization procedures by eliminating the need to determine the orientation of the source longitudinal-axis from imaging studies. D" (r)= SK ⋅ Λ ⋅
G X (r,θ0 )
⋅ g (r)⋅ φan (r) G X (r0 ,θ0 ) X
(9)
Users should adopt one of the following implementations of equation (9): 2
r " r = S ⋅ Λ ⋅ 0 ⋅ g r ⋅ φ (r) , D K P an r
()
()
(10)
or 2
r " ( r) = S ⋅ Λ ⋅ 0 ⋅ g ( r) ⋅ φ (r) . D K P an r
(11)
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While commercial treatment planning systems generally utilize the formalism in equation (10), that presented in equation (11) is the recommended formalism due to its improved accuracy at small distances, e.g., less than r = 1 cm. These formulations require consistency between the geometry function used for dose calculation and the geometry function used for extracting gX(r) from the transverse-plane dose distribution. Furthermore, these revised formulations correct an inconsistency in equation (11) of the original TG-43 protocol that indirectly recommended the following incorrect equation: " = S Λ ⋅ GP (r, θ 0 ) ⋅ g (r) ⋅ φ (r ) . NOT RECOMMENDED D(r) K L an GP (r0 , θ 0 )
(12)
While use of the wrong gX(r) datasets will typically give errors in the calculated dose rate of less than 2% at distances beyond 1 cm, average errors of 3%, 15%, and 74% arise at distances of 0.5, 0.25, and 0.1 cm, respectively. Clinical utilization of the 1-D dosimetry formalism presented in equation (12), or other formalisms that inconsistently apply the geometry function, are not recommended. The 1-D anisotropy function, φan(r), is identical to the anisotropy factor defined by the original TG43 protocol. At a given radial distance, φan(r) is the ratio of the solid angle-weighted dose rate, averaged over the entire 4 π steradian space, to the dose rate at the same distance r on the transverse-plane, see equation (13). π
φan(r)=
∫ D" ( r,θ ) sin (θ ) dθ 0
(13)
2 D" ( r, θ 0 )
Note that one should integrate dose rate, not the values of the 2-D anisotropy function to arrive at φan(r). Equation (11) will exactly reproduce the solid-angle weighted dose rate at a given r and is recommended because the line-source geometry function will provide more accurate interpolation and extrapolation at small distances. The accuracy achievable using the 1-D formalism for prostate implants has been reported (Lindsay, Battista, and Van Dyk 2001; Corbett et al., 2001). For brachytherapy treatment planning systems that do not permit entry of φan(r), Equations (10) or (11) can still be implemented by carefully modifying gX(r) to include φan(r) as shown in equation (14). Linear interpolation may be used to match the gX(r) and φan(r) grid spacing. These modified dosimetry parame– ters, g´(r) and φ 'an, are defined as: g ' ( r ) = g X ( r ) ⋅ φan ( r ) φ 'an = 1
(14)
Consensus Datasets for Clinical Implementation In the TG-43U1 protocol and subsequent updates, a deadline was established for inclusion of sources which had satisfied the AAPM recommendations that comprehensive (2-D) reference-quality dose-rate distribution data be accepted for publication by a peer-reviewed scientific journal. Appropriate publications can report either Monte Carlo, or experimentally derived TG-43 dosimetry parameters. For each source model, a single consensus dataset was derived from multiple published datasets according to the
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following methodology (Nath et al., 2002). If items essential to critical evaluation were omitted, the authors were contacted for information or clarification. (a) Experimentally determined values for the dose-rate constant (EXPΛ) were averaged. Separately, Λ values obtained using Monte Carlo techniques (MCΛ) were averaged. The CONΛ value recommended in this protocol (CONΛ) is the equally weighted average of the separately averaged experimental and Monte Carlo L values. In cases where there is only 1 experimental result and 1 Monte Carlo result: CONΛ = [EXPΛ + MCΛ]/2. (b) Each candidate dataset was examined separately and eliminated from consideration if it was determined to have a problem, e.g., data inconsistency. Corrections for use of a non-liquid water measurement phantom were applied if not included in the original investigators’ analysis. (c) For the 2-D anisotropy function, F(r,θ), and the radial dose function, g(r), all candidate datasets for a given source model were transformed using identical line-source geometry functions to permit fair comparison. The radial dose function was corrected for non-liquid water measurement medium if necessary. Assuming that the different datasets agreed within experimental uncertainties, the consensus data were defined as the ideal candidate dataset having the highest resolution, covering the largest distance range, and having the highest degree of smoothness. For most source models examined in this protocol, the consensus F(r,θ) and g(r) data, CONF(r,θ) and CONg(r), were taken from the transformed Monte Carlo dataset. (d) A few entries in the tabulated consensus datasets were taken from the non-ideal candidate dataset(s) to cover a larger range of distances and angles. These data were italicized to indicate that they were not directly confirmed by other measurements or calculations. (e) The 1-D anisotropy function, φan(r), was derived using numerical integration of the dose rate, as – calculated from CONF(r,θ) dataset, with respect to solid angle. Use of the anisotropy constant, φ 'an, is discouraged. (f) When scientifically justified for a given source model, exceptions or modifications to these rules were made, and are described later. For example, if the datasets were too noisy, they were rejected. (g) Following tabulation of g(r) and F(r,θ) for all eight source models, overly dense datasets were down-sampled to permit reasonable comparisons. Removal of a dataset point was deemed reasonable if linear interpolation using adjacent points resulted in a difference no larger than ±2% of the dataset point in question. Similarly, because the various authors used different table grids, it was necessary to interpolate some of the data into the common mesh selected for presenting all eight datasets. Linear-linear interpolation was used for F(r,θ) datasets, and log-linear interpolation was used for g(r) datasets. In this case, interpolated data are indicated by boldface. The details used to evaluate dosimetry parameters for each source were: 1. Internal source geometry and a description of the source, 2. Review of the pertinent literature for the source, 3. Correction coefficients for 1999 anomaly in NIST air-kerma strength measurements (if applicable), 4. Solid water-to-liquid water corrections, 5. Experimental method used: TLD or diode, 6. Active length assumed for the geometry function line-source approximation,
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10. Agreement between Monte Carlo calculations and experimental measurement.
Source Geometry Variations Source geometry and internal construction are highly manufacturer-specific and can affect the dosimetric characteristics of the source. Source models vary with regard to weld thickness and type, radioactivity carrier construction, presence of radio-opaque material with sharp or rounded edges, the presence of silver (which produces characteristic x-rays that modify the photon spectrum), and capsule wall thickness. Radioactive carriers may consist of a radio-transparent matrix, a radio-opaque object coated with radioactivity, or a radiotransparent matrix with highly attenuating radioactive coating. For example, the Amersham model 6702 and NASI model 3631-A/M sources utilize spherical resin carriers coated or impregnated with radioactivity. The number of spheres varies from 3 or more per source. Other sources, such as the Amersham model 6711, utilize a silver rod carrier. The amount of silver, or the length of silver rod, varies by the source model. Graphite pellets are also used. For example, in the Theragenics Corporation model 200 103 Pd source, the pellets are coated with a mixture of radioactive and nonradioactive palladium. All 125I and 103Pd source models, except for the now-obsolete model 6702 source, contain some type of radio-opaque marker to facilitate radiographic localization. For example, the graphite pellets of the Theragenics Corporation source are separated by a cylindrical lead marker. Beside the obvious dependence of photon spectrum on the radioisotope used, the backing material (e.g., the radiopaque marker) may further perturb the spectrum. For the sources containing 125I deposited on silver, the resultant silver x-rays significantly modify the effective photon spectrum. These source construction features influence the resultant dose rate distribution and the TG-43 dosimetry parameters to varying degrees. Accurate knowledge of internal source geometry and construction details is especially important for Monte Carlo modeling. While the dosimetry formalism was presented, its applicability to the derivation of consensus datasets is given later.
General Discussion of TG-43 Dosimetry Parameters Air-Kerma Strength Standards The NIST Wide-Angle Free-Air Chamber or WAFAC-based primary standard became available in 1998, and was used to standardize the 125I sources then available. The WAFAC standard shifted for unknown reasons in 1999, and was corrected in the first half of 2000. In early 2000, NIST noticed a shift in wellchamber calibration coefficients, with an apparent downward shift in the WAFAC air-kerma strength of approximately 3% for 125I sources, 5% for 125I-on-Ag sources, and 5% for 103Pd sources. NIST conducted a thorough investigation of this anomaly, completely checking all systems, mechanical, electrical, environmental, software, etc., particularly any factor that could equally affect both WAFAC instruments. However, a convincing explanation for this temporary shift was not determined. Thus, seed calibrations performed at NIST after mid 1998 and before January 1, 2000, need correction due to this measurement anomaly present in 1999. These dates are subject to revision should changes in manufacturing procedures, source geometry, or the WAFAC standard itself occur that affect the accuracy of vendor or ADCL secondary standards. Future source model-specific revisions to the calibration standard could require corresponding corrections to the recommended dose-rate constant. For this reason, regular calibration comparisons among NIST, ADCL, and vendors are required.
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In summary, there were two possible situations regarding the calibration of seeds at NIST using the WAFAC-based air-kerma strength standard. First, seed calibrations performed at NIST during the 1999 calendar year need correction due to a measurement anomaly present in 1999 only. This correction was determined by another WAFAC measurement for each seed model made at a designated date in 2000 or later. Secondly, WAFAC measurements made prior to 1999 and after January 1, 2000, needed no correction (Williamson et al., 1999a). Since the notation SK,N99 represents the NIST WAFAC-based air-kerma strength standard as officially introduced on January 1, 1999, this notation is used for all WAFAC measurements, regardless of the date of calibration. Thus, all measured dose-rate constant data given in this protocol have been normalized to the SK,N99 standard. Any measured dose-rate constants normalized to NIST calibrations performed in 1999 have been appropriately corrected for WAFAC measurement anomalies. Dose-Rate Constant Specifying the dose-rate constant as accurately as possible is essential, as it is used to transform the relative TG-43 dose distribution into absolute dose rates given the air-kerma strength of the sources deployed in the patient. As discussed in more detail later, Monte Carlo simulations have a freedom from detector positioning and response artifacts, smaller estimated uncertainty, and can yield artifact-free dose-rate estimates at distances shorter or longer than those accessible by TLD measurement techniques. However, the accuracy of Monte Carlo is inherently limited by the investigators’ ability to accurately delineate the source internal geometry. Few Monte Carlo studies have systematically evaluated the effects of geometric uncertainty, internal component mobility, tolerances in the fabrication of sources, and small manufacturing changes on the uncertainty of calculated dose-rate distributions. Therefore, the use of Monte Carlo values without confirmation by experimental studies is highly undesirable. Drawbacks of TLD dosimetry include (a) limited precision of repeated readings and spatial resolution; (b) a large and somewhat uncertain relative energy response correction; (c) failure of most investigators to monitor or control the composition of the measurement medium (Williamson 1991). For these reasons, the low-energy interstitial brachytherapy dosimetry subcommittee (LIBD) recommends using an equally-weighted average of the average measured (e.g., using TLDs) and average calculated (e.g., Monte Carlo-derived) values since the two recommended dosimetry characterization techniques have complementary strengths and limitations. Radial Dose Function For each source, Monte Carlo values of g(r) were graphically compared with experimental values. A comparison of the Monte Carlo and experimental g(r) results were expected to show an average agreement of ±10%. While the observed differences were typically < 5% for r ≤ 5 cm, systematic differences as large as 10% were observed due to use of outdated Monte Carlo cross-section libraries. Experimental values are difficult to measure at r < 1 cm, but Monte Carlo calculation of dose rate values are often available at smaller distances. In each case, the most complete dataset (typically Monte Carlo values) was used since values were readily available over a larger range of distances (especially at clinically significant distances closer than 1 cm) than provided by experimental measurements. 2-D Anisotropy Function Because Monte Carlo based datasets generally have superior smoothness, spatial and angular resolution, and distance range, all anisotropy functions recommended in this protocol are derived from Monte Carlo results which have been validated by comparison to less complete experimental datasets. A graphical comparison of datasets was performed, and the agreement between the Monte Carlo datasets and the experimental datasets was again expected to be ±10%. For θ > 30°, observed differences between the datasets were typically < 5% with a maximum of about 9%. For θ ≤30°, differences were larger (typi-
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cally ~10% with maximum ~17%), and are attributed to volume averaging and the high-dose-rate gradient near the source longitudinal-axis as well as uncertainties in the source geometry assumed by Monte Carlo simulations.
Uncertainty Analysis Most of the experimental and computational investigations, especially those published prior to 1999, failed to include a rigorous uncertainty analysis, and the AAPM recommends that henceforth dosimetry investigators include rigorous uncertainty analyses, specific to their methodology employed, in their published articles (i.e., Table 1). Based on results of Monroe and Williamson (Monroe and Williamson 2002), purely Monte Carlo estimates of the transverse-axis dose-rate per unit air-kerma strength typically have uncertainties of 2% to 3% at 1 cm and 3% to 5% at 5 cm, depending on the type and magnitude of internal seed geometric uncertainties. Since relatively little has been published on estimation of systematic (type B) uncertainties of Monte Carlo-based dose estimation, the following sections apply the principles of uncertainty analysis, as outlined in NIST Technical Note 1297 (Taylor and Kuyatt 1994), to estimation of total uncertainty of Monte Carlo dose-rate constants, MCΛ, Monte Carlo radial dose functions MCg(r), consensus dose-rate constants, CONΛ, and absolute transverse-axis dose as evaluated by the dosimetric parameters recommended in the report. Implementation of a practical uncertainty analysis (Rivard, Melhus, and Kirk 2004) may be considered.
Recommended Dosimetry Parameter Acquisition Methodology The AAPM recommends a list of methodological details that should be described in brachytherapy dosimetry publications based upon either experimental or theoretical methods, along with more prescriptive
Table 1. Generic Uncertainty Assessment for Experimental Measurements Using TLDs, and Monte Carlo Methods for Radiation Transport Calculations. Type A and B uncertainties correspond to statistical and systematic uncertainties, respectively. All values provided are for 1 σ. TLD uncertainties Component
Type
Repetitive measurements
4.5%
TLD dose calibration
A Type B
Monte Carlo uncertainties Component Statistics
2.0%
(including linac calibration)
r = 1 cm
r = 5 cm
Photo-ionization
0.3% · d (1,θ0) =
1.0% · d (5,θ0) =
cross-sections
1.5%
4.5% 2% · d (5,θ0) =
(2.3%) LiF energy correction
5.0%
Seed geometry
Measurement medium
3.0%
Source energy
2% · d (1,θ0) =
spectrum
0.1%
0.3%
Quadrature sum
2.5%
5.0%
Correction factor Seed/TLD positioning
4.0%
Quadrature sum
4.5%
Total uncertainty
8.6%
ADCL SK uncertainty
1.5%
Total combined uncertainty in Λ
8.7%
7.3%
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guidelines on performing such studies. To better appreciate results from a particular dosimetric measurement and its uncertainties, the reader is referred to a listing of parameters needed to assess data for TLD measurements (Kron et al., 1999). Unfortunately, this level of description was not realized in many of the papers cited. When key data or methodological details were missing from a published paper, the author was contacted and asked to provide the missing information.
General Recommendations Since publication of TG-43 (Nath et al., 1995), the AAPM has published guidelines on dosimetric prerequisites (Williamson et al., 1998) for low-energy photon-emitting interstitial brachytherapy sources. However, they provided few technical recommendations to investigators publishing reference-quality doserate distributions derived from measurements or radiation transport calculations. Based on the AAPM experience (Nath et al., 2002) of analyzing publications and preparing consensus dosimetry datasets, more detailed recommendations on dosimetry methodology and data analysis were presented to define minimum requirements for future source dosimetry studies.
Preparation of Dosimetry Parameters Dosimetric parameters should be tabulated for both 1-D and 2-D dose-calculation models. This will require the investigator to calculate the geometry function and the radial dose function using both point-source (1-D) and line-source (2-D) geometry functions. Consequently, the investigator should always specify the active length used for the 2-D line-source geometry function. It is necessary to recommend minimum spatial resolutions and ranges for which these parameters should be defined since specification at a few distances or angles will not allow a sufficiently complete or accurate dose reconstruction of the 2-D dose distribution. The underlying dose distribution may have high gradients and result in inaccurate interpolation by brachytherapy treatment-planning systems, unnecessarily giving rise to dose-delivery errors. Air-Kerma Strength For experimental measurement of absolute dose rates to water, at least one source should have direct traceability of SK to the 1999 NIST WAFAC calibration standard. Other sources should have a precisely transferred air-kerma strength using high-precision transfer devices such as well-characterized well-ionization chambers and secondary standards maintained by the investigator as well as the manufacturer’s laboratories. The investigator using experimental techniques should state the NIST SK calibration uncertainty in the evaluation of Λ. Dose-Rate Constant The experimental investigator should rigorously control and try to minimize all detector response artifacts such as: dose-rate dependence, dose response nonlinearity, energy dependence, volumetric averaging, temporal stability of readings and calibration coefficients, and accuracy of detector positioning both in the source measurement setup and the detector calibration setup. These issues should be discussed in the measurement methodology section of the published paper, and a rigorous uncertainty analysis should also be provided. · Experimentally, Λ is evaluated by taking the ratio of the absolute dose rate, D (r0 ,θ 0) (the only absolute dose rate required to define TG-43 dosimetry parameters) and the measured air-kerma strength of the source, decayed to the time of dose-rate measurement. Typically 8 to 10 sources are used, with at least · one source having direct traceability to a NIST calibration. At least 15 measurements of D (r0 ,θ 0) are · generally performed. For example, multiple measurements of D (r0 ,θ 0) around a single NIST WAFACcalibrated source could be made by placing multiple TLDs in different quadrants of the transverse plane.
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Monte Carlo radiation transport codes commonly provide direct estimates of absorbed dose or collision kerma per number of histories simulated (or some other internal normalization quantity, e.g., number of disintegrations, proportional to the number of primary photons simulated). Two simulations are necessary: one with the source model embedded in a phantom, yielding estimates of dose at specified points, and a second simulation with the source model positioned within a vacuum or a large air sphere. The doserate constant can then be estimated using the following ratio [equation (15)] since the numerator and denominator are similarly normalized, and the normalization constant used by the Monte Carlo code is irrelevant (Williamson 1988). Λ=
d" (r0 , θ 0 )
(15)
sK
· The parameter d (r0 ,θ 0) is the dose rate per history estimated using Monte Carlo methods at the reference position, and sK is the air-kerma strength per history estimated using Monte Carlo methods. Note the lower-case notation used to differentiate the normalized parameter, e.g., dose rate per history [cGy·h–1·history–1] as compared to absolute dose rate [cGy·h–1]. Although Monte Carlo studies are potentially free from experimental artifacts such as positioning uncertainties, energy response corrections, and signal-to-noise ratio limitations, such simulations require an accurate and complete geometric model of the source, selection of an appropriate cross-section library, and careful selection of dose tallying (estimation) and variance-reduction strategies. As with experimental studies, Monte Carlo-based dosimetry studies should include a complete uncertainty analysis. Care must be taken in evaluating dose rates at distances falling outside the range of tabulated data, rmin and rmax, especially at r < 0.5 cm. While difficult to measure, modern Monte Carlo techniques can easily calculate dose rates at distances as small as 0.1 cm from the center of the source (Williamson and Li 1995; Rivard 2002). This protocol recommends that gX(r) data be extracted and tabulated from Monte Carloderived dose rates for r ≥ 0.1 cm if possible. Users are warned that when working at small distances it is essential to use the same geometry function and L for evaluating absorbed dose as when used in the conversion of measured or Monte Carlo data in the TG-43 parameter ratios. Often, FX(r,θ) and gX(r) will exhibit non-physical properties due to breakdown of the line-source GL(r,θ) very near the source. Radial Dose Function The minimum distance range over which transverse-plane dose-rate data shall be measured or calculated should be from 0.5 to 7 cm for 125I and from 0.5 to 5 cm for 103Pd. Special attention to accuracy and completeness should be given to the 0.5 cm to 1.5 cm distance range, which is the most critical for accurately calculating dose for typical prostate and other volume implants and for accurately relating absolute dose rates (via the dose-rate constant) to the relative dose distribution. However, accurate dose characterization at shorter distances is critical for some procedures (e.g., episcleral eye plaque therapy) and for estimating hotspots in all implants. Accurate dose distributions at larger distances also contribute to overall dose-calculation accuracy and are important for assessing dose to organs at risk. Thus dose rates should be estimated down to the smallest distance supported by the methodology used. Some investigators have reported g(r) data down to 1 mm or less. Monte Carlo simulation can easily estimate dose at sub-millimeter distances as well as distances of 10 to 15 cm. In addition, some investigators have reported TLD or diode measurements at distances less than 1 mm. Since a 5th order polynomial is frequently used for smoothing transverse-plane data, one should present a minimum of 6 data points to fully-constrain the fit. Because these fits tend to perform poorly at small distances where the dose rate is highest, care should be taken to assure the fit is in good agreement at these positions. Acceptable levels of agreement are outlined in the following section.
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2-D Anisotropy Function When reporting 2-D anisotropy function data, at a minimum, F(r,θ) should be tabulated at radial distances, r = {0.5, 1, 2, 3, 5, and 7 cm} for 125I and {0.5, 1, 2, 3, and 5 cm} for 103Pd, and from θ = {0° to 90° with 10° increments} for all sources that are symmetric about the transverse-plane. To minimize extrapolation errors, data should be determined over as wide a radial distance range as reasonably possible given the experimental method or calculation technique. To minimize interpolation errors, data should also be obtained such that bilinear interpolation between various F(r,θ) data points produce errors less than 2%. When measuring 2-D anisotropy function data with sources that are cylindrically symmetric and symmetric about the transverse-plane (four identical quadrants), it is recommended that investigators determine relative dose rates a minimum of three times at each position analyzed. For source designs that are asymmetric about the transverse-plane or exhibit internal component shifts that may result in asymmetric dose distributions, F(r,θ) should be similarly tabulated, except with θ = {0° to 180° with 10° increments}. Because of the increased sampling capabilities using Monte Carlo methods in comparison to experimental methods, investigators using Monte Carlo methods should consider calculating dose rates using much finer grids in high-gradient regions such as near the source ends (e.g., θ near 0° or 180°). Angular resolution of a few degrees near the ends may be needed, with 10° resolution elsewhere. For those source designs with anisotropic photon-fluence distributions near the transverse plane, measurements and calculations having higher angular resolution are required to ensure that experimentally determined anisotropy functions are accurately normalized and that air-kerma strength per simulated history for Monte Carlo simulations is accurately calculated. The anisotropy constant does not exactly reproduce either the measured or Monte Carlo dosimetry data for r < 1 cm. TG-43U1 provides a detailed explanation for implementation of dose-calculation errors at small distances, e.g., r ≤ 1 cm. 1-D Anisotropy Function To derive 1-D anisotropy function data, a solid-angle weighted-average of the relative dose rates, uncorrected by the geometry function, should be performed over all angles. When examining small radii where θ=0° or 180° would place the calculation point within the source, the weighting should exclude the capsule/source volume and include only the volume outside the encapsulation. This is easily calculated for radii, r, less than half the capsule length where r sin θ > rcap, where rcap is the radius of the capsule.
Reference Data and Conditions for Brachytherapy Dosimetry Radionuclide Data The currently recommended values for half-lives, abundances and energies of 125I and 103Pd are presented in Table 2 (NNDC 2005; Lund/LBNL 1999; Seltzer 1993, Seltzer and Hubbell 1995). These values should be used to interpret future experimental measurements and as source spectra in Monte Carlo calculations. The 125I spectrum should be described in terms of five different photon energies. The 103Pd emission spectrum should be described in terms of eight discrete photon emissions. Although the relative number of high-energy photons emitted by 103Pd is low, their contribution to dose at distances beyond 10 cm can be clinically relevant and should also be considered for shielding calculations and exposure-control procedures (Rivard, Waid, and Wierzbicki 1999). Reference Media Water continues to be the recommended medium for reference dosimetry of interstitial brachytherapy sources. For dosimetry calculations and measurements, it may be necessary to know the composition of various forms of water and air. Pure, degassed water is composed of two parts hydrogen atoms and one
312
Mark J. Rivard Table 2. Recommended Nuclear Data for 125I and 103Pd for Brachytherapy Dosimetry
125
I (half-life = 59.40 ± 0.01 days) photon energy (keV) photons per disintegration
103 Pd (half-life = 16.991 ± 0.019 days) photon energy (keV) Photons per disintegration
27.202
0.406
20.074
0.224
27.472
0.757
20.216
0.423
30.98
0.202
22.72
0.104
31.71
0.0439
23.18
0.0194
35.492
0.0668
39.75
0.00068
294.98
0.00003
357.5
0.00022
497.1
0.00004
Weighted mean energy = 28.37 keV
Total = 1.476
I Γ10 keV = 0.0355 µGy·m2·h-1·MBq-1
125
Weighted mean energy = 20.73 keV
Total = 0.7713
Pd Γ10 keV = 0.0361 µGy ·m2·h-1·MBq–1
103
part oxygen atoms, with a mass density of 0.998 g ·cm–3 at 22 °C. Reference conditions for dry air are taken as 22°C and 101.325 kPa (760 mm Hg) with a mass density of 0.001196 g · cm–3. The composition of air may change as a function of relative humidity, and one must account for this effect (ICRU 37, 1989; Seltzer 2001). The proportion by weight of water in air of 100% relative humidity varies only between 1% and 2%, for temperatures between 16 °C and 26 °C and pressures between 735 mm Hg and 780 mm Hg. The change in mass density of saturated air is no more than a 1% reduction with respect to that for dry air, over this range of temperatures and pressures. Thus, the mass density will be set at 0.00120 g·cm–3 for both dry and moist air. For Monte Carlo calculations, the recommended relative humidity is 40%, which corresponds to the relative humidity in an air-conditioned environment where measurements should be carried out (Table 3).
Methodological Recommendations for Experimental Dosimetry Compared to Monte Carlo theorists who may idealize reality by a theoretic construct, the experimental investigator should address the variability that represents the clinical environment. The experimental study should investigate a reasonably large sample of sources received from multiple shipments at different stages of the production stream from the manufacturer. Detector Choice LiF TLD remains the method of choice for the experimental determination of TG-43 dosimetry parameters for low-energy photon-emitting brachytherapy sources. While a variety of other experimental dosimeters such as diodes, diamond detectors, miniature ionization chambers, plastic scintillators, liquid ionization chambers, polymer gels, radiographic and radiochromic film, and chemical dosimeters have been used for brachytherapy dosimetry, their validity for obtaining brachytherapy dosimetry parameters has not yet been convincingly demonstrated. The reader is referred to the TG-43U1 report (Rivard et al., 2004a) for a more in-depth discussion with references justifying the current view.
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Table 3. Composition (percent mass) of air as a function of relative humidity at a pressure of 101.325 kPa. Relative humidity (%)
Hydrogen
Carbon
Nitrogen
Oxygen
Argon
0
0.0000
0.0124
75.5268
23.1781
1.2827
10
0.0181
0.0124
75.4048
23.2841
1.2806
40
0.0732
0.0123
75.0325
23.6077
1.2743
60
0.1101
0.0123
74.7837
23.8238
1.2701
100
0.1842
0.0122
74.2835
24.2585
1.2616
Use of the aforementioned experimental dosimeters is an area of future research of significant scientific value. For measuring brachytherapy dosimetry parameters, detectors should have the following properties: (a) Detectors should have a relatively small active volume such that averaging effects of high-gradient dose fields over this volume are negligible or are accurately accounted for by correction coefficients. (b) A well-characterized energy-response function such that differences between the calibration energy and experimentally measured energy are either negligible or may be quantitatively accounted for. (c) Sufficient precision and reproducibility to permit dose-rate estimation with 1s statistical (Type A) uncertainties ≤ 5%, and 1s systematic uncertainties ≤ 7%. For example, TLD statistical uncertainties may be improved through repeated measurement at a given location, and systematic uncertainties may be improved through measuring chip-specific calibration coefficients. Typical statistical and systematic uncertainties for 1 × 1 × 1 mm3 TLD-100 chips are 4% and 7%, respectively, with total combined uncertainties of 7% to 9% (Meigooni et al., 2001). Therefore, 1 × 1 × 1 mm3 TLD-100 chips are considered a valid detector to perform the aforementioned absolute and relative measurements. Medium and Energy Response Characterization The measurement medium should also be well characterized (Meigooni et al., 1994). While epoxy-based substitutes for water, such as Solid Water™ by Gammex-RMI or Virtual Water™ by MED-CAL Inc., have liquid-water conversion coefficients that are less than 5% for high-energy teletherapy beams, coefficients range from 5% to 15% for low-energy photon-emitting sources. Recently, the measured calcium concentration of Solid Water‘ was found to have deviated from the vendor’s specification by as much as 30% (Patel et al., 2001). Therefore, when Solid Water is used in experimental dosimetry, the atomic composition of the material used should be measured and correction coefficients based on the measured composition of Solid Water should be used. Although Solid Water is the most widely used material for TG-43 reference dosimetry, it has several shortcomings. In addition to concerns over the constancy of its composition, Solid Water and similar water substitutes require solid-to-liquid water conversion corrections ranging from 4% to 15% in the 1 to 5 cm range. Alternative materials including polystyrene, polymethylmethacrylate, or plastic water (model PW2030 by Computerized Imaging Reference Systems, Inc.) may be possible candidates for future low-energy photon-emitting brachytherapy dosimetry studies. The relative energy response correction, E(r), is the largest single source of type B (systematic) uncertainty for TLD and other secondary dosimeters used in brachytherapy dosimetry. It is defined as the ratio of TLD response per unit dose in water medium at position r in the brachytherapy source geometry, to its
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response per unit dose in the calibration geometry, usually a calibrated 60Co or 6 MV x-ray beam. In general, E(r) depends on source-to-detector distance, r, and may include corrections for volume averaging (influence of dose gradients in the TLD volume), detector self-absorption, medium displacement, and conversion from the measurement medium to liquid water (Williamson and Meigooni 1995). Most investigators treat E(r) as a distance-independent constant, although when it includes volume-averaging and Solid-to-liquid water corrections, as is often the case for Monte Carlo estimates, E(r) varies significantly with distance (Patel et al., 2001). A more in-depth discussion of these issues is given in the TG-43U1 report (Rivard et al., 2004a). In utilizing measured or Monte Carlo E(r) estimates published by others, the AAPM recommends that TLD experimentalists confirm that the associated measurement methodology matches their dosimetry technique with regard to TLD detector type and size, annealing and readout technique, and megavoltage beam calibration technique. The latter requires accounting for differences in calibration phantom material and dose-specification media used by the experimentalist and assumed by the selected E(r) estimate. Finally, further research is needed to resolve the discrepancy between published E(r) values, to identify the appropriate role for transport calculations in TLD dosimetry, and to reduce the large uncertainty associated with relative energy-response corrections. Specification of Measurement Methodology The experimental investigator should describe the important features of the measurement materials and methods to permit assessment of the results: 1. Description of the external and internal source geometry, 2. Brachytherapy source irradiation geometry, orientation, and irradiation timeline, 3. Radiation detector calibration technique (including protocol from which the technique is derived) and energy response function, E(r), 4. Radiation detector (dimensions, model #, and vendor) and readout system (e.g., electrometer unit model # and settings, or TLD readout unit model #, vendor, time-temperature profiles, and annealing program), 5. Measurement phantom (composition, mass density, dimensions, model #, and vendor), 6. Phantom dimensions and use of backscatter (at least 5 cm backscatter is recommended for 125I and 103 Pd dosimetry measurements), 7. Estimation of the impact of volume averaging on the results at all detector positions, 8. Number of repeated readings at each position, the number of different sources used, and the standard deviation of the repeated readings, 9. NIST SK value and uncertainty used for the measured source(s), and 10. Uncertainty analysis section assessing statistical and systematic uncertainties and their cumulative impact.
Methodological Recommendations for Monte Carlo-based Dosimetry Monte Carlo codes used to model photon transport for brachytherapy dose calculation should be able to support detailed 3-D modeling of source geometry and appropriate dose-estimation techniques. In addition, they should be based upon modern cross-section libraries and a sufficiently complete model of photon scattering, absorption, and secondary photon creation. Codes that have been widely used for interstitial
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brachytherapy dosimetry include EGS, MCNP, and Williamson’s PTRAN code. These codes have been widely benchmarked against experimental measurements or each other, so that their appropriate operating parameters and limitations can be considered to be well understood (Williamson et al., 1993). In general, the AAPM recommends Monte Carlo investigators utilize such well-benchmarked codes for brachytherapy dosimetry studies intended to produce reference-quality dose-rate distributions for clinical use. However, regardless of the transport code chosen and its pedigree, all investigators should assure themselves that they are able to reproduce previously published dose distributions for at least one widely used brachytherapy source model. This exercise should be repeated whenever new features of the code are explored, upon installing a new code version, or as part of orienting a new user. Other radiation transport codes, including Monte Carlo codes not previously used in brachytherapy dosimetry, should be more rigorously tested and documented in the peer-reviewed literature before proposing to use their results clinically. This is especially true for other types of transport equation solutions, including multigroup Monte Carlo, discrete ordinates methods, and integral transport solutions that have been proposed for brachytherapy dosimetry. Due to the short range of the secondary electrons produced by interactions from photons emitted by the radionuclides covered in this protocol, electron transport is not required and collision kerma closely approximates absorbed dose. Since the investigator performing Monte Carlo analysis can control many features of the transport calculations, it is imperative that the salient details be described in publications presenting Monte Carlo-derived brachytherapy dosimetry data. For instance, the collisional physics model should be described. The standard model used by experienced Monte Carlo users includes incoherent scattering corrected for electron binding by means of the incoherent scattering factor, coherent scattering derived by applying the atomic form factor to the Thompson cross-section, and explicit simulation of characteristic x-ray emission following photoelectric absorption in medium- and high-atomic number media. For sources containing Ag or Pd, it is imperative that, if characteristic x-ray production is not explicitly simulated, the primary source spectrum be appropriately augmented to include their presence. Specification of Monte Carlo Calculation Methodology A list of key features that should be specified by the investigator in the publication follows: 1. Radiation transport code name and version number, and describe major options, if any, invoked, 2. Cross-section library name and version number, and describe customizations performed if any, 3. Radiation spectrum of the source, 4. Manner in which dose-to-water and air-kerma strength are calculated: specify name of estimator or tally, whether or not transport was performed in air and how attenuation correction coefficients · were applied, and how suppression of contaminant x-ray production for Kδ (d) calculations was performed to be compliant with the NIST SK,N99 standard, 5. Source geometry, phantom geometry, and sampling space within the phantom, 6. Composition and mass density of the materials used in the brachytherapy source, 7. Composition and mass density of the phantom media, 8. Physical distribution of the radioisotope within the source (point-source modeling is unacceptable), and 9. Uncertainty analysis section assessing statistical and systematic uncertainties and their cumulative impact.
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Good Practice for Monte Carlo Calculations For calculating brachytherapy dosimetry parameters, the following requirements should be adhered to: 1. Primary dosimetry calculations should be performed in a 30-cm diameter liquid water phantom, but calculations in Solid Water may also be performed to supplement experimental results, e.g., calculation of E(r), performed in Solid Water or other solid water substitutes. Typical calculations will produce dosimetry results extending out to r ~ 10 cm, with at least 5 cm of backscatter material for 125I and 103Pd dosimetry calculations. 2. Enough histories should be calculated to ensure that dosimetry results have a 1 σ (k=1, 67% confidence index) ≤ 2% at r ≤ 5 cm, and that calculations for derivation of SK, have 1 σ ≤ 1 % at the point of interest. 3. Modern, post–1980 cross-section libraries should be used, preferably those equivalent to the current NIST XCOM database such as DLC-146 or EPDL97. Exclude or appropriately modify (DeMarco, Wallace, and Boedeker 2002) older cross-section libraries (Storm and Israel 1970). Note that EGS4, EGSnrc, and MCNP all currently require modification or replacement of their default photo-ionization cross-sections to meet this requirement. Furthermore, moist air bestdescribes experimental conditions in comparison to dry air, and mass-energy absorption coefficients for moist air are recommended to minimize systematic uncertainties. 4. Manufacturer-reported source dimensions and compositions of encapsulation and internal components should be verified through the use of physical measurements, transmission radiography, and autoradiography. Just as the TLD experimentalist should measure an appropriate sample of sources, the Monte Carlo investigator should quantify the geometric variations in a sample of similar size. 5. The impact of volume-averaging artifacts should be limited to ≤1% through the appropriate choice of estimators (tallies) and scoring voxels if used. · 6. Calculations of d (r0 ,θ 0) to derive F(r,θ) should include high-resolution sampling in high-gradient regions such as near the source ends or in regions where internal source shielding causes abrupt · changes in d (r0 ,θ 0) and subsequently F(r,θ). · 7. K(d) should be modeled as a function of polar angle for sK simulation and Λ derivation. Williamson has shown that for some sources, detectors with large angular sampling volumes (such as the NIST WAFAC) will have a significantly different response than point-kerma detectors positioned on the transverse-plane. When the radioactivity is dispersed within or on the surface of a high-density core with sharp corners and edges, it may be necessary to simulate, if only approximately, the WAFAC geometry (dimensions and composition) to permit investigators the opportunity to directly compare Monte Carlo calculations of Λ with NIST-based measurements of Λ. 8. Mechanical mobility of the internal source structures, which has the potential to significantly affect the dose distribution, should be considered by the Monte Carlo investigator in developing both the geometric model of the source and the uncertainty budget.
Publication of Dosimetry Results Previous AAPM recommendations (Williamson et al., 1998) stated that dosimetry results should be published preceding clinical implementation. However, the journal Medical Physics established a “seed policy” in 2001 that, in effect, limits printing of articles to Technical Notes unless they contain significant
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new science. In order to comply with this restriction imposed by the journal, the AAPM will accept technical notes with limited details as acceptable, provided the full details as listed above are available to the committee at the time of evaluation. This policy in no way prevents publication of the article in other journals, as other scientific journals of interest to medical physicists are appropriate venues for publication of these dosimetry parameters. In an AAPM report, Williamson et al. (1998) recommended that dosimetry results be published by independent investigators, but did not offer a strict definition of what this independence entails. The spirit of the initial recommendation was to prompt publication of multiple studies to assess all the TG-43 brachytherapy dosimetry parameters, e.g., Λ, g(r), F(r,θ), and φ(r). Through determining the consensus datasets for the brachytherapy sources evaluated in this protocol, a definition of “Independence Policy” was adopted. There are two aspects of this policy, and both shall be met for full compliance. The first meaning of “independent studies” is that they are performed, written, and published by investigators who are affiliated with institutions independent of the source vendor and who have no major conflicts-of-interest with that vendor. The second meaning of “independent studies” is that they are scientifically independent of one another. In the case of two measurement-based studies, this will usually mean that two different investigators have used their own methodologies for measuring Λ and sampling the relative dose distribution, as TLD dosimetry is highly technique and investigator dependent. In the case of an empirical study and a Monte Carlo study, if properly executed, they will yield scientifically independent estimates of the TG-43 parameters. Thus, so long as the two studies are successfully scrutinized by the peer-review process and satisfy the AAPM scientific requirements, the empirical and Monte Carlo investigator rosters can overlap or even be identical. It is permissible to publish the Monte Carlo and measured estimates in the same paper so long as the two datasets are independently tabulated. In this context, “Not independent” means that the one study is used to modify the outcomes and methods of the other to improve agreement between the two datasets in a manner that is not scientifically justified. When possible, the authors should cite previous publications where the measurement system or techniques were first described, and illustrate only the key features. It does not benefit either the reader or the journal in question to continually restate the definition of TG-43 parameters or their formalism. Simply citing this protocol or the original TG-43 publication will suffice.
Clinical Implementation Dose distributions in and around clinical interstitial implants are calculated using computerized radiotherapy treatment planning (RTP) systems. For sources with radio-opaque markers, the 3-D coordinates of the centers (or the two ends) of the markers in implanted sources are determined using multiple-view radiographs or CT scans. The dose-rate contributions from each source at the points of interest are calculated using a one-dimensional or two-dimensional dose-calculation algorithm. These contributions are then summed to determine the total dose rate. This procedure assumes that there are no source-to-source shielding effects, that all tissue is water equivalent, and that the scattering volume within the patient is equivalent to that used in the consensus datasets. The term equivalent in this context means at least 5 cm of waterequivalent material surrounds the point of calculation. Many RTP systems are available commercially and use a variety of methods to calculate clinical dose-rate distributions. Some of the RTP systems use the single-source dosimetry data in a tabular form as input, whereas others represent the data by means of a mathematical formula that requires input of certain coefficients. Some use the TG-43 dose-calculation formalism and others do not. Prior to adopting the recommendations presented in this report, the medical physicist should implement the dose-calculation data and technique recommended by this report on his/her treatment planning system and quantitatively assess the influence of this action on dose delivery. This is best performed by
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comparing the dose distribution for typical implants based on the revised dose-calculation procedure with those based upon the currently implemented algorithm for the same seed locations, source strengths, and dose-calculation grid. The potential impact of these dose-calculation modifications on dose delivery relative to the current dose-calculation technique should be discussed with the appropriate radiation oncologist before clinically implementation. Finally, the comparison of old and new dose-calculation algorithms for the same seed input data, and the resultant decisions that may impact clinical dose delivery, should be documented for future reference and for regulatory purposes.
Dose-Calculation Formalism RTP systems are divided into those that comply fully with the TG-43 formalism, and those that do not. Full compliance is defined here as the use of equation (1) or one of its two approximations, given in equations (10) and (11). Full compliance also requires the use of the dose-rate constant, appropriate values of the radial dose function, and the 1-D or 2-D anisotropy functions that are provided in this protocol. For RTP systems that use the TG-43 dose-calculation formalism and permit customer input of dosimetry parameters, one should enter (or verify the correct entry of) the recommended parameters, and check the accuracy of the dose calculation algorithm. The tabulated data provided in this protocol should be used with such systems. In other cases, coefficients in an equation, e.g., 3rd to 5th order polynomial fits, describing the TG-43 parameters may be required. In these cases, the RTP-specific model or formula should be fit to the data provided by this protocol. For example, for systems that require a third-order (or higher) polynomial fit of the radial dose function, the clinical physicist is responsible for obtaining the best-fit coefficients by fitting the polynomial equation to appropriate gX(r) data from this protocol. Before implementing the dose-calculation model, it is necessary to evaluate the quality of the parametric fit. Deviation of the fitted data from those recommended in this protocol should be less than ±2% which will yield a dose-rate calculation accuracy of ±2% over the distance range of interest. The medical physicist shall take responsibility for verifying the accuracy of dosimetry data, whether the manufacturer or the user actually enters the data. The AAPM recommends using RTP systems that comply fully with the TG-43 formalism, whenever possible. However, some RTP systems do not use the TG-43 dose-calculation algorithm. In most cases, one can devise a method to force the algorithm to generate the single-source dose-rate distributions recommended here by using modified values for the dosimetry parameters required by the RTP system. This conversion should be performed with care. As with RTP systems based on the TG-43 dose calculation formalism, one should assure that the RTP system is generating correct single-source dose-rate data by creating a single-source treatment plan with the modified parameters before clinical use. Meigooni et al. (1995) have described an example of this approach. The methods used to arrive at modified data, as well as records of the evaluation of the RTP system, should be documented carefully and retained for use following installation of upgrades and for inspection by regulatory authorities. Extreme caution should be exercised whenever parameters should be entered or displayed that have units that do not match the units on documentation printed by the RTP system or displayed on its monitor. Procedures should be developed and documented to describe exactly how the modified data and parameters are related to the non-TG-43 parameters assumed by the RTP system. These procedures should address both clinical treatment planning practices and chart-checking procedures. Ratios of the unconventional units to the conventional units should be supplied, to facilitate review of the planning method. Because this approach is prone to errors in implementation or interpretation, this method should be used as the last resort.
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Acceptance Testing and Commissioning Thorough acceptance testing and commissioning shall be carried out prior to clinical implementation of a new RTP system or new source model. The user should document the results of these tests both for later reference, and for compliance with applicable regulations. As a minimum, calculations of the dose-rate distribution shall be performed for a single source of each type to be used clinically. The recommendations of the AAPM TG-40 (Kutcher et al., 1994), AAPM TG-53 (Fraass et al., 1998), and AAPM TG-56 (AAPM 1997) should be followed. The user should determine the range of distances from the source over which the RTP calculations meet the recommended level of accuracy (±2%). If deviations between calculated results and the published data exceed ±2%, the deficiencies of the RTP system should be documented and further investigated by the user. This is especially important for RTP systems that fit a model to published data, because such models are prone to large errors outside the range of the reference data. In the high-dose-gradient regions close to a source, particularly near the ends of a source, the acceptable error may need to be larger. These deficiencies should be considered when evaluating treatment plans, and when considering the purchase of an RTP system. To perform comparisons at the recommended level of precision, numerical point dose rates calculated by the RTP system, rather than measured diameters of plotted isodose contours, should be used. The user should evaluate any deviations from the recommended data, taking into consideration the uncertainty of calculations at very small distances. The calculation matrix should be adjusted to a value appropriate for the high gradients near a source. For low-energy photon-emitting sources, grid spacing should not exceed 1 × 1 × 1 mm3; this size generally represents a reasonable compromise between calculation speed and accuracy. Isodose curves should be generated as part of the commissioning and continuing quality assurance procedures, but should be viewed as a test of the spatial accuracy of the graphic dose display function of the RTP system, rather than a definitive test of the underlying dose-calculation algorithm. Because comparisons should include both point dose-rate calculations and the placement of isodose lines, the user should also ensure that the RTP system and its graphical output devices cause isodose curves to appear in the correct locations relative to corresponding point calculations.
Source Calibrations For calibrating radioactive sources, the AAPM has previously recommended that users not rely on the manufacturer’s calibrations, but instead confirm the accuracy of source-strength certificates themselves by making independent measurements of source-strength that are secondarily traceable to the primary standard maintained at NIST according to AAPM TG-40. For patient treatments, AAPM further recommended that all clinically used sources bear calibrations that are secondarily traceable to the primary standard. AAPM TG-56 (AAPM 1997) defines “direct traceability,” “secondary traceability,” and “secondary traceability with statistical inference” as follows: “Direct traceability is established when either a source or a transfer instrument (e.g., well chamber) is calibrated against a national standard at an ADCL or at NIST itself.” “Secondary traceability is established when the source is calibrated by comparison with the same radionuclide and design that has a directly traceable calibration or by a transfer instrument that bears a directly traceable calibration.” “Secondary traceability by statistical inference is established when a source is one of a group of sources of which a suitable random sample has direct or secondary traceability.”
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This protocol, in accord with the previously published AAPM guidance (TG-56, TG-40), also recommends that all sources implanted into patients have measurements based upon secondary traceability. Normally, vendors should be expected to provide calibration certificates that document secondary traceability to NIST for their sources. Each institution should maintain a means for verifying vendor calibrations by air-kerma strength measurements with secondary traceability. Source sampling and instrument quality assurance guidelines are documented in TG-56. Source strength verification with secondary traceability can be achieved either by means of a chamber bearing a calibration that is directly traceable to the appropriate national standard, or by comparison to a source of the same model having a calibration that is directly traceable to the national standard. These methods are described. 1. Secondarily traceable calibrations at the institution using a transfer instrument. For brachytherapy sources, calibrating sources with secondary traceability is best done with a well-type ionization chamber having an ADCL-assigned, directly traceable calibration coefficient determined for the source model to be used. TG-56 recommends returning the chamber to the ADCL at intervals of two years for recalibration. 2. Secondarily traceable calibrations at the institution by source intercomparison. When NIST or an ADCL provides an air kerma strength value for a given source, that source is then said to have a directly traceable calibration. The user can then use this source to calibrate his/her transfer instrument, such as a well chamber. The well chamber in this situation is said to have a calibration coefficient bearing secondary traceability. To accomplish this, the user may obtain a source from a vendor, send it to NIST or an ADCL and obtain an air kerma strength value for that source. The user can then use this source to calibrate his or her well chamber. The well chamber can then be used to confirm the vendor-stated source strengths of other sources of the same model before they are used clinically. If the chamber is calibrated by the ADCL, the calibration coefficient is directly traceable. If the calibrated source method is used, the user is responsible for transferring the calibration to his or her instrument. Regardless of which method is used, the well chamber should be checked for constancy on a regular basis using a long-lived source such as 137Cs. The AAPM TG-40 has recommended that a constancy check be performed at each used that the well chamber exhibit constancy to within ±2%. The clinical user is cautioned that the secondary traceability calibration procedure should meet some minimum standard of quality, i.e., that for the intercomparisons method, the total uncertainty should be comparable to that achievable using an instrument with a directly traceable calibration. For example, for the calibrated chamber or the calibrated source approach, the total expanded uncertainty (2σ = 95% confidence level) is obtained by adding in quadrature the uncertainties of instrument or source calibration at the ADCL (typically 2.4% including the uncertainty at NIST). It is emphasized that the precision of measurement in the well chambers is better than this (generally within ±0.5%). The typical total, expanded uncertainty at the user facility using a source transfer to a chamber is typically 3.0%. This is the maximum uncertainty that is relevant for secondarily traceable calibration. Any additional steps in the intercomparison process will increase the total uncertainty and thus will not meet the minimum standard of quality recommended in this protocol. The user is further warned that under no circumstances should a vendor or user calibration be used as a basis of institutional verification calibrations. Finally, it is important not to confuse a source bearing a calibration with direct or secondary traceability with a “calibrated source” obtained from the source manufacturer. Use of a source calibrated by a manufacturer is not an acceptable alternative for providing in-house calibrations with traceability to NIST or an ADCL.
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Williamson, J. F., B. M. Coursey, L. A. DeWerd, W. H. Hanson, R. Nath, and G. Ibbott. (1999a). “Guidance to users of Nycomed Amersham and North American Scientific, Inc., I-125 interstitial sources: Dosimetry and calibration changes: Recommendations of the American Association of Physicists in Medicine Radiation Therapy Committee Ad Hoc Subcommittee on Low-Energy Seed Dosimetry.” Med Phys 26:570–573. Williamson, J. F., B. M. Coursey, L. A. DeWerd, W. F. Hanson, R. Nath, M. J. Rivard, and G. Ibbott. (1999b). “On the use of apparent activity (Aapp) for treatment planning of 125I and 103Pd interstitial brachytherapy sources: Recommendations of the American Association of Physicists in Medicine Radiation Therapy Subcommittee on Low-Energy Brachytherapy Source Dosimetry.” Med Phys 26:2529–2530. Williamson, J. F., B. M. Coursey, L. A. DeWerd, W. F. Hanson, R. Nath, M. J. Rivard, and G. Ibbott. (2000). “Recommendations of the American Association of Physicists in Medicine on 103Pd interstitial source calibration and dosimetry: Implications for dose specification and prescription (AAPM Report No. 69).” Med Phys 27:634–642. Williamson, J. F., W. Butler, L. A. DeWerd, M. S. Huq, G. S. Ibbott, Z. Li, M. G. Mitch, R. Nath, M. J. Rivard, and D. Todor. (2005). “Recommendations of the American Association of Physicists in Medicine regarding the impact of implementing the 2004 Task Group 43 report on dose specification for 103Pd and 125I interstitial brachytherapy.” Med Phys 32:1424–1439. Yu,Y., L. L. Anderson, Z. Li, D. E. Mellenberg, R. Nath, M. Schell, F. M. Waterman, A. Wu, and J. C. Blasko. (1999). “Permanent prostate seed implant brachytherapy: Report of the American Association of Physicists in Medicine Task Group Report No. 64.” Med Phys 26:2054–2076. Zelefsky, M. J., and W. F. Whitmore, Jr. (1997). “Long-term results of retropubic permanent 125iodine implantation of the prostate for clinically localized prostatic cancer.” J Urol 158:23–30.
Chapter 17
Treatment Planning Considerations of Brachytherapy Procedures Ali S. Meigooni, Ph.D.,1 and Robert E. Wallace, Ph.D.2 1 University of Kentucky, Lexington, KY 2 Cedar-Sinai Medical Center, Los Angeles, CA Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 326 Calculation Algorithm . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 326 Linear Source Approximation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 326 TG-43 Algorithm . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 326 Sievert Integral . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 327 Interpolation Method . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 329 Point Source Approximation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 329 TG-43 Formalism . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 329 Traditional Algorithm . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 330 Curvilinear Line Source Approximation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 330 Line Source . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 331 Stranded Source . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 331 Source Data Entry . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 331 Linear Source Approximation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 331 2-D TG43U1 Parameters . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 331 Tissue Attenuation Coefficients . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 333 Look-up Table . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 333 Point Source Approximation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 333 1-D TG-43U1 Parameters . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 333 Gamma Factors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 334 Look-up Table . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 334 Specific Features in Planning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 334 Calculation Geometries . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 334 Imaging Support . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 335 Radiographic Reconstruction: Orthogonal Films . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 335 Radiographic Reconstruction: Linear Stereo-Shift . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 335 Radiographic Reconstruction: Rotation Stereo-Shift . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 336 Radiographic Reconstruction: Three or More Film Noncoplanar Film and Fiducial Jigs . . . . . . 336 Volumetric Reconstruction: DICOM Image Sources . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 337 Volumetric Reconstruction from CT Image Series . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 337 Real-Time Planning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 338 Quality Control of Treatment Planning Systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 338 AAPM Task Group 43 Recommendations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 339 AAPM Task Group 53 Recommendations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 340 AAPM Task Group 56 Recommendations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 342 AAPM Task Group 64 Recommendations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 342 Implementations and Factors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 344 Recommendations for Data Entry in Varian Planning Systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . 344 Recommendations for Data Entry in Prowess Planning System . . . . . . . . . . . . . . . . . . . . . . . . . . . 344 Recommendations for Data Entry in the ADAC Pinnacle Planning System . . . . . . . . . . . . . . . . . . 345 Recommendations for Data Entry in Nucletron TheraplanTM and SPOTTM Brachytherapy Planning Systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 345 Shortcomings and Recommendations in the Present Planning Systems . . . . . . . . . . . . . . . . . . 345 Linear Source Calculations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 345
325
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Ali S. Meigooni and Robert E. Wallace Point Source Calculations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 346 Interpolation and Extrapolations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 346 Strategies To Implement TG43U1 Parameters in Systems that Do Not Support the TG-43 . . . . . . 346 Nomogram . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 346 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 347
Introduction The continuous improvements of the brachytherapy treatment modality is partly due to the advancement of the technical aspects of source design and treatment delivery, and partly due to the development in the treatment planning systems and quality controls in the program (Russel and Blasko 1993; Nori 1992; Porrazzo et al., 1992; Blasko et al.,1995; Sommerkamp, Ruppercht, and Wannenmacher 1988; Willins and Wallner 1998; Merrick et al., 1998, 2000; Fuks et al., 1991; Kumar and Good 1986; Tapen et al., 1998). Although there are still several shortcomings in these aspects of brachytherapy, a review of the existing programs and protocols will give us guidelines for future activities. In this chapter, the existing treatment planning methods will be reviewed and the shortcomings of each technique will be presented.
Calculation Algorithm Dose calculations in brachytherapy are divided into two general categories of linear and point source approximations. The main distinctions between these two categories can be related to the fact that if the distribution of the activities within the source and the its self absorption can be neglected or it has to be considered in the dose calculations. Generally, when the calculation points was located at a distance greater than twice of the active length of the source, or the energy of the photon emission was greater than 100 keV, the source can be assumed as an isotropically emitting radiation source, and otherwise it will be assumed as a linear source. In the following sections, the dose calculation methods for each of these two categories are presented.
Linear Source Approximation Dose distributions around a linear source are generally calculated using one of the following three methods. TG-43 Algorithm In 1988, Task Group 43 (TG-43) of the American Association of Physicists in Medicine (AAPM) was started, and introduced a new protocol for dosimetric characterization and dose calculation formalism for brachytherapy sources (Nath et al., 1995). Since then, many investigators have determined the dosimetric characteristics of various source models and hence, some shortcomings of the original TG-43 protocol have been discovered and hence an updated AAPM TG-43 protocol (also known as AAPM TG-43U1) has been introduced (Rivard et al., 2004). The formalism introduced by this protocol is as follows:
D" (r, θ ) = SK ⋅ Λ ⋅
GL (r , θ ) GL (r0 , π / 2 )
⋅ gL (r ) ⋅ F (r, θ ),
(1)
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where · D (r,θ ) is the dose rate at point P(r,θ) as shown in Figure 1; SK is the air kerma strength; Λ is the dose rate constant; GL(r,θ) is the geometric function; gL(r) is the radial dose function; and F(r,θ) is the 2-D anisotropy function. The readers are referred to the AAPM TG-43U1 protocol for a detailed description of these parameters (Rivard et al., 2004). Moreover, the members of this task group are reviewing the dosimetric characteristics of the different brachytherapy source models, and derive the consensus of the multiple measured and Monte Carlo simulated data for each source type. These consensus datasets for source parameters will be available in their future publications. The TG-43 formalism can easily be adapted in any computerized treatment planning systems. Sievert Integral For many years dose distributions around a linear brachytherapy source were calculated using the Sievert Integral (Sievert 1921). In this method, a linear source subdivided into several equal segments (Figure 2) where each segment was small enough that could be assumed as a point source. The integration of the exposure rate contributions from each segment to a given point was then calculated using point source approximation. The final integrated exposure rate was presented as: I ( x, y ) =
M eq .Γ Ra Ly
eµ t
'
∫
θ2
θ1
e− µ t sec θ dθ '
'
(2)
where Meq = Total source strength in term of mgRaEq. ΓRa = Exposure rate constant of 226Ra source, with 0.5 mm of platinum filter. L = Active length of the source. µ′ = Effective attenuation coefficient of the filter (source capsule). t = Thickness of the source capsule. If the source strength is specified in term of air kerma strength (SK), this equation will changed to I ( x, y ) =
SK eµ t W Ly( ) e
'
∫
θ2
θ1
e− µ t sec θ dθ ' , '
(3)
where W/e is defined as work function that is the average energy absorbed per unit charge of ionization in air. The original Sievert integral has corrected the exposure for the attenuation of the radiation by the source capsule and eliminated the self-absorption by the core of the source. However, Shalek and Stovall (1969) had modified this integral to consider the self-absorption of the source, as well. Depending on the type
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Figure 1. Schematic diagram of a brachytherapy source, used in calculation of dose rate at point P, from a linear source, using the AAPM TG-43 formalism.
Figure 2. Schematic diagram of a brachytherapy source, illustrating geometric relationships used in calculation of exposure at point P, from a linear source, using the Sievert Integral method.
of source and filtration, the energy spectrum may be significantly altered by the filter. This Sievert Integral can be evaluated by numerical analysis. Because this integral uses the energy absorption coefficient, the underlining assumption in this method is that the emitted energy fluence is approximately attenuated by filter thickness traversed by the photon. This is an approximation that has been shown to work well for 226Ra and 192Ir sources in region bounded by the active source ends (Shalek and Stovall 1969; Williamson, Morin, and Khan 1983). However, Monte Carlo simulations (Williamson, Morin, and Khan 1983) have shown that beyond the end of the active source region, the Sievert approach introduces significant error and practically breaks down in the extreme oblique directions. The treatment-planning systems using the Sievert integrals, acquire the exposure rate constant of the source, its geometric information such as the active length, physical length, capsule thickness, and the attenuation coefficients of the capsule and the core of the source. In addition, the values of exposure to dose conversion factors (f-factor), half-life, and the radial dose function or tissue air ratio of the source are needed to generate a dose distribution matrix in an along-away format. The tissue air ratio for some of the initial sources introduced for brachytherapy treatments were determined by Meisberger et al. (Meisberger, Keller, and Shalek 1968). These values were given in the forms of a third order polynomial coefficient, more commonly known as Meisberger coefficients. If the source strength is given in term of the air-kerma strength, the exposure rate constant value, Γ, should be replaced (W/e)–1= 1.142.
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Interpolation Method A third method of dose calculations around a linear source is interpolation of a set of data points that has been determined using either experimental method or theoretical calculations. The along-and-away table of Krishnaswamy (1972) for the 137Cs source or tabulated data in Paterson-Parker’s system (Meredith 1967) and Quimby system (Glasser et al., 1961) for radium equivalent sources are among the most commonly used data in calculations of the dose rate from a linear source, using the interpolation methods. These tabulated data are in a matrix form with a given format that can easily be applied for clinical applications, particularly for a quick double check dose calculation with a simple calculator. It should be noted that some of these tables are only applicable for specific source geometry and radionuclides. For example, the along-and-away table of Krishnaswamy (1972) is only valid for the 137Cs radionuclide with the source geometry that has been specified in that publication. Although this method of calculation is simple, its accuracy is not warranted. Moreover, there are different tabulated data from various source models. Thus, it is not easy to prepare consensus datasets.
Point Source Approximation Although most of the brachytherapy sources are linear in shape, however, under the following conditions the source geometry can be ignored and an isotropic point source can be utilized during the treatment planning of an implant patient. These conditions are: (a) The distance from the source is far enough (r > 2L) that the distribution of the activity within the source will not have any impact to dose calculation at a given point. (b) The random orientation of the brachytherapy sources within the implant volume will minimize the effects of dose anisotropy within the implant volume. The following methods are describing the dose calculations using the point source approximation technique. TG-43 Formalism The 2-D formalism introduced in the AAPM TG-43U1 protocol can be simplified as D" (r ) = SK ⋅ Λ ⋅
GL (r, θ ) GL (r0 , π / 2 )
⋅ gL (r ) ⋅ ϕ ave (r ),
(4a)
Or
r D" (r ) = SK ⋅ Λ ⋅ o g p (r ) ⋅ ϕ ave (r ), r 2
(4b)
Where the ϕan(r) is called 1-D anisotropy function at the radial distance of r which is calculated using the following equation as: π •
ϕ an (r ) =
∫ D(r,θ ) Sin(θ ) dθ 0
•
2 D (r, θ o )
(5)
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where gp and gL are the radial dose functions of the source determined using point or line source approximations, respectively. Although the updated TG-43 formalism recommends use of equation (4a) in order to achieve better accuracy at short distances (r < 1 cm), commercially available treatment planning systems generally utilize equation (4b). Therefore, dose calculations with the point source approximation require the dose rate constant of the source, 1-D anisotropy function, and radial dose function (i.e., gp) of the source. In case of the permanent implant, the half-life of the source also would be required. Traditional Algorithm The traditional algorithm for dose calculations around brachytherapy sources were based on Γ of the source corrected for the source anisotropy constant measured in air, apparent source activity (App), and tissue air ratio (B(r)) as shown below: 1 D" (r ) = Gamma ⋅ App ⋅ 2 ⋅ f ⋅ T ⋅ DC ⋅ ⋅B(r ) ⋅ 100 , r
(6)
where, f is the exposure to dose conversion factor (R/cGy), T is the exposure time (hours), and DC is the source decay correction during the treatment defined as: DC = 1.443 ⋅ Half − Life ⋅ [1 − exp(0.693 ⋅ T / Half − Life)
B(r ) = [ exp(− µ ⋅ r )
]
]
(7)
(8)
For low-energy photon-emitting brachytherapy sources, this calculation method has been considered to be in error since the exposure rate measured in air and its conversion to dose to tissue has been inconsistent between the user and vendors. Moreover, the vendor determined Aapp by measuring the exposure rate of the source at a distance and dividing that by the gamma constant of the source, which was not necessarily the same as the Γ values used by the physicists.
Curvilinear Line Source Approximation Line Source Flexible sources such as 192Ir wires may be bent into arbitrary shape. There is no simple or general mathematical procedure to calculate the dose at a point. The most conventional method is to proceed by subdividing the curved source into a train of small source segments. The total dose to a given point is calculated by superposition of the dose contributions from each source segments (Rosenwald 1997; Marinello and Pierquin 1997; Meigooni et al., 2004). It is not necessary for all the segments to have the same length, rather the length of each segment should be selected in order to reproduce the curvature of the wire. Also, the tabulated dose distributions around a source wire, as a function of the length of the wire are available for some of the brachytherapy sources (Rosenwald 1997). One may also present such results graphically in an easy-to-use form. Figure 3 shows a sample of isodose curves (also known as isodose “snail”) for 192 Ir wires as a function of active lengths ranging from 1 to 7 cm (Pierquin and Marinello 1997; Schlienger et al., 1970). From these data, one can calculate the doses actually delivered by multiplying the tabulated data by the reference source strength and the treatment duration time. The isodose values obtained for the
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Figure 3. The “Snail” isodose for 192Ir wires of length ranging from 1 to 7 cm. Each curve corresponds to a given dose rate (cGy/day) for sources of unit linear reference kerma rate in central plane of the wire. (Reproduced from Pierquin and Marinello (1997) with permission of Medical Physics Publishing.)
central plane are valid as long as the plane of calculations is at a distance less than one-quarter of the length of the wire. Stranded Source A stranded source with the center-to-center spacing between the adjacent sources of S, can be closely approximated by a continuous linear source of length L = N*S, where N = number of souces in the strand. It should be noted that the length L is larger that greater than the distance stranded sources. The excess distances on each separating the extreme point sources of the end of the equivalent continuous line source and stranded sources are S/2 (Figure 4).
Source Data Entry Linear Source Approximation 2-D TG43U1 Parameters The AAPM TG-43U1 2-D dosimetric characteristics (i.e., dose rate constant, radial dose function, and 2D anisotropy function) can be easily entered in the source files of any treatment planning systems. These parameters are given either as a tabulated data or in the form of polynomial coefficients. It should be noted that, the polynomial coefficients are only warranted within the range of the data that the coefficients have been fitted. Any extrapolation from these coefficients may lead to large and unexpected errors. The users are recommended to verify the algorithm and methodology of dose calculations outside of the given range of data. The 1995 TG-43 protocol recommends the 5th order polynomial fit [equation (9)] to the radial dose function of the sources as, g(r ) = a0 + a1r + a2 r 2 + a3r 3 + a4 r 4 + a5 r 5
(9)
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Figure 4. A tandem of N sources in a strand form compared with an 192Ir wire with continuous activity distribution. (Reproduced from Pierquin and Marinello (1997) with Permission of Medical Physics Publishing.)
However, several investigators have shown that the polynomial fit could not accurately represent the radial dose functions with large variations at short or extended distances (Furhang and Anderson 1999; Meigooni et al., 2003). Furhang and Anderson (1999) have introduced the double exponential fit [equation (10)] and Meigooni et al. (2003) have introduced the modified 5th order polynomial fit [equation (11)] for more accurate representation of these radial dose functions. g(r ) = C1e( − µ r ) + C2 e( − µ r ) 1
(
(10)
2
)
g(r ) = a0 + a1r + a2 r 2 + a3r 3 + a4 r 4 + a5 r 5 e− br
(11)
Figures 5a and 5b shows a sample of these types of curve fits. It should be noted that in equation (10), C2 was recommended to be a negative value. Therefore, introducing a positive number for the initial value of this parameter during the curve fit may lead to an erroneous result. These authors recommend using either equation (10) or (11) in the brachytherapy treatment planning systems. Meanwhile, the physicist could utilize these fitting routines in order to generate a larger number of data points from a coarse dataset. This dataset can be used in treatment planning systems, which incorporates interpolation methods in patient dosimetry. The 2-D anisotropy functions can be introduced in tabulated form. However, it should be noted that, some of the treatment planning systems require the 2-D anisotropy functions for the angles ranging from 0 to 90° degrees, and some require data from 0 to 180°. It is interesting to know that, although the 2-D dosimetric parameters of the source can be entered in the treatment planning systems, they may not be using that properly. For instance, some prostate implant treatment planning systems assume that the source directions are normal to the CT or ultrasound image plane, and do not use true source directions for the treatment planning. For true linear source calculations, the users should be able to introduce the location and direction of the seeds by digitizing either the two ends of each source, or introduce the center and orientation of the source (e.g., digitizing the catheter or applicators holding the source).
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Figure 5. Comparison of the 5th order polynomial fit with (a) modified polynomial fit and (b) double exponential fit for radial dose function of BrachySeed Pd-103 model Pd-1 source.
Tissue Attenuation Coefficients The traditional treatment planning systems acquire the source geometric information such as the active length, physical length, capsule thickness, and the attenuation coefficients of the capsule and the core of the source. In addition to Γ, f-factors, and half-lives, Meisberger coefficients are needed to generate a matrix of dose distribution in an along-away. Look-up Table A look-up table configured in “along and away” Cartesian rectangular coordinates centered on a line source center can be used to represent the two-dimensional distribution of dose rate about a prototype source of a given design, having unit strength. Such tables are interderivable with the polar-coordinate based TG43 dose-rate tables for unit air-kerma strength. The advantage of such tables is that dose calculation points are typically distributed on rectangular grids in treatment planning systems. The process of mapping the dose-rate contributed by a given source is simplified by reduction to a process of interpolation, and the planning system speed improves. Printed look-up tables allow a clinical physicist to estimate the dose to selected points by a “hand calculation.” Look-up tables for other quantities than dose rate are possible. For example, a treatment planning system may not fully support TG-43 parameters for dose calculation. Suppose that such a system did not support the geometry factor, G(r,θ), and the anisotropy function, F(r,θ), as distinct entities but supported one normalized relative dose table in two dimensions (either “along and away” or the polar coordinate systems). Given the TG-43 parameters, one could compute the product, G(r,θ)*F(r,θ), and fill in the table. Thus, one may combine TG-43 parameters for a given source to fit the requirements of the specific treatment planning system.
Point Source Approximation 1-D TG-43U1 Parameters The 1-D TG43U1 dosimetric characteristics (i.e., dose-rate constant, radial dose function, and 1D anisotropy function) of the brachytherapy sources can be easily entered in the source files of any treatment planning systems. Similar to the linear source data entry, these parameters can be given either as a tabulated data or in the form of polynomial coefficients. Although the 2004 TG-43U1 report recommends use of equation (4a) in order to achieve a dose calculation accuracy at short distances (r < 1 cm), commercially available treatment planning systems generally utilize equation (4b). Therefore, dose calculations
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with this methodology require the point source approximation radial dose function (i.e., gp) of the source. Similar to the above description for data entry of the linear source models, care should be taken for introducing the radial dose function of the brachytherapy sources. Both, the original and updated TG-43 protocols, recommend the tabulated radial dose function using the 5th order polynomial fit. The 1-D anisotropy functions can be introduced in tabulated form. The updated TG-43U1 has given the 1-D anisotropy of brachytherapy seeds for radial distances ranging from 0.5 to 7 cm. This protocol – recommend using the modified dosimetry parameters, g′(r) and φ a′n defined as, g′(r ) = g (r ) ⋅ φ an
(12)
φ an′ = 1
(13)
Gamma Factors Γ values for a brachytherapy source are defined as the exposure rate at a reference distance along the transverse bisector of the source, per unit source strength. Γ has units of either R◊cm2/(h◊mCi), if the source strength is expressed in terms of Aapp, or R⋅cm2/(h⋅mg) if the source strength is expressed in terms milligram radium equivalent. Γ values for commonly available brachytherapy sources can be found in most of the radiation therapy physics textbooks (Meigooni, Saw, and Nath 1997; Khan 1992). Source strength units must be consistent with Γ units entered into a treatment-planning system. Look-up Table A look-up table configured in “along and away” Cartesian rectangular coordinates centered on a line source center can be used to represent the two-dimensional distribution of dose-rate about a prototype source of a given design, having unit strength. In principle, such a table could be derived for a point source. Such tables are interderivable with the polar-coordinate based TG-43 dose-rate tables for unit air-kerma strength. For the case of a line source, look-up tables for quantities other than dose-rate are possible. For example, a treatment planning system may not fully support TG-43 parameters for dose calculation. Suppose that such a system did not support the radial dose function g(r), and the anisotropy factor, φan(r), as distinct entities, but instead supported a relative dose-rate table in one dimension. Given the TG-43 data, one could compute the product, g(r)* φan(r), and fill in the table by calculating the distance r(x,y) for points on a rectangular grid. Thus, requirements of the treatment-planning system can be met with recommended TG43 consensus dataset values without requiring that the planning system be modified.
Specific Features in Planning Calculation Geometries The geometry of the brachytherapy implants is normally determined by the shape and location of the target volume. For instance, brachytherapy treatment of endobroncheal cancer is normally performed using a linear source, while an interstitial breast implant is most commonly either a single or double plane implant. Whether calculations are pursued at selected points, in a plane, in a set of parallel or orthogonal planes, or at a number of points regularly or randomly distributed in a volume, each approach will requires a coordinate system representing points that map from the virtual world of the plan to the real world of patient
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anatomy. The dimensions, locations, and often the orientations of the brachytherapy sources relative to the treatment volume must therefore be clearly identified. The relation between the implant sources and target volume can be identified using one of the following imaging techniques. The relationship of the target volume to the plan coordinates can be similarly defined.
Imaging Support Radiographic Reconstruction: Orthogonal Films Orthogonal radiographs are taken at right angles, with the central axis of the x-ray beams meeting approximately in the middle of the implant. Usually, AP and lateral films, exposed isocentrically, provide such geometry. The coordinate system is conventionally established with the x-axis from the right to the left of the patient, the y-axis from inferior to superior, and z-axis from posterior to anterior. The AP film represents a magnified view of the implant image projected onto x-y plane, while the lateral film presents the image projected onto the z-y plane. The origin of the coordinate system is chosen to be a common point on both films such as one end of a source or applicator. For a typical treatment planning, the origin and the x-axis on the AP film and y-axis on the lateral film must be identified before brachytherapy source localization can be performed. The points of interest must be identified relative to the applicators (i.e., point A) or relative to the fiducial markers (i.e., bladder) placed in the patient. The orthogonal radiographic technique is not applicable for implants where the sources are masked by an object or skeletal bone. For example, the brachytherapy sources in the pelvic area for the GYN or prostate implant patients with femoral prostheses cannot be seen in the lateral film. Moreover, the brachytherapy seeds in the prostate implant patients are not clearly visible in the lateral film since the femoral bone easily blocks them. The alternative methods for these patients are linear or rotational stereoshift techniques in additional to CT, MRI, and ultrasound imaging techniques as described below. Radiographic Reconstruction: Linear Stereo-Shift The linear stereo-shift method of source localization consists of taking two radiographs of the patients from the same view, but shifting either the patient or x-ray tube by a certain distance (i.e., 20 cm) between the two exposures. Figure 6 shows the principle of the linear stereo-shift technique, for identification of two sources. Using the similar triangle rules for the two triangles of T1BT2 and B1BB2, the coordinates of point A can be related to those of point B as: d y1 + s − y2
=
F−Z Z
.
(14)
Also, from similar triangles of T1AT2 and A2AA1 it can be found that: d s
=
F− f f
.
(15)
From these two equations one can find: Z=F
( F − f )( y2 − y1 ) − d . f ( F − f )( y2 − y1 ) − d .F
.
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Figure 6. Schematic diagram of the source and target geometry in the linear stereo-shift radiographic technique.
Radiographic Reconstruction: Rotation Stereo-Shift The stereo-shift technique can be achieved by rotating the x-ray tube, instead of shifting that in a linear fashion. This procedure is performed with a x-ray machine with a rotational capability. Figure 7 demonstrates a typical rotational stereo-shift technique. The calculations of the coordinates is the same as the linear stereo-shift, except for the distances between the two target position, d, which is calculated as: d = 2 R .sin(θ / 2 )
(17)
Other rotational methods exist, using films that are orthogonal to the beam central axis (i.e., the dashed lines in Figure 7). An excellent resource for x-ray photogrammetry methods can be found in Hallert (1970). Radiographic Reconstruction: Three or More Noncoplanar Films Determination of dose distribution in complex implant geometry, such as prostate implant, may require more complex imaging technique than just orthogonal film imaging. In principle, a set of three (or more) non-coplanar images can be used with a fiducial jig having four or more radiopaque markers to reconstruct the locations of sources in three dimensions (Altschuler, Findla and Epperson 1983; Altschuler and Kassaee 1997; Amols and Rosen 1981; Biggs and Kelly 1983; Jackson 1983; Rosenthal and Nath 1983; Su et al., 2004; Todor et al., 2002). The non-coplanarity of the beams provides views that can be useful to sort seeds, which may be occluded in other images by anatomical structures or other sources. Use of a jig with embedded markers set out in a three-axis grid provides a coordinate system. The requirement of four or more markers improves the determination of the three-dimensional (3-D) coordinate system from the radiographic images of the fiducial markers. Since the geometry of the jig is known, the coordinate transformations can be determined via linear least squares or algebraic estimation. The greater the number of fiducial markers, the better is the estimation. The same is true of the number of films. No attempt
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Figure 7. Schematic diagram of the source and target geometry in the rotational stereo-shift radiographic technique.
will be made here to provide details of such methods due to their complexity. However, there has been much work in the field on such methods and Su et al. (2004) is a clear point of departure. Volumetric Reconstruction: DICOM Image Sources The Digital Imaging and Communications in Medicine (DICOM) standard was created to aid the distribution and viewing of medical images, such as CT scans, MRIs, and ultrasound (NEMA 2004). The National Electrical Manufacturers Association (NEMA) introduced this format of distribution of images. A single DICOM file contains both a header (which stores information about the patient’s name, the type of scan, image dimensions, etc), as well as all of the image data (which can contain information in three dimensions). Most modern brachytherapy planning systems have facilities to import and use images (and image sets) that are presented via some DICOM mechanism. The DICOM standard specifies the format of files containing the medical (or other, dose and 3-D structures, for example) images and specifies the requirements for file transport and interchange using digital media. Exchange may be facilitated by copying to a portable, interchangeable Compact-Disk (CD) ROM, through the agency of Universal Serial Bus (USB) based “solid-state disks”, or by direct network connection, email, etc. While such systems may have complex design and implementation, their use is relatively simple by virtue of the DICOM standard. Volumetric Reconstruction from CT Image Series There has been much effort over many years in defining and automating seed detection, localization, and (unique) identification in CT axial series (Bice et al., 1999; Brinkmann and Kline 1998; Li et al., 2001; Liu et al., 2003; Roy et al., 1993; Tutar, Managuli, and Shamdasani 2003). Tubic et al. (2001) described simulated annealing methods to sort and cull possible seed locations as an optimization problem. Holupka et al. (2004) describe a machine vision application using the elliptical Hough transform to find and sort
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possible seed locations. 2004). The common denominator of all these methods is use of axial CT source images that are reconstructions using sampled raw CT detector data. The spatial resolution is therefore limited by CT voxel dimensions and the Nyquist rate for sampling in and among the axial images. In this setting, the achievable spatial resolution of detected seed positions is on the order of one to two millimeters with no information about seed orientation. This derives from the fact that the pixel sizes are approximately the same diameter of cylindrical brachytherapy sources (about 0.8 mm) and the interslice spacing is the order of the length of the sources (i.e., 5 mm). Recently, Tubic and Beaulieu have reported promising initial findings using raw CT detector sinogram data to provide unambiguous seed localization with positional accuracy of 0.2 mm and orientation determination to within 3° about the source principal axis (Tubic and Beaulieu 2005). This method has logistical issues in that CT sinogram data , while a likely DICOM object, is not typically provided for use beyond image reconstruction in CT scanners. Real-Time Planning Traditionally, the ultrasound guided prostate seed implants were performed in three separate steps: (1) Volume study, a series of ultrasound images were collected 1 or 2 weeks prior to the implant. These images were used to pre-plane the treatment, and determine the seed activity, number of seeds and needles, and locations of the seeds. (2) Brachytherapy implant, the patient was prepared in the same position as the volume study, using the patient landmarks such as urethra and bladder. Then the needles, which were loaded by the seeds following the pre-plane, were placed in the patient. (3) Post implants CT-based evaluation. The patient dosimetries were evaluated using the CT images taken a month after the implant procedure. However, with the advancement of the technological aspects of the brachytherapy procedures, such as the Mick Applicator (Mick Radio-Nuclear Instruments Inc., Mount Vernon, NY) also known as “Mick Gun,” the first two procedures are combined into one. In this case, seeds of fixed activities are purchased prior to the implant, in special magazine that can be placed in Mick applicator. In this implant, the patient dosimetry is been performed as each needle is inserted in the patients, which is called “read time.” With this technique, once the treatment is over, the patient dosimetry is also completed.
Quality Control of Treatment-Planning Systems Several reports of AAPM task groups provide guidance on the quality assurance (QA) of treatment-planning systems in general with some mention regarding such systems for brachytherapy. The general recommendation is that each component of the system be tested with appropriate independent and standard methods. In each clinical use of a brachytherapy planning system, the resulting plans should be verified using an independent, if idealized, system. Hand calculation of dose at selected points assuming exposure and/or dwell times, source strengths, realistic source positions should be performed to verify a specific plan. Alternatively, a second computer-based system that implements the hand calculation (i.e., in a general purpose spreadsheet program) could suffice for the check if it had itself been earlier verified. A patient treatment can be defined by specific data: the input data for source position (and orientation), shield position (and orientation), source strength, prescription dose to a point or to an isodose surface enclosing a volume, and the dose-rate data for single sources of the design used (i.e., TG-43 or similar data). Using appropriate patient and treatment specific information, the responsible medical physicist can perform an end-to-end secondary calculation that validates the more complex plans generated by a dedicated treatment-planning system. Thus, the general statement is one of prudence and caution. While the present discussion revolves around treatment-planning systems, the QA of a part of the brachytherapy system cannot realistically be performed in a vacuum except to prove consistency of that system. The goal is to ensure that all compo-
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nents used in a brachytherapy treatment perform as intended and as expected. Each needs be tested, alone and as a part of the complex whole.
AAPM Task Group 43 Recommendations The updated TG-43 report (Rivard et al., 2004) suggests that, before clinical application of a new radiotherapy treatment planning (RTP) system or a new source model on an established RTP system, a thorough acceptance testing and commissioning must be carried out. This protocol recommends that the users document the results of their tests for later reference, as well as the compliance with applicable regulations. Moreover, it recommends that the dose rates calculated by the RTP system from a single source should be compared with the dose-rate distribution derived from the tabulated parameters and equations presented in the protocol. In addition, this recommendation suggests that the user should compare the RTP system calculations with the dose-rate distributions derived from the appropriate 1-D or 2-D data tables from the protocol. To facilitate this comparison for the brachytherapy sources included in this protocol, a tabulated transverse-plane dose rates as a function of distance was provided (Table 1). As has been recommended previously by the AAPM, this comparison should yield agreement within ±2% over all angles and over the range of radial distances commissioned (Kutcher et al. 1994; Fraass et al., 1998; Nath et al., 1997). If deviations between calculated results and the published data exceed ±2%, the deficiencies of the RTP system should be documented and further investigated by the user. This is especially important for RTP systems that fit a model to the published data, because such models are prone to large errors outside the range of the reference data. Moreover, this protocol suggests that, in the high dose-gradient regions close to a source, particularly near the ends of a source, the acceptable error may need to be larger. These deficiencies should be considered when evaluating treatment plans, and when considering the purchase of an RTP system.
Table 1. Dose Rates (cGy·h–1U–1) as a Function of Distance for Eight Brachytherapy Sources Using the 1-D Dosimetry Formalism of Equation (11) with Interpolation for gL(r) and φan(r) r (cm)
Amersham Amersham model model 6702 6711
Best model 2301
NASI model MED3631A/M
Bebig model I25.S06
Imagyn model IS-12501
Theragenics NASI model model 200 MED3633
0.5
4.119
3.937
3.978
4.112
3.922
3.426
3.014
3.184
1
0.995
0.911
1.004
0.986
0.95
0.815
0.587
0.626
1.5
0.413
0.368
0.419
0.42
0.398
0.334
0.199
0.215
2
0.213
0.186
0.217
0.207
0.205
0.169
0.0837
0.0914
3
0.0768
0.0643
0.0783
0.0746
0.0733
0.0582
0.0206
0.0227
4
0.0344
0.0284
0.0347
0.0325
0.0323
0.0246
0.00634
0.00697
5
0.0169
0.0134
0.0171
0.0157
0.0157
0.0118
0.00221
0.00247
6
0.0089
0.00688
0.00908
0.00811
0.0084
0.00592
0.000846
0.000933
7
0.0049
0.00373
0.00506
0.00429
0.00459
0.00328
0.000342
0.000364
(Reproduced from Table XV in the AAPM TG43U1 report: “Update of AAPM Task Group No. 43 Report: A revised AAPM protocol for brachytherapy dose calculations.” Med Phys 31: 633–674. M. J. Rivard, B. M. Coursey, L. A. DeWerd, W. F. Hanson, M. S. Huq, G. S. Ibbott, M. G. Mitch, R. Nath, and J. F. Williamson. Also available as AAPM Report No. 84. © 2004, with permission from AAPM.)
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To perform comparisons at the recommended level of precision, numerical point dose rates calculated by the RTP system, rather than measured diameters of plotted isodose contours, should be used. The user should evaluate any deviations from the recommended data, taking into consideration the uncertainty of calculations at very small distances. The calculation matrix should be adjusted to a value appropriate for the high gradients near a source. For low-energy photon-emitting sources, voxel size should not exceed 1×1×1 mm3; this size generally represents a reasonable compromise between calculation speed and accuracy. Isodose curves should be generated as part of the commissioning and continuing quality assurance procedures, but should be viewed as a test of the spatial accuracy of the graphic dose display function of the RTP system, rather than a definitive test of the underlying dose-calculation algorithm. Because comparisons should include both point dose-rate calculations and the placement of isodose lines, the user should also ensure that the RTP system and its graphical output devices cause isodose curves to appear in the correct locations relative to corresponding point calculations.
AAPM Task Group 53 Recommendations The AAPM Task Group 53 report assists the medical physicist in developing and implementing a comprehensive program of quality assurance for radiotherapy treatment planning (Fraass et al., 1998). This report was the first guidance on the topic of treatment planning QA from the AAPM, although there are several related reports, including the recent report from Task Group 40 on Comprehensive QA for Radiation Oncology (Kutcher et al., 1994). Further expansion of AAPM recommendations regarding treatment planning quality assurance is likely after the radiation oncology community accumulates some experience with the approach recommended in this report. In recent years, the increased complexity of the treatment-planning process required to support such procedures as conformal radiotherapy has led to the need for a comprehensive set of quality assurance guidelines that can be applied to treatment-planning systems that support this complex process. This task group has been charged by the AAPM to prepare this report recommending the scope and content of necessary quality assurance procedures and the frequency of tests, from acceptance testing, characterization and commissioning to routine QA of clinical system use. Radiotherapy treatment planning (RTP) has long been an important part of the radiotherapy treatment process, so assuring that the treatment-planning process is being performed correctly is thus an important responsibility of the radiation oncology physicist. In recent years, as 3-D and image-based treatment planning has begun to be practiced in numerous clinics, the need for a comprehensive program for treatment-planning QA has become even more apparent. The AAPM TG-40 report published an overall approach to QA for the therapy process (Kutcher et al., 1994), but this work includes only a very general discussion of brachytherapy treatment-planning QA issues. In this report, radiation oncology physicists were advised to create the appropriate QA program for the treatment-planning systems and processes used in their clinics. Although this QA program will vary widely between different clinics, use of this report should allow each clinic to concentrate its QA efforts on those areas of most importance. General Definitions and Aims The radiotherapy treatment-planning process is defined to be the process used to determine the number, orientation, type, and characteristics of the radiation beams or brachytherapy sources used to deliver a large dose of radiation to a patient in order to control or cure a cancerous tumor or other problem. Most often, treatment planning is performed with the assistance of a computerized treatment-planning system that helps the treatment planner and physician define the target volume, determine beam directions and shapes, calculate, and evaluate dose distributions. The AAPM TG-53 report codifies the planning process and identifies those that may require attention to maintain correct functionality. Their recommendations are found in Table 2.
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Table 2. The Clinical Treatment Planning Process 1. Patient Positioning and Immobilization • Establish patient reference marks/patient coordinate system. 2. Image Acquisition and Input • Acquire and input CT, MR, and other imaging information into the planning system. 3. Anatomy Definition • Define and display contours and surfaces for normal and critical structures. • Geometrically register all input data (CT, MR), including registration with initial simulation contours, films, patient position, etc. • Define target contours, generate 3-D target surface using surface expansion, import target information from multiple imaging modalities. • Generate electron density representation from CT or from assigned bulk density information. 4. Beam/Source Technique • Determine beam or source arrangements. • Generate beam’s-eye-view displays. • Design field shape (blocks, MLC). • Determine beam modifiers (compensators, wedges). • Determine beam or source weighting. 5. Dose Calculations • Select dose calculation algorithm and methodology, calculation grid and window, etc. • Perform dose calculations. • Set relative and absolute dose normalizations. • Input the dose prescription. 6. Plan Evaluation • Generate 2-D and 3-D dose displays. • Perform visual comparisons. • Use DVH analysis. • Calculate NTCP/TCP values, and analyze. • Use automated optimization tools. 7. Plan Implementation • Align (register) the real patient with the plan (often performed at a plan verification simulation). • Calculate Monitor Units or implant duration. • Generate hardcopy output. • Transfer plan into record and verify system. • Transfer plan to treatment machine. 8. Plan Review • Perform overall review of all aspects of plan before implementation. (Reproduced from Table 1-2 in the Task Group 53 report: Med Phys 25:1773–1829, “American Association of Physicists in Medicine Radiation Therapy Committee Task Group 53: Quality assurance for clinical radiotherapy treatment planning,” B. Fraass, K. Doppke, M. Hunt, G. Kutcher, G. Starkschall, R. Stern, and J. Van Dyk. Also available as AAPM Report No. 62. © 1998, with permission from AAPM.)
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While general in definition, each step identified above has either identical or direct analogues in brachytherapy treatment planning. Other AAPM task groups provide more recommendations.
AAPM Task Group 56 Recommendations The recommendations of the AAPM Task Group 56 for quality assurance of treatment-planning systems are scant (Nath et al., 1997). In fact, the authors reported that they found that “Relatively little has been written on QA of clinical treatment planning systems in general and even less is available specifically for brachytherapy treatment planning systems.” This report identifies those areas specific to brachytherapy treatment planning that need to be addressed in a QA program: 1. Methods to reconstruct source positions and orientations in 3D from a number of imaging sources: radiograph sets, and 3D CT or ultrasound image sets. 2. Methods to predict or project catheter trajectory from the same image sets. 3. Graphics and image display system(s) linearity and correctness. 4. Mechanisms to assign source strengths and placement durations (HDR dwell through permanent implant) to each individual source. 5. Dose calculation algorithms. 6. Dose distribution optimization, evaluation, and representation algorithms. 7. Hard-copy documentation: accuracy of numeric and text information, geometric accuracy of printed representations of source locations and dose distributions. To these we add: 8. Methods to reconstruct shield or filter positions and orientations in 3-D from a number of imaging sources: radiographs, and 3-D CT, or ultrasound image sets. TG-56 report summarizes QA recommendations in its Table VIII, reproduced here at Table 3, and also recommends that treatment specifications, times, positions, and dose should be verified by a secondary calculation. This is similar to the practice of checking monitor unit calculations for each planned external beam therapy. An independent check of each treatment is often possible using a copy of the data that was entered into the treatment planning system (in the form of look-up tables of dose rate at an array of positions about a single source) and the data defining the specific treatment being checked.
AAPM Task Group 64 Recommendations The recommendations of the AAPM Task Group 64 regarding treatment-planning systems are that the medical physicist shall verify that the treatment planning system reproduces the values published in TG43 (et seq.) for single source dose at total decay for an 125I model 6711 or a 103Pd model 200 seed in the point source approximation (Yu et al., 1999). This serves to verify that the dose planning system complies with the dosimetric formalism of TG-43 (Nath et al., 1995). This corresponds to the recommendations of Task Group 56 for initial acceptance testing where the results of a planning system are compared to manual calculations. The Task Group 64 report further recommends that the medical physicist shall verify that the system similarly provides valid results for simple configurations of multiple seeds. This echoes the recommendations of the Task Group 56 report. Task Group 64 also recommends QA for imaging sources, equipment, implant templates, applicators, and accessories, and physical dosimeters. Well-chambers shall be ADCL calibrated on a schedule and
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Table 3. Brachytherapy Computer Planning System Quality Assurance Function
Benchmark Data
Frequency
Verify geometric accuracy of I/O peripherals: digitizer, CT or ultrasound interface, and plotter.
Digitize/plot pattern of known geometry; for CT/US, image and reconstruct phantom implant
Monthly
Verify input parameters for all pre-calculated single source arrays
Published recommendations, source vendor’s mechanical drawings
Initially and annually
Verify dose, dwell time, and treatment time calculations at representative points for all source files
Published dose-rate tables; manual calculations
Initially, annually, new software version or source identity
Accuracy of single source isodoses
Point source output
Initially, new software version
Accuracy of multiple-source isodose contouring
Point source data for symmetric source arrays
Initially, new software version
Accuracy of plan rotation matrix
Constancy of point doses, source positions, and isodoses under repeated orthogonal rotations for symmetric source arrays
Initially, new software version
Consistency of printed plan documentation
Assumed input parameters
Every clinical use
Accuracy of coordinate reconstruction
Radiograph phantom with known catheter geometry (added this writing: phantom with known source position geometry)
Initially, new software version
Accuracy of electronic downloading of treatment parameters of afterloader
Comparison of treatment unit and planning system printed output
Initially, new software version, each treatment
Dose-volume histogram/implant figures of merit
* Use isotropic point source or segment of line source allowing analytic calculation of DVH
* Initially, new software version
* Constancy of test case DVH
* Annually
Optimization software
Run series of test cases based in idealized implant geometries of various sizes; Develop a sense of what optimization does tot an implant compared to uniform loading before trying it on patients
Initially, spot check when software changes by duplicating old cases
Overall system test
Run series of standardized plans to globally test all clinically used features
Initially, new software version, annually
(Reproduced from Table VIII in the Task Group 56 report: Med Phys 24:1557–1598, “Code of practice for brachytherapy physics: Report of the AAPM Radiation Therapy Committee Task Group No. 56,” Nath, R., L. L. Anderson, J. A. Meli, A. J. Olch, J. A. Stitt, and J. F. Williamson. (1997).” Also available as AAPM Report No. 59. © 1997, with permission from AAPM.)
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constancy checks using a long-lived radionuclide (preferably with energy comparable to the clinical source) should be performed prior to each use. Ion chambers, Geiger-Müller counters, and survey dosimeters should be calibrated on a schedule. Phantom-based verification of ultrasound imaging systems should be performed to ensure correct alignment of the physical template and the software generated template grid display. Fluoroscopy and radiographic units should be verified for linearity and minimized distortion on display. In particular, ultrasound systems QA should follow those laid out in the AAPM TG-1 report (Goodsitt et al., 1998).
Implementations and Factors Recommendations for Data Entry in Varian Planning Systems VariSeedTM 7.1, is the most recent version of this treatment planning system (Varian Medical Systems, Charlottesville, VA) available for clinical applications. This treatment-planning system utilizes 3-D dosimetry technique to generate dose distributions around a brachytherapy source using either the line or point source approximation. VariSeed accepts all the TG-43 dosimetric parameters (dose rate constant, radial dose function, 2-D anisotropy functions, 1-D anisotropy function, and anisotropy constant) for a particular source in a single library. This planning system only accepts one set of radial dose functions for point and linear source models, rather than two separate data, as recommended in TG-43U1. Therefore, two separate source files must be generated for each source, one for point source and one for linear source calculations. It should be noted that, as per TG-43U1 recommendations, one could use linear radial dose function for point source calculation, as long as the proper geometric function is used. At this stage, the VariSeed planning system doses not have this capability. EclipseTM v7.2.24 (Vision v6.5) is a general-purpose treatment planning system (Varian Medical Systems, Charlottesville, VA) that includes brachytherapy planning in its BrachyVisionTM module. Source data entry and considerations in Eclipse/BrachyVision mimic those for the VariSeed product except that the radial dose function is entered only by 5th order polynomial expansion.
Recommendations for Data Entry in Prowess Planning system There are three different versions of the Prowess planning systems (1370 Ridgewood Dr., Ste. 20 Chico, California) commonly in the United States. The Prowess 2D-planning system is a very user-friendly treatment-planning system for different types of implants using both points and linear sources. In this version of the treatment-planning systems, the source data can be entered for seeds or linear sources. The source parameters in this program can portion of this program required the geometric structures of the source, Γ values, exposure-to-dose conversion factors, and tissue-to-air ratio in the form of the 3rd order polynomial fit coefficients. For a linear source approximation, the program generates a 2-D matrix of dose distributions, using the Sievert Integral method. Although, the program has the option of the TG-43 formalism, this option is only used by replacing the gamma factors by the dose rate constant, placing f-factor=1, and including the anisotropy factors in the calculations. The attenuation coefficients for the source capsule and core must be introduced for linear source approximation. The tabulated data are not editable for the 2-D anisotropy of the radiation distribution around the source. Moreover, the tissue attenuation coefficients of the sources can be only introduced as 3rd order polynomial fits, rather than 5th order polynomial fit as suggested by the TG-43 recommendation. Recently the 3D-Version of the Prowess planning system has become available but provides only for point source approximation. The implementation for linear sources remains in development.
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Recommendations for Data Entry in the ADAC Pinnacle Planning System The Pinnacle3 treatment planning system by Philips Medical Systems is a general-purpose radiation therapy-planning package that includes brachytherapy planning. Source data entry is by one of several methods that support legacy source data as well as TG-43U1 point and line source approximations for more modern sources. Once source data are entered, interconversion can be made between TG-43 parameters and other representations of dose rate about a generalized source. TG-43 radial dose functions may be entered in tabular form or via (up to) 5th order polynomial. The 1-D and 2-D anisotropy function data is entered in tabular TG43U1 form with a variable mesh size. Source strength in a wide variety of units and specific to the clinic inventory may also be entered. Once all data are entered and commissioned, an along-away table for each source type is generated.
Recommendations for Data Entry in Nucletron TheraplanTM and SPOTTM Brachytherapy Planning Systems The TheraplanTM software is a general-purpose RTP system that includes modules for brachytherapy planning. Sources are entered (“modeled” in Theraplan vernacular) by one of several methods providing back-compatibility with legacy source data and providing TG-43 point and line source(s). Once source data is entered, inter-conversion can be made between TG-43 style and other representations of dose-rate about a generalized source. Notably, Theraplan supports both the point and line source approximations of TG43U1 as separate sources. TG-43 radial dose functions may be entered in tabular form or via the Meisberger (or other) 5th order polynomial. Anisotropy data are entered in tabular TG43U1 form and the (outdated) anisotropy constant is supported for point source approximation modeled. The SPOT ProTM system is a special purpose intra-OP planning system for prostate implant planning available from Nucletron Corporation (Columbia, MD). Typical sources are 125I and 103Pd seeds and SPOT Pro takes only TG43U1 formatted data. The TG43U1 recommendation for having separate and different g(r) data for point and line source approximations is supported in SPOT Pro creating two versions of data for a given seed design, one using point source data, the other using line source data. All TG43U1 data are entered in tabular form.
Shortcomings and Recommendations in the Present Planning Systems Linear Source Calculations Despite the vast development of the TG-43 protocol in brachytherapy source dosimetry, the present algorithm is mostly applicable for short brachytherapy sources. Patel et al. (2001) have shown the use of cylindrical coordinates for treatment planning parameters of an elongated 192Ir source. Also they have introduced modified TG-43 algorithm for elongated sources, using the cylindrical coordinate system. Moreover, at the close vicinity of some brachytherapy sources, such as 192Ir, there may be significant contribution of dose from the transmitted electron from the core and the capsule of the source (Baltas et al., 2001; Patel et al., 2001). The present TG-43 algorithm does not cover the dosimetry of such mixed beams. In addition, the tissue heterogeneity corrections have not been adequately addressed in any brachytherapy dosimetry protocols or commercially available planning systems. The 5th order polynomial fit to the radial dose function fails in describing the data at short distances relative to the source axis, particularly for sources with activities distributed only at the two ends of the source. Double exponential method recommended by Furhang and Anderson (1999) and the modified polynomial fit recommended by Meigooni et al. (2003) are more accurate techniques for fitting the data.
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Point Source Calculations Application of point source approximation may be suitable for random orientation of the sources distributed within the implant volume. However, in implants with the fixed orientation of the sources using an applicator (i.e., eye-plaque) or ribbons (i.e., stranded seeds) this method is not suitable. Meanwhile, the limitation of the digitization of the source geometry in an implant volume, does not allow us to utilize the full source dosimetry, as it is needed.
Interpolation and Extrapolations There is no clear recommendation for the methods of extrapolation of the dose distribution, outside of the zone for which dosimetric data are available. There are only limited dosimetric evaluations of the brachytherapy sources, beyond the range that is commonly called “clinical range,” and more work is needed to clarify our techniques of dose extrapolation.
Strategies To Implement Dosimetry Parameters in Systems that Do Not Support TG-43 Format Treatment-planning systems do not always support the latest formulations dosimetry and data. This can nearly always be remedied by combining TG-43 parameters into composite factors and/or rendering into either “along and away” tables or TG-43–like tables. Such tables can be used in treatment-planning systems or for “hand calculation” of dose to selected points. In planning systems that provide point-source only models and that use the 1995 TG-43 anisotropy constant, there are two paths to use the TG43U1 (or equivalent) anisotropy factors, φan(r). The first is to combine this factor with the point source radial dose function, gP(r), to create an anisotropy weighted function, gP(r) φan(r), that one enters in the array provided for the radial dose function (Williamson 2002). When the planning system only supports polynomial expansions of the radial dose function (or gP(r) φan(r) ), then the g(r) function must be expanded into tabular form, the product (with φan(r)) calculated, and the resulting composite function refit to a polynomial of appropriate degree for use in the planning system. In this case, the anisotropy constant entered into the planning system would be unity. The second method is to calculate an anisotropy constant, φan, from the TG43U1 (or equivalent) anisotropy factor data as the inverse-square (of distance) weighted average of φan(r). Formulation and discussion of this option are discussed in greater detail in appendix D of TG-43U1.
Nomogram In general, a nomogram is a novel method of graphic correlation between the parameters in the field of radiation therapy for a quick determination of quantities such as treatment time. Early nomograms were developed to assist in planning 125I prostate brachytherapy (Anderson 1976). Anderson has further expanded the application of this methodology in brachytherapy for determination of the number of 192Ir ribbons and seeds, used in planar implant (Anderson 1984; Anderson, Hilaris, and Wagner 1985). Figure 8 shows a sample nomogram by Lowell Anderson for LDR 192Ir sources. Such nomograms exist for several different radionuclides, particularly for 125I and 103Pd brachytherapy sources (Anderson 1976). In commonly used clinical application, the nomogram, particularly in prostate implant with 125I and 103Pd sources, is one of the most common treatment aids. With an appropriate nomogram, one can readily find the number and strength of seeds to order for patient treatment within clinically acceptable accuracy. One also may find recommendations regarding source spacing (Anderson 1976).
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Figure 8. A sample nomogram by Lowell Anderson for planar implant with 192Ir ribbons with peripheral dose rate of 10 Gy d–1. (Reproduced from Endocuriether/Hypertherm Oncol, vol 1, “A nomograph for planar implant planning,” L. L. Anderson, B. S. Hilaris, and L. K. Wagner, pp. 9–15. © 1985, with permission of the American Institute of Physics.)
These nomograms consist of multiple lines, each scaled for a particular set of implant parameters, such as volume, total and individual seed activities, number of seeds, and seed spacing. These nomograms are designed for a particular application with a particular prescribed dose. For instance, some of the earlier nomograms were designed for prostate implants with 160 Gy prescribed dose, using uniform distribution of the seeds. With some simple steps and calculations, one can determine the number of sources for the low-energy as well as the high-energy brachytherapy in day-to-day clinical applications. Therefore, this calculation method at a given point for a given brachytherapy source provides recommendations based on the accumulated experience of many cases and constitutes of a sample of this experience.
References Altschuler, M. D., P. A. Findlay, and R. D. Epperson. (1983). “Rapid, accurate, three-dimensional location of multiple seeds in implant radiotherapy treatment planning.” Phys Med Biol 28:1305–1318. Altschuler, M. D., and A. Kassaee. (1997). “Automated matching of corresponding seed images of three simulator radiographs to allow 3D triangulation of implanted seeds.” Phys Med Biol 42:293–302. Amols, H. I., and I. I. Rosen. (1981). “A three-film technique for reconstruction of radioactive seed implants.” Med Phys 8:210–214. Anderson, L. L. (1976). “Spacing nomogram for interstitial implants of 125I seeds.” Med Phys 3:48–51. Anderson, L. L. “Computerized Treatment Planning Systems.” F. Bagne (ed). Rockville, MD: US DHHS, 1984. Anderson, L. L., B. S. Hilaris, and L. K. Wagner. (1985). “A nomograph for planar implant planning.” Endocuriether/Hypertherm Oncol 1:9–15. Baltas, D., P. Karaiskos, P. Papagiannis. L. Sakelliou, E. Loeffler, and N. Zamboglou. (2001). “Beta versus gamma dosimetry close to Ir-192 brachytherapy sources.” Med Phys 28 (9): 1875–1882. Bice, W. S. J., D. F. Dubois, J. J. Prete, and B. R. Prestidge. (1999). “Source localization from axial image sets by iterative relaxation of the nearest neighbor criterion.” Med Phys 26:1919–1924. Biggs, P. J., and D. M. Kelly. (1983). “Geometric reconstruction of seed implants using a three-film technique.” Med Phys 10:701–704.
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Blasko, J. C., K. Wallner, P. D. Grimm, and R. Ragde. (1995). “Prostate specific antigen based control following ultrasound-guided 125I implantation for stage T1/T2 prostate carcinoma.” J Urol 154:1096. Brinkmann, D., and R. W. Kline. (1998). “Automated seed localization from CT datasets of the prostate.” Med Phys 25(9):1667–1672. Butler, W. M., and G. S. Merrick. (1996). “I-125 Rapid Strand™ loading technique.” Radiat Oncol Invest 4:48–49. Fraass, B., K. Doppke, M. Hunt, G. Kutcher, G. Starkschall, R. Stern, and J. Van Dyk. (1998). “American Association of Physicists in Medicine Radiation Therapy Committee Task Group 53: Quality assurance for clinical radiotherapy treatment planning.” Med Phys 25:1773–1829. Also available as AAPM Report No. 62. Fuks, Z., S. A. Leibel, K. E. Wallner, C. B. Begg, W. R. Fair, L. L. Anderson, B. S. Hilaris, and W. F. Whitmore. (1991). “The effect of local control on metastatic dissemination in carcinoma of the prostate: long-term results in patients treated with 125I implantation.” Int J Radiat Oncol Biol Phys 21:537. Furhang, E. E., and L. L. Anderson. (1999). “Functional fitting of interstitial brachytherapy dosimetry data recommended by the AAPM Radiation Therapy Committee Task Group 43.” Med Phys 26(2):153–160. Glasser, O., E. H. Quimby, L. S. Taylor, J. L. Weatherwax, and R. H. Morgan. Physical Foundation of Radiology, 3rd ed. New York: Harper & Row, 1961. Goodsitt, M. M., P. J. Carson, S. Witt, D. L. Hykes, and J. M. Kofler. (1998). “Real-time B-mode ultrasound quality control test procedures: Report of AAPM Ultrasound Task Group No. 1.” Med Phys 25:1385–1406. Hallert, B. X-Ray Photogrammetry: Basic Geometry and Quality. New York: Elsevier Publishing Company, 1970. Holupka, E. J., P. M. Meskell, E. C. Burdette, and I. D. Kaplan. (2004). “An automatic seed finder for brachytherapy CT postplans based on the Hough transform.” Med Phys 31(9):2672–2679. Jackson, D. D. (1983). “An automatic method for localizing radioactive seeds in implant dosimetry.” Med Phys 10:370–372. Khan, F. The Physics of Radiation Therapy. Baltimore, MD: Williams and Wilkins, 2003. Krishnaswamy, V. (1972). “Dose distribution about 137Cs sources in tissue.” Radiology 105:181–184. Kutcher, G. J., L. Coia, M. Gillin, W. F. Hanson, S. Leibel, R. J. Morton, J. R. Palta, J. A. Purdy, L. R. Reinstein, G. K. Svensson, M. Weller, and L. Wingfield. (1994). “Comprehensive QA for radiation oncology: Report of the AAPM Radiation Therapy Committee Task Group No. 40.” Med Phys 21:581–615. Also available as AAPM Report No. 46. Kumar, P. P., and R. R. Good. (1986). “Vicryl carrier for I-125 seeds: Percutanious transperineal insertion.” Radiology 159:276. Li, Z., I. A. Nalcacioglu, S. Ranka, S. K. Sahni, J. R. Palta, W. Tome, and S. Kim. (2001). “An algorithm for automatic, computed-tomography-based source localization after prostate implant.” Med Phys 28(7):1410–1415. Liu, H., G. Cheng, Y. Yu, R. Brasacchio, D. Rubens, J. Strang, L. Liao, and E. Messing. (2003). “Automatic localization of implanted seeds from postimplant CT images.” Phys Med Biol 48:1191–1203. Marinello, G., and B. Pierquin. “The Paris Systems, Optimization of Dose, and Calculation of Treatment Time.” Chapter 4 of A Practical Manual of Brachytherapy. B. Pierquin and G. Marinello. Translated by Frank Wilson, Beth Erickson, and Jack Cunningham. Madison, WI: Medical Physics Publishing, 1997. Meigooni, A. S., H. Zhang, C. Perry, S. A. Dini, and R. A. Koona. (2003). “Theoretical and experimental determination of dosimetric characteristics for BrachySeed Pd-103 model Pd-1 source.” Appl Radiat Isot 58:533–541. Meigooni, A., C. B. Saw, and R. Nath. “Basic Physics of Brachytherapy” in Principles and Practice of Brachytherapy. Subir Nag (ed.). Armonk, NY: Futura Publishing Company, Inc., 1997. Meigooni, A. S., V. Rachabatthula, S. B. Awan, and R. A. Koona. (2004). “Treatment planning considerations for prostate implants with the new linear RadioCoil™ 103Pd brachytherapy source.” To be submitted to Medical Physics. Meisberger, L. L., R. J. Keller, and R. J. Shalek. (1968) “The effective attenuation in water of the gamma rays of gold 198, iridium 192, cesium 137, radium 226, and cobalt 60.” Radiology 90:953–957. Meredith, W. J. (ed.). Radium Dosage: The Manchester System. Edinburgh: Livingston, 1967. Merrick, G. S., W. M. Butler, A. T. Dorsey, and H. L. Walbert. (1998). “The influence of timing on the dosimetric analysis of transperineal ultrasound guided prostatic conformal brachytherapy.” Radiat Oncol Invest 6:182. Merrick, G. S., W. M. Butler, A. T. Dorsey, J. H. Lief, and M. L. Benson. (2000). “Seed fixity in the prostate/periprostatic region following brachytherapy.” Int J Radiat Oncol Biol Phys 46(1):215–220.
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Nath, R., L. L. Anderson, G. Luxton, K. A. Weaver, J. F. Williamson, and A. S. Meigooni. (1995). “Dosimetry of interstitial brachytherapy sources: Recommendations of the AAPM Radiation Therapy Committee, Task Group No. 43.” Med Phys 22:209–234. Also available as AAPM Report No. 51. Nath, R., L. L. Anderson, J. A. Meli, A. J. Olch, J. A. Stitt, and J. F. Williamson. (1997). “Code of practice for brachytherapy physics: Report of the AAPM Radiation Therapy Committee Task Group No. 56.” Med Phys 24(10):1557–1598. Also available as AAPM Report No. 59. NEMA 2004. “The DICOM Standard” online at http://medical.nema.org/dicom/2004.html, National Electrical Manufacturers Association, Rosslyn, VA. Nori, D. (1992). “Conformal brachytherapy of prostate cancer: An effective outpatient treatment.” Cancer Watch 1:124). Patel, N. S., S.-T. Chiu-Tsao, P. Fan, H. S. Tsao, S. F. Liprie, and L. B. Harrison. (2001). “The use of cylindrical coordinates for treatment planning parameters of an elongated 192Ir source.” Int J Radiat Oncol Biol Phys 51(4):1093–1102. Pierquin, B., and G. Marinello. A Practical Manual of Brachytherapy. Translated by Frank Wilson, Beth Erickson, and Jack Cunningham. Madison, WI: Medical Physics Publishing, 1997. Porrazzo, M. S., B. S. Hilaris, C. R. Moorthy, A. E. Tchelebi, C. A. Mastoras, L. L. Shih, L. Stabile, and N. Salvaras. (1992). “Permanent interstitial implantation using palladium-103: The New York Medical College preliminary experience.” Int J Radiat Oncol Biol Phys 23:1033. Rivard, M. J., B. M. Coursey, L. A. DeWerd, W. F. Hanson, M. S. Huq, G. S. Ibbott, M. G. Mitch, R. Nath, and J. F. Williamson. (2004). “Update of AAPM Task Group No. 43 Report: A revised AAPM protocol for brachytherapy dose calculations.” Med Phys 31:633–674. Also available as AAPM Report No. 84. Rosenthal, M. S., and R. Nath. (1983). “An automatic seed identification technique for interstitial implants using three isocentric radiographs.” Med Phys 10(4):475–479. Rosenwald, J. C. “Calculation of Dose Distribution,” Chapter 3 of A Practical Manual of Brachytherapy. B. Pierquin and G. Marinello. Translated by Frank Wilson, Beth Erickson, and Jack Cunningham. Madison, WI: Medical Physics Publishing, 1997. Roy, J. N., K. E. Wallner, P. J. Harrington, C. C. Ling, and L. L. Anderson. (1993). “A CT-based evaluation method for permanent implants: application to prostate.” Int J Radiat Oncol Biol Phys 26(1):163–169. Russel, K. J., and J. C. Blasko. (1993). “Recent advances in interstitial brachytherapy for localized prostate cancer.” Probl Urol 7:260. Schlienger, M., J. C. Rosenwald, M. Miclutia, R. Quint, and B. Pierquin. (1970). “Controle dosimetrique en brachycuritherapie par les isodoses,’escargot’.” Acta Radiol 9:282—288. Shalek, R. J., and M. Stovall. “Dosimetry in Implant Therapy” in Radiation Dosimetry, vol III. F. H. Attix and E. Tochlin (eds.). New York: Academic Press, pp. 776–798, 1969. Sievert, R. M. (1921). “Die Intensitätsverteilung der primären g-Strahlung in der Nähe medizinischer Radiumpräparate.” Acta Radiol 1:135. Sommerkamp, H., M. Ruppercht, and M. Wannenmacher. (1988). “Seed Loss in Interstitial radiotherapy of prostatic carcinoma with I-125.” Int J Radiat Oncol Biol Phys 14:389. Su, Y., B. J. Davis, M. G. Herman, and R. A. Robb. (2004). “Prostate brachytherapy seed localization by analysis of multiple projections: Identifying and addressing the seed overlap problem.” Med Phys 31(5):1277–1287. Tapen, E. M., J. C. Blasko, P. D. Grimm, H. Radge, L. Ray, S. Cliford, J. Sylvester, and T. Griffin. (1998). “Reduction of radioactive seed embolization to lung following prostate brachytherapy.” Int J Radiat Oncol Biol Phys 42:1063. Todor, D. A., G. N. Cohen, H. I. Amols, and M. Zaider. (2002). “Operator-free, film based 3D seed reconstruction in brachytherapy.” Phys Med Biol 47:2031–2048. Tubic, D., A. Zaccarin, J. Pouliot, and L. Beaulieu. (2001b). “Automated seed detection and three-dimensional reconstruction. II. Reconstruction of permanent prostate implants using simulated annealing.” Med Phys 28:2272–2279. Tubic, D., and L. Beaulieu. (2005). “Sliding slice: A novel approach for high accuracy and automatic 3D localization of seeds from CT scans.” Med Phys 32:163–174. Tutar, I. B., R. Managuli, and V. Shamdasani. (2003). “Tomothynthesis-based localization of radioactive seeds in prostate brachytherapy.” Med Phys 30:3135–3142.
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Williamson, J. F. “Dosimetric characteristics of the DRAXIMAGE model LS-1 I-125 interstitial brachytherapy source design: A Monte Carlo investigation.” Med Phys 29:509–521. Williamson, J. F., R. I. Morin, and F. M. Khan. (1983). “Monte Carlo evaluation of the Sievert integral for brachytherapy dosimetry.” Phys Med Biol 28:1021. Willins, J., and K. Wallner. (1998). “Time-dependent changes in CT-based dosimetry for I-125 prostate brachytherapy.” Radiat Oncol Invest 6:157–160. Yu, Y., L. L. Anderson, Z. Li, D. E. Mellenberg, R. Nath, M. C. Schell, F. M. Waterman, and A. Wu. (1999). “Permanent prostate seed implant brachytherapy: Report of the American Association of Physicists in Medicine Task Group No. 64.” Med Phys 26:2054–2076. Also available as AAPm Report No. 68.
Chapter 18
Systems 1B Manchester Planar and Volume Implants and the Paris System Michael T. Gillin, Ph.D., and Firas Mourtada, Ph.D. The University of Texas M.D. Anderson Cancer Center Department of Radiation Physics Houston, Texas Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 351 The Manchester System for Surface Applicators . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 352 The Manchester Interstitial System . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 354 The Manchester Planar System . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 354 The Manchester Volume System . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 360 Paris System of Single and Double Plane Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 363 Optimized Stepping Source Dosimetry System . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 369 Actual Practice . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 370 Summary . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 371 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 371
Introduction The development of systems for interstitial brachytherapy began early in the practice of radiation oncology. It is now a challenge for the contemporary radiation oncology physicist to comprehend the practice of radiation oncology before the current digital universe. Different priorities existed in the pre-digital era. Some of these priorities addressed the need to consistently produce clinically acceptable dose patterns in the absence of both visual dose distributions and significant medical physics support. Madame Curie prepared the first brachytherapy calibration source of radium salt. The unit of milligrams of 226Ra became a measure of source strength for intercomparison of brachytherapy sources. Sievert developed a calculation approach in 1921 for filtered sources. In 1932 Quimby published the first article on the system, which now bears her name. The Quimby system is based upon a uniform distribution of sources to produce a non-uniform radiation distribution. In 1934 the first paper, which described the Manchester System, also commonly known as the Paterson and Parker system, was published. In the Manchester System, the sources are arranged in a non-uniform pattern in order to produce a “uniform” dose distribution. In 1960 Pierquin published the first paper on the Paris System, which was based on the use of 192Ir. In some ways the Paris System, with its variable source spacing, can be viewed as an evolution from the Manchester System with its fixed source spacing. These systems were developed to serve as a guide to the oncologist. If the sources were arranged according to the rules of the system, then the oncologist had a sense of the resultant dose distribution and could provide consistency in patient care. Image-based computerized dose distributions developed later. These systems still have a role in the thought process, which should precede every implant. Given a specific volume to be treated, the systems suggest the pattern of the source placement, including the separation between the sources and the source length. The initial sources used in the Manchester System were radium (226Ra) sources. These sources had both a physical length and a shorter active length, which is addressed by the Manchester System. The classic Manchester dose prescription, translated to the modern concept of dose, was 65 Gy in 6 to 8 days. This lead to full strength 226Ra needles of 0.66 mg cm–1 and half strength 226Ra needles of 0.33 mg cm–1. Eventually 137Cs needles were available as a replacement for 226Ra needles. These needles were also characterized by the physical and active length. Many institutions possessed inventories of various needles that had different lengths and source strengths. There were needles with a linear activity distribution,
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needles with a high activity at one end in the event that one end could not be crossed (Indian club needles), and needles with high activity on both ends (dumbbell needles). Today, less descriptive but more politically correct names would most likely be used. Modern basic physics references for interstitial brachytherapy include the chapter by Shalek and Stovall, “Brachytherapy Dosimetry,” in volume III of The Dosimetry of Ionizing Radiation (Shalek and Stovall 1990) and Physical Aspects of Brachytherapy: Medical Physics Handbook 19 by Godden (1988). In addition, there is chapter 13 in The Physics of Radiology, 4th Edition by Johns and Cunningham (1983). There are multiple clinical/physics texts, which are devoted in whole or in part to brachytherapy. Chapter 12, “Physics of Brachytherapy” by Glasgow and Perez, in Principles and Practice of Radiation Oncology, Second Edition, by Perez and Brady is an outstanding summary of brachytherapy in a clinical textbook (Perez and Brady 1992). A Practical Manual of Brachytherapy by Pierquin and Marinello has been translated into English by Wilson, Erickson, and Cunningham (Pierquin and Marinello 1997) and is the most complete English language description of the Paris System. There are some classic books from the 1930s and 1940s, which describe the Manchester System, e.g., Radium Dosage, The Manchester System, edited by W. J. Meredith (Paterson et al. 1947). Readers of the older works should remember that some numeric values found in these texts have been updated, as various parameters have been refined. The importance of units does not have to be stressed to physicists. Units are especially important in brachytherapy. It may be necessary to be able to convert from one system of units to another, especially as they pertain to the statement of source activity. The definition of terms is also important. The various systems have specific terms, which are explicitly defined by that system.
The Manchester System for Surface Applicators The Manchester System for surface applicators is based upon the assumption that the source distribution rules are followed. The result is a ±10% dose uniformity over the treatment surface which is located at a distance, h, from the plane of the implant. The original dose tables from the 1930s gave the amount of radium in milligram-hours (mg-h) to give a “dose” of 1000 R to the treatment area, which is at a distance h from the surface of the applicator. Over the years, a number of authors have recommended corrections to convert the original tables to “modern” units. The resultant correction factor, which ranges between 1.064 (Johns and Cunningham) to 1.08 (Godden) and 1.10 (Shalek), was used to multiply the existing tables to create new tables, which provide the amount of milligram-hours to deliver 10 Gy at various treatment distances as a function of the area defined in square centimeters. For example, consider a 16-cm2 area and a treatment distance of 10 mm. The table as modified by Gibb and Massey (1980) shows that 611 mg-h are required to deliver 10 Gy. These tables can be used for other isotopes by substituting mCi-h or MBq-h for mg-h as determined by the ratio of the exposure rate constants. The exposure rate constant, Γ, is defined as the exposure rate in R/hr at a point 1 meter from a 1 Ci source. The exposure rate constant for 226Ra, filtered by 0.5 mm Pt, a classic number in brachytherapy, is 0.825 R m2 hr–1 Ci–1, or 8.25 mR cm2 hr–1 mg–1. The exposure rate constants for different sources can be found in the texts, which have been previously described. Thus, if the implant is to be performed using 192Ir, the 611 mg-h of radium would be converted to: 611 mg-h (radium) × 8.25 (226Ra)/4.69 (192Ir) = 1075 mCi-h (192Ir) , where Ra–226Γδ (8.25) and Ir–192Γδ (4.69) are the exposure rate constants in units of R cm2/mCi-h.
(1)
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If the prescribed dose of 10 Gy is to be delivered over 24 hours, then the required activity is: 611 mg-h/24 h
=
25.5 mg 226Ra
or
(2) 1075 mCi-h/24 h
=
44.8 mCi 192Ir .
It is also possible to correct the tables for rapidly decaying isotopes, such as 222Ra or 198Au. Paterson and Parker first described a dosimetry system for surface applicators. They then extended their work to propose a system for interstitial applications. There are some differences between these two systems. However, the basic approach is the same in that source distribution rules are defined and tables are presented that provide the product of the activity and treatment time as a function of area and distance to deliver a specific amount of dose. The Manchester System permits the sources to be arranged in a crossing pattern, i.e., the sources can be perpendicular to each other as they define the plane of interest. The source distribution rules for surface applicators, as presented by Godden (1988), which are based upon chapter II, “Mould Treatments” in the Paterson et al. (1947) book for rectangular applicators (a × b) and a treatment distance h, are as follows: 1. The distance between the active source ends should not be greater than 2h. 2. For rectangles and/or squares of area a × b, where b > a, sources should be placed around the periphery and possibly some in lines parallel to b, the number of lines being such that the area is divided into strips of width 2h. 3. If the activity per unit length on the periphery is ρ, then for a 4h two or more lines are added whose linear density is 2/3 ρ. 4. The number of mg-h for 10 Gy must be increased by a correction factor to account for any elongation factor, i.e., the ratio of b to a. This elongation correction factor ranges from 1.025 for a b/a ratio of 1.5 to 1.12 for a ratio of 4 to 1. Consider a simple example for a rectangular applicator of a 2 cm × 3 cm applicator that is designed to treat to a depth of 5 mm. Thus twice the depth, 2h, is 10 mm, so that the sources would be placed 10 mm apart. The applicator is designed such that there are needles around the entire periphery (no uncrossed ends). For 6 cm2 at a treating distance of 5 mm, Table 5.1 in Godden (1988) gives 191 mg-h per 10 Gy, while Table 13-4 of Johns, which is Table 2 in this chapter, gives 188 mg-h per 10 Gy. If the prescribed dose is 60 Gy, then the number of milligram-hours required is: 191 mg-h/10 Gy × 60 Gy × 1.025 = 1175 mg-h .
(3)
If the patient will be asked to wear this applicator for 7 hours per day for a total of 5 days, then the number of milligrams is: 1175 mg-h/(7 hours/day × 5 days) = 33.6 mg .
(4)
Assume that the applicator is constructed with four needles (2 cm active length each), which are separated by 10 mm, and two needles (3 cm active length each), which are crossing (see Figure 1). The total
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Figure 1. Representation of an applicator constructed with four needles (separated by 10 mm) and two needles crossing.
active length is 14 cm. The total length of needles on the periphery is 10 cm and 4 cm of total needle length in the center. In this example, the shorter distance is exactly equal to 4h and the applicator is defined with two central sources. Following the rules for applicators, the linear intensity can be calculated as follows: (3 + 3 + 2 + 2)ρ + (2/3)ρ(2 + 2) = 12.67 ρ ,
(5)
where ρ = 2.65 mg/cm for 226Ra or 2.65 (8.25/4.69) 4.66 mCi (192Ir)/cm. Thus assuming the use of 192Ir sources, the four peripheral sources would have a linear activity of 4.67 mCi/cm and the two central sources would have a linear activity of 3.13 mCi/cm.
The Manchester Interstitial System In the interstitial system, the “Stated Dose” is defined as the dose, which is 10% greater than the minimum dose in the plane of the calculation. The maximum dose in the volume being treated will be substantially higher than the minimum dose as the point of calculation approaches a source. The Manchester System provides guidance for both planar and volume implants.
The Manchester Planar System For single plane implants, Manchester/Paterson and Parker (1947) considered the uniformity of the dose in a plane, which is located 5 mm from the source plane and is parallel to the source plane. Thus a 10mm thick volume of tissue with the sources located in the center of this volume is treated with a single plane implant. The area of the implant is defined as the width of the implant times the distance between the crossing needles, assuming that there are crossing needles on both ends of the implant. The area to be used in the Manchester System tables should be reduced by 10% for each uncrossed end. Figure 2 presents a drawing of the areas involved for implants with crossing needles, implants with one end crossed, and implants with no crossing sources. The Manchester System permits double plane implants of up to 2.5 cm thickness. Both Godden (1988, Table 6.1) and Johns and Cunningham (1983, p. 474) present separation factors for separation distances of 1.5 cm, 2.0 cm, and 2.5 cm. Johns suggests that these factors are only approximate and provides an alternative approach by calculating the doses in planes which are 5, 10, and 15 mm from the source planes. Chapter V, “Interstitial Treatments” of Paterson et al. (1947) presents the Manchester rules for planar and volume implants. Johns and Cunningham (1983) Table 13-5 is surprisingly weak on the Manchester rules, while Shalek presents a concise statement of the rules (Shalek and Stovall 1990). The following is a modern version of these rules as presented by Godden (1988):
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1. Sources should be placed in a single plane with a certain amount of the activity distributed on the periphery and the remaining amount distributed evenly in the center. The relative distribution depends upon the area, as shown below. 2. The needles should be in parallel rows at a spacing not greater than 1.0 cm. 3. Crossing needles are permitted and they should be placed across the active ends if possible, but not more than 1 cm from the active ends. 4. If the ends of the implant cannot be across, 10% of the area is deducted from each uncrossed end. 5. For implants, which use seeds or short sources, the distance between the active ends should not exceed 1 cm. 6. If two planes are used, the separate planes should be arranged as for a single implant, parallel to each other. If they differ in area, then the area used to determine the amount of activity is the average area and the activity is divided between the planes in a ratio of the areas of the planes. In addition, there are rules for the distribution of activity, which have been memorized by oncologists and physicists for decades (see Table 1). While the exact ratios may not be obvious, the fact that more activity is placed in the central part of the implant as the area being implanted increases seems intuitive. Consider a 2 cm × 3 cm interstitial implant (as opposed to the applicator, which was described above; see Figure 3). Assuming two crossed ends, one would use the same number of mg-h per 10 Gy, i.e., 191. The reason for this is that for a single plane implant, the prescribed dose is in a plane 5 mm from the plane of the implant. The implant is scheduled to be in place 7 days or 168 hours to deliver a total dose of 60 Gy. The number of milligrams required is:
Figure 2. Areas involved for implants with crossing needles, implants with one end crossed, and implants with no crossing sources. (Reprinted from Biomedical Uses of Radiation, W. Hendee (ed), Fig. 5.39. © 1999, with permission of Wiley-VCH, Weinheim.)
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Michael T. Gillin and Firas Mourtada Table 1. Manchester Distribution of Activity for Planar Implants Area
Relative Proportion of Activity
Periphery
Center
< 25 cm2
2/3
1/3
> 25 cm2 and < 100 cm2
1/2
1/2
> 100 cm
1/3
2/3
2
161 mg-h/10 Gy × 60 Gy/168 h = 5.75 mg 226Ra or 10.1 mCi 192Ir .
(6)
The distribution rules call for two-thirds of the activity to be placed on the periphery or 6.8 mCi (192Ir) on the periphery and one-third of the activity in the center or 3.3 mCi (192Ir). The linear activity for the peripheral sources is 0.68 mCi/cm and for the central sources is 0.82 mCi/cm. Consider a classic single plane implant using 137Cs needles (physical length 42 mm and active length 30 mm.) As shown in Figure 4, this implant has one uncrossed plane. The inventory of the 137Cs needles contains three 2.00 mg Ra equivalent needles and two 1.00 mg Ra equivalent needles, which are available for use with this implant. The area of the implant is: 3 cm × 3 cm × 0.9 = 8.1 cm2.
(7)
For this area, approximately 221 mg-h/10 Gy is required (see Table 2). Two-thirds of the activity should be placed on the periphery and thus the three 2.00 mg Ra eq needles are used on the periphery. One third of the activity should be centrally located and thus the two 1.00 mg Ra equivalent needles are used in the center of this implant. The dose rate is: (8) 8 mg ⋅ 1000 cGy 221 mg-h
= 36.2 cGy/h.
Figure 3. 2 cm × 3 cm interstitial implant representation.
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Table 2. Surface Applicators and Planar Implants The table gives RA, the number of milligram hours required to deliver 10 Gy to muscle tissue for different areas and treatment distances. Filtration 0.5 mm Pt. The table may be used for planar implants by using a treatment distance of 0.5 cm. Treatment Distance (cm) 2
Area cm 0 1 2 3 4 5 6 7 8 9 10 12 14 16 18 20 22 24 26 28 30 32 34 36 38 40 42 44 46 48 50 60 70 80
.5
1.0
1.5
2.0
2.5
3.0
3.5
4.0
4.5
5.0
32 72 103 128 150 170 188 204 219 235 250 278 306 335 364 392 418 444 470 496 521 546 571 594 618 642 664 685 708 729 750 851 947 1044
127 182 227 263 296 326 354 382 409 434 461 511 557 602 644 682 717 752 784 816 846 876 909 935 967 994 1024 1053 1080 1110 1141 1283 1426 1567
285 343 399 448 492 531 570 603 637 667 697 755 813 866 918 968 1021 1072 1122 1170 1215 1261 1305 1349 1392 1432 1472 1511 1550 1585 1619 1790 1944 2092
506 571 632 689 743 787 832 870 910 946 982 1053 1120 1184 1245 1303 1362 1420 1477 1530 1582 1635 1688 1743 1793 1843 1894 1942 1990 2037 2083 2319 2532 2726
792 856 920 978 1032 1083 1134 1182 1229 1272 1314 1396 1475 1553 1622 1690 1755 1821 1881 1943 2000 2060 2119 2179 2234 2290 2344 2399 2452 2504 2556 2815 3059 3301
1139 1204 1274 1331 1388 1436 1495 1547 1596 1645 1692 1780 1865 1947 2027 2106 2180 2252 2328 2398 2468 2532 2598 2662 2726 2787 2848 2908 2966 3025 3082 3362 3628 3891
1551 1625 1697 1760 1823 1881 1938 1993 2047 2099 2149 2247 2341 2429 2514 2601 2683 2764 2841 2917 2997 3073 3145 3215 3285 3351 3421 3484 3548 3612 3676 3974 4257 4532
2026 2100 2172 2241 2307 2369 2432 2490 2548 2605 2660 2769 2870 2968 3063 3155 3242 3326 3405 3484 3562 3639 3713 3787 3859 3931 4003 4071 4139 4207 4275 4605 4913 5213
2566 2636 2708 2772 2835 2896 2956 3011 3067 3123 3178 3284 3389 3490 3585 3682 3777 3872 3962 4047 4131 4220 4306 4389 4466 4546 4626 4706 4781 4857 4929 5288 5632 5958
3166 3295 3349 3383 3450 3513 3575 3634 3694 3752 3809 3917 4027 4131 4240 4341 4441 4540 4634 4730 4824 4915 5000 5089 5174 5258 5341 5422 5505 5586 5668 6054 6419 6756
Filtration (mm Pt) Correction to mg hrs
0.3 –4%
0.5 0
0.6 +2%
0.8 +6%
1.0 +10%
This table was prepared from the original by Meredith (M12) by multiplying his values by C=1.064. (Reprinted from Johns, H.E., and J. R. Cunningham, The Physics of Radiology, 4th Edition, ©1983. Courtesy of Charles C Thomas Publisher, Ltd., Springfield, IL.)
1.5 +20%
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If the total dose to be delivered is 7000 cGy, then the total time of the implant is: 7000 cGy/(36.2 cGy/h) = 193.4 hours.
(9)
Plan a simple two-plane implant with a separation of 2.0 cm between the planes (see Figure 5). The separation factor is 1.4. The planned dose is 70 Gy to be delivered in 168 hours. Both planes have the same area, i.e., 4 cm × 3 cm. There are no crossing sources. The area to be used in the Manchester tables is: 4 cm × 3 cm × 0.8 = 9.6 cm2.
(10)
Interpolating the Manchester tables for this area yields 244 mg-h/10 Gy. This factor now needs to be adjusted by the separation factor. 244 mg-h ⋅ 1.4 10 Gy
=
342 mg-h
.
10 Gy
(11)
Note that the number of milligram hours has been increased as a result of the separation factor. 244 mg-h ⋅ 1.4 10 Gy
=
342 mg-h
.
10 Gy
Figure 4. Single plane implant with one crossing source.
(12)
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359
Each plane, which contains five sources, should contain a total 7.15 mg. The two peripheral sources contain two-thirds of this activity, i.e., 2.8 mg per source. The three central sources should contain one third of this activity, i.e., 0.79 mg per source. The next step in this planning exercise is to consult the inventory of 137Cs needles that are available to be used. Assume that the inventory contains 3-cm active length needles with activities of 1.0 mg Ra eq and 3.0 mg Ra eq. Thus, due to limitations in the inventory of sources, the total activity per plane is 9 mg. This still results in a source pattern, which obeys the Manchester rules, in that two-thirds of the activity is still on the periphery. If 192Ir was used for this implant, as opposed to 137Cs or 226Ra, approximately 25 mCi would be required using a conversion factor based on the ratio of the gamma rate constants (or the exposure rate constants, as has been previously demonstrated.) It is now possible to calculate the dose rate in various planes, using the Manchester surface applicators and planar implant table. 5 mm from a source plane:
10 mm from a source plane:
15 mm from a source plane:
10 Gy ⋅ 9 mg 244 mg h 10 Gy ⋅ 9 mg 450 mg-h 10 Gy ⋅ 9 mg 685 mg-h
=36.9 cGy/h
=20.0 cGy/h
(13)
=13.1cGy/h
The total dose rate in a plane, which is 10 mm from each source plane, is 40 cGy/h. The total dose rate in a plane, which is 5 mm from one source plane and 15 mm from the other source plane, is 50 cGy/h. The time required to deliver 70 Gy is 175 hours (7000 cGy/40 cGy/h).
Figure 5. Double plane implant.
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The Manchester Volume System In the Manchester Volume System, the “Stated Dose” is taken to be the dose, which is 10% higher than the absolute minimum dose in the effective volume. Consider a line radiating from a spherical or cylindrical volume, which passes between the sources. If the Manchester source distribution rules are followed, then the minimum dose at the surface of the volume is 10% below the stated dose. The maximum dose on the line increases as the point of interest moves toward the center of the volume and may be as much as 20% above the stated dose. Volumes are considered to be composed of a rind, which covers the entire surface, like a skin, and a central core. For a sphere, the rind is the surface of the sphere, while for a cylinder, the rind is a belt, the curved surface of the cylinder, plus the two flat ends. The total activity is divided into a number of equal parts and distributed according to the rules, which are presented below. This is summarized in Johns and Cunningham (1983, Table 13-8) and in Godden (1988, Table 6.5). See Table 3. Chapter V of Paterson et al. (1947) also contains the rules for volume implants. The following rules for volume implants are taken from Godden (1988): 1. The sources on each face must be spaced as evenly as possible. 2. The sources must be distributed throughout the core. 3. The sources on each surface must be spaced evenly and not more than 1.0 to 1.5 cm apart. 4. For a cylindrical implant, there must not be less than eight sources in the belt and four in the core. 5. For cuboid implants arranged as a multiplane implant, the distribution of sources is as for planar implants. 6. The geometric volume of the implant must be reduced by 7.5% for each uncrossed end. The geometric volume is defined as the cross-sectional area multiplied by the active length, which is the active length of the belt. There are separate tables, which provide the number of mg-h for 10 Gy for volume implants. Modern versions of these tables can be found in Johns and Cunningham (1983, Table 13-8) or in Godden (1988, Table 6.3). See Table 3. Consider planning a volume implant, using the Manchester approach, for a floor of the mouth lesion. For planning purposes, this lesion can be viewed as a square of 2.7 cm × 2.7 cm with a 3.0 cm height (see Figure 6). One approach would be to place needles in a 1 cm matrix with an active length of 3.0 cm. This would result in 12 needles in the belt and 4 needles in the core. The volume of the implant is 21.9 cm3. Based upon an interpolation of Table 13-8 (see Table 3), 284 mg-h/10 Gy would be required. In a 1967 monograph entitled Implant Dosimetry by Shalek and Stovall (1967) describes an implant resembling the above (see Figure 7). In this actual clinical example, 15 needles were used in total. Twelve of the needles had an active length of 3.5 cm and were of Indian club design with a total activity of 1.5 mg each. Two needles had an active length of 3.0 cm with a total activity of 1.0 mg each. One needle had an active length of 3.5 cm with a total activity of 1.16 mg. There were crossing needles on one end and the Indian club needles were implanted such that the higher activity portion of the needle was placed deeper into the patient, which essentially represented another crossed end. There were eight needles in the belt and four needles in the core. The total activity implanted was 20.74 mg. The dose rate for this implant was: 20.74 mg/(284 mg-h/10 Gy) = 73 cGy/h .
(14)
Systems 1B Manchester Planar and Volume Implants and the Paris System Table 3. Volume Implants Volume cm3 cm3
Rv mg hr
5 10 15 20 30 40 50 60 80 100 140 180 220 260 300 340 380
106 168 220 267 350 425 493 556 673 782 979 1156 1322 1479 1627 1768 1902
Distribution Rules Volume should be considered as a surface with 75% activity and core with 125% Rules for cylinders Belt—50% activity with minimum 8 needles Ends—12.5% of activity on each end Core—25% with minimum of 4 needles For each uncrossed end, reduce volume by 7.5% Length Diameter
1.5% 2.0
Increase mg hr 3%
6%
2.5
3.0
10%
15%
This table was prepared from the original by Meredith (M12) by multiplying his values by C=1.064. (Reprinted from Johns, H.E., and J. R. Cunningham, The Physics of Radiology, 4th Edition, ©1983. Courtesy of Charles C Thomas Publisher, Ltd., Springfield, IL.)
Figure 6. Theoretical volume implant.
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Michael T. Gillin and Firas Mourtada
An example of a permanent implant is presented in chapter V, “Interstitial Treatments” in Paterson et al. (1947). This old Manchester example used radon (222Ra) seeds that were encapsulated in gold. The following example uses 198Au seeds, half-life of 2.698 days, for the same implant, which was a 4.0 cm diameter sphere, volume 33.5 cm3, designed to deliver 65 Gy. Interpolating the volume Manchester table, Johns (Johns and Cunningham 1983, Table 13-8) yields 376 mg-h/10 Gy. Thus the total number of mg-h needed to deliver 65 Gy is 2444 mg-h. It is now necessary to convert this to the appropriate 198Au activity. Using the classic exposure rate constants (226Ra 0.825 R m2 h–1 Ci–1 and 198Au 0.238 R m2 h–1 Ci–1), it can be determined that: 1 mCi (198Au) = 0.288 mg (226Ra) 1 mCi decayed (198Au) = 1.44 × 2.698 days × 24 hours/day × 0.288 mg
(15)
= 26.9 mg-h (226Ra) . To determine the required activity of 198Au: 2444 mg-h . = 90.9 mCi . mg-h 26.9 mCi
Figure 7. Implant using 15 needles (cf., Shalek and Stovall 1967).
(16)
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363
Following the classic Manchester rules, the rind would contain six parts of the activity (68 mCi) and the core 2 parts (23 mCi) (see Table 4). For planning purposes, the implant would consist of 12 seeds, 5.7 mCi per seed, on the surface of the 4 cm diameter sphere and 6 seeds, 3.8 mCi each, evenly distributed throughout the volume of the sphere.
Paris System of Single and Double Plane Implants The French, led by Pierquin and Chassagne, began proposing a different system of interstitial implants in the early 1960s. This system used 192Ir in the form of wire, straight pins, or double pins. In the 1970s in the United States, it was possible to purchase either wire or pins from Europe as well as 192Ir seeds from various American companies. Today, U.T. M.D. Anderson (MDACC) continues to use 192Ir wire by purchasing the wire from the United Kingdom and activating the wire in a reactor, which is located in Texas. Americans who made major early contributions to the use of 192Ir included both Henschke and Delclos. Customized 192Ir sources, which could be obtained in one or two days, meant that institutions were no longer dependent upon maintaining a large inventory of mock 226Ra sources, such as 137Cs. (In the 1970s, the federal government had a program that encouraged hospitals to dispose of their 226Ra sources.) The Paris System is defined for both single- and double plane implants. The maximum thickness allowed for a single plane implant is 12 mm and the maximum thickness for a double plane implant is 25 mm. The Paris System of dosimetry is based upon determining the minimum dose in the region between the sources as arranged in allowed geometries and then defining the reference dose as a percentage of this minimum dose. The net result is to define a prescription dose that encompasses the entire treatment volume, assuming that the rules of the Paris System have been followed. The allowed geometries are a single plane with equally spaced sources and a double plane with sources either in a triangular or a square pattern. The active length of the sources is larger than the treated length by a factor of 1.4. The spacing between the sources is varied, depending upon the thickness of the target to be treated. Basic concepts in the Paris System include: 1. The basal dose or basal dose rate, which is the minimum dose rate in the central plane between two or more sources. (The central plane is generally the plane, which bisects the sources.) 2. The reference dose or reference dose rate, which is taken as 85% of the average of the basal doses. (The 85% criterion is based upon the clinical experience of the team of oncologists and physicists who developed this system.) 3. The treatment volume, which is the volume enclosed by the reference dose.
Table 4. Distribution of Activity for Volume Implants Volume Shape
Distribution of Activity
Sphere Cylinder Crossed at active ends 1 end uncrossed 2 ends uncrossed Crossed at needle tips 1 end uncrossed
Rind 6 Parts Belt 4 4 4 4 4
Core 2 Parts Core 2 2 2 2 2
End A 1 1 0 2 2
End B 1 0 0 2 0
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Michael T. Gillin and Firas Mourtada 4. A central plane, which is a plane perpendicular to the long axis of the sources and which bisects the sources. 5. A hyperdose volume, which is the volume surrounded by the 170% isodose surface. Note that it is twice the reference dose value. 6. The use of the same linear activity sources for the implant that is the same linear reference kerma rate or the same mCi/cm for the entire implant. Marinello (Marinello et al. 1985) has defined the Paris System criteria for the use of seeds, namely that the spacing between the seeds should be less than or equal to 1.5 times the active length of the seed. Thus for seeds which are 3 mm in length, the spacing between seeds, if the Paris System is being used, should be less than or equal to 4.5 mm.
Chapter 4 in Pierquin and Marinello (1997) contains a comprehensive description of the Paris System. There are multiple figures in this chapter that present the basic geometric concepts of the Paris System. Table 5 of this chapter presents the relationships between source length and source spacing for single and double plane implants. There is also a summary of these rules in an article by Gillin et al. (1984). An initial example using the Paris System is a simple single plane implant. Consider a target volume of the following dimensions, length 2.0 cm, width 1.2 cm, and thickness 0.5 cm (see Figure 8). The ratio of the treated thickness to the spacing for two sources is 0.5. Thus the separation between the sources is 1.0 cm. The ratio of the treated length to the active length is 0.7. Thus the source length is approximately 2.8 cm. The basal dose is calculated at one point, midway between the sources in the central plane. If the
Figure 8. Example of Paris system single plane implant.
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Table 5. Predictive Relationships in the Paris System for Rectilinear Sources Equal in Length Patterns
2 Lines N lines in 1 plane N lines in squares N lines in triangles
Ratio of Treated Length to Active Length
Ratio of Treated Thickness to Spacing
Ratio of Lateral Margin to Spacing
Ratio of Safety Margin to Spacing
0.7 0.7 0.7 0.7
0.5 0.6 1.55–1.60 1.3
0.37 0.33 — —
— — 0.27 0.20
(Reprinted from Marinello, G. and B. Pierquin, A Practical Manual of Brachytherapy, Table 4.1. © (1997) with permission of Medical Physics Publishing, Madison, WI.)
prescription dose is 60 Gy, delivered at 0.6 Gy/h, the required linear activity can be determined. Appendix 2 in Pierquin and Marinello (1997) provides absorbed dose rates in water for 192Ir wire of two different diameters, 0.3 mm and 0.5 mm for a 1 U/cm (1 µGy h–1 m2 cm–1) source as a function of wire length and distance. The calculated dose rate for a 1 U/cm source for this two-source implant is 11.4 cGy/h. To obtain the 60 cGy/h dose rate, a source of 5.24 U/cm is required. AAPM Task Group 43 (TG 43) report (Nath et al. 1995), Table IV provides a conversion factor of 4.2 U/mCi (apparent activity), so that in older units the linear apparent activity is 1.25 mCi/cm. However, in the United States 192Ir seeds would most likely be used as opposed to wire. Following the Paris rules, these seeds would be spaced at 5 mm center-tocenter spacing and each seed would contain 0.62 mCi. Historically, American vendors wanted the 192Ir activity expressed in terms of milligram radium equivalents (mg RA eq), so that each seed would contain approximately 0.35 mg Ra eq. Consider the Paris System approach to the 2.5 cm × 3.5 cm × 2.5 cm thick target volume, described in the Manchester System double plane implant by Godden. The solution offered by Godden involves the use of five needles per plane, spaced 1.0 cm apart. The separation between the source planes is 2.0 cm. There are no crossing needles (see Figure 9). The peripheral needles contain twice the activity of the central needles. A total of 18.5 mg Ra eq is used, which converts to 32.5 mCi of 192Ir. The Paris System would propose a double plane implant with a square pattern (see Figure 10). Using the relationships presented in Table 5, the Paris System would do the following: Active length = Treated Length/0.7 = 3.5/0.7 = 5.0 cm. Source Spacing = Treated Thickness/1.6 = 2.5/1.6 = 1.6 cm.
(17)
Safety margin = 0.27 to source spacing = 0.27 × 1.6 = 0.4 cm. Thus the Paris System requires eight sources, four in each plane, with a separation of 1.6 cm between the sources and 1.6 cm between the planes. The basal dose would be calculated in the center of the three squares that are formed by this implant, and averaged. The reference dose would be 0.85 times the basal dose. The 5.0 cm long sources will extend in each direction approximately 7.5 mm beyond the 3.5 cm long target. The outer two pairs of sources would be outside the target volume, while the inner two pairs would be centered in the target volume. What is the linear activity needed if the Paris System is to deliver the same dose, 60 Gy, over the same time period, 168 hours, using 192Ir wires? One could perform an unfiltered line source calculation, assuming a dose rate of 1 mCi/cm and then scale the result. Given the fixed geometries and the calculation in the central plane, the manual calculations for the three basal dose rates are quite easy.
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Figure 9. Example of Manchester System double plane implant.
Figure 10. Example of Paris system double plane implant.
Systems 1B Manchester Planar and Volume Implants and the Paris System
DR =
Γ A −1 l ⋅ ⋅ 2 ⋅ tan , 2h h l
367
(18)
where Γ is the dose rate constant, h is the perpendicular distance to the center of the square, A is the activity, and l is the active length, 5 cm. In this example, the perpendicular distances, h, are approximately 1.13 cm, 2.6 cm, and 4.2 cm. A linear activity of approximately 0.67 mCi/cm would be required. Thus, the total activity would be approximately 26.7 mCi, which is close to the 32.5 mCi if the Manchester approach were used. It is also possible to perform manual calculations using the TG-43 approach. " θ ) = D" (r ,θ ) [G(r, θ) / G(r , θ ) ] g(r)F(r,θ ) . D(r, 0 0 0 0 0
(19)
TG 43 (Nath et al. 1995) defines the geometry factor, G(r,Θ), which accounts for the spatial distribution of activity within the source, ignoring photon absorption and scattering in the source structure. For a line source, TG 43 presents the geometry factor as: G(r,θ ) =
β Lr sin(θ )
.
(20)
where L is the active length of the source, β is the angle subtended by the active source with respect to the point (r, Θ), and r is the distance from the point of interest to the center of the source. In chapter 9 of Pierquin and Marinello (1997), Calitchi and Marinello describe a single plane implant that was used to treat a recurrence in a chest wall following mastectomy. The thickness of this implant was 1.0 cm, while the length was 3.0 cm and the width 6.0 cm (see Figure 11). The spacing between the sources, as calculated by the Paris System rules for this 1.0 cm thick implant, is 1.0/0.6 or 1.7 cm. Four sources with an active length of 4.3 cm (3/0.7) would generously cover this implant. This example displays the flexibility of the Paris System with its variable source spacing. The following calculation uses a TG43 approach to calculate the dose rate in the central plane between sources 2 and 3. The determination of the dose rate constant in water, Λ, for a 4.3-cm long 192Ir source is a challenge. The dose rate constant presented in Nath et al. (1997, Table VI) for 192Ir, 1.12 cGy h–1 U–1 , is for a seed. Dose rate constants are a function of source composition and source length. Karaiskos et al. (2001) provide Monte Carlo and Sievert calculated dose rate constants for three different diameter 192Ir sources as a function of source length. Figure 3 of Karaiskos et al. (2001) can be interpolated to yield a dose rate constant of 0.6 cGy h–1 U–1 for a 4.3 cm long 192Ir source with an outer diameter of 0.3 mm. Ballester et al. (1997) earlier calculated Monte Carlo dose rate distributions around 192Ir wires. Table IV of that work can be interpolated to yield a dose rate constant for this 4.3 cm source of 0.61 cGy h–1 U–1.
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Figure 11. Paris System single plane implant with variable source spacing.
D(r,Θ) = SkΛ[G(r,Θ)/G(r0,Θ0]g( r)F(r,Θ) ,
(21)
where Sk equals 1 U, Λ equals approximately 0.60 cGy h–1 U–1, g(r ) is essentially 1.0 over the distances being considered, F(r,Θ) is 1.0 for calculations in the central plane, G(r,Θ) is equal to β/Lr sin θ, r0 equals 1.0 cm, Θ0 equals Π/2, r is the distance from the sources to the point of calculation, which is this case is 0.85 cm and 2.55 cm, G(0.85, Π/2) = 0.651 cm–2, G(1.00, Π/2) = 0.528 cm–2, G(2.55, Π/2) = 0.128 cm–2, D = 1 U × 0.6 cGy h–1 U–1 {2 × 0.651/0.528 + 2 × 0.128/0.528}, and D = 1.77 cGy/h for 1 U. Using a conversion factor of 4.205 mCi–1, the total activity of each source is 0.238 mCi. Assume that the desired activity is 1.0 mCi/cm. Then the TG-43 calculated dose rate is D = 1.77 × 4.3/.238 = 32 cGy/h . This calculated value is reassuringly close to the unfiltered line source result of 30.8 cGy/hr.
(22)
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What limits the separation between the sources? Why not use three sources, which are placed 3.0 cm apart, to cover the 6.0 cm width? The answer is that such a distribution would result in large hyperdose sleeves, which most likely would lead to necrosis. The Paris System recommendation, based upon clinical experience, is to limit the diameter of the hyperdose sleeve to 8 to 10 mm or less. In chapter 8 of Pierquin and Marinello (1997), Pernot and Marinello describe two mobile tongue implants. The larger implant has target volume dimensions of 4 cm in width by 3.5 cm in length by 2.5 cm in thickness. Their solution is to use three 192Ir loops in a pattern of squares. The suggested spacing is 1.6 cm. Each wire is approximately 9 cm in length. This would result in an active length of each leg being approximately 3.7 cm in length and a crossing piece of 1.6 cm. If the three loops are centered in the target volume, then there is a safety margin of 0.43 cm (0.27 × 1.6 cm) on each end of the implant. The total width of the implant, according to the resultant Paris System distribution, is 3.2 cm + 2 × 0.43 cm, or 4.06 cm. The target dimensions of the smaller implant are 2.5 cm width, 3.0 cm in length, and 1.5 cm in thickness. Such an implant is ideal for the use of 192Ir pins, which have a 1.2 cm separation between their legs. A pattern in 12-mm squares produces an appropriate solution. The only minor challenge is determining the length of the legs of the pin. In the Paris System, the sources should be 10+% longer than the target volume on each side. Thus, in order to cover the 3.0 cm length, the sources should extend approximately 3 mm longer on the side that does not have the crossing piece of the pin.
Optimized Stepping Source Dosimetry System An optimized stepping source dosimetry system has been described by van der Laarse and Prins from Nucletron Research (van der Laarse 1994). This stepping source system has evolved from the Paris System and includes computer-based planning. The significant features of this system include: 1. Equidistant dwell positions have optimized dwell times. 2. Dose specification points for the purposes of optimization are defined midway between the catheters along the active length of the catheter. 3. Active dwell positions remain inside the target volume. Major et al. (2002) compared the Paris System and the dose point system in terms of conformality and homogeneity of dose distributions. This paper addressed the active source length question, i.e., should active dwell positions remain within the target volume, as suggested in the optimized stepping source system, or should they extend past the target volume by some amount. They conclude that the optimal active lengths depend upon the catheter separation, the method of optimization, and the geometrical shape of the target volume. The distance between the outer source position in the catheter and the surface of the Planning Target Volume (PTV) is between 2.5 and 5 mm, with the source positions being in the PTV. They also conclude that in optimized systems an isodose value, which is higher than the 85% of the mean central dose, should be used to define the reference dose. As shown in Figure 12, a simple comparison, which involves two catheters separated by 1.0 cm, between the Paris System and the optimized stepping source dosimetry system has been performed. In the Paris System, the equal linear activity sources extended past the target volume and in the optimized stepping source system, the sources (dwell positions) were confined to the target volume. Dose points were defined halfway between the catheters and opposite each dwell position. The optimization was performed on all dose points on volume, which is equivalent to optimizing on dose points at a distance. The dwell times in the optimized system varied by approximately a factor of three, with the shortest times in the center and the longest times being at the first and last dwell positions.
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Figure 12. Comparison of optimized stepping source system (above) with the Paris System (below).
The introduction of remote afterloading devices, both high dose rate and pulsed dose rate, raises many clinical questions, especially with respect to time-dose relationships. A comprehensive article on head and neck interstitial implants by Mazeron et al. (2002) discusses some of these questions as part of an overall review.
Actual Practice Several recent interstitial implants from U.T. MDACC have been reviewed. These implants were performed by different oncologists. The first implant was a single plane implant for recurrent malignant fibrohisteocytoma of the right upper extremity. The prescription was 45 Gy at 60 cGy/h. Fifteen catheters were implanted with spacing that ranged between 5 mm and 15 mm. (The last 2 catheters on one side had the 15 mm separation, while the separation between the remaining 13 was 5 to 10 mm.) The active length varied between 4.5 cm to 5.5 cm. 192Ir-wire was used with a linear activity of approximately 1 mCi/cm. The total activity was 68.8 mCi. No particular system was used. However, the implant closely resembles a Paris System implant. A manual, unfiltered line source calculation produces a basal dose of approximately 77 cGy/h, while the computer calculation yields 82 cGy/h at the same location. The difference between the computer calculation and the manual calculation reflects distance approximations made in the manual calculation. The Paris System suggests a reference dose rate of 65 cGy/h, while in fact 60 cGy/h was used. Given the larger separation between the last two catheters, the reference dose surface is not contiguous. The reference dose volume breaks up around catheter 14 and 15. The hyperdose regions do not exceed the Paris guidelines. In the opinion of the oncologist, this is an acceptable implant. The second implant was essentially a two-plane implant plus an additional source for the nasal septum. There were two sources in one plane, three sources in the second plane, which was separated by 2 cm, and one source in the middle essentially at 1 cm from each plane. The source were 2.2 cm long Ir-wire with a linear activity of 1.01 mCi/cm. A prescription dose rate of 50 cGy/h was used to a total dose of 67 Gy. This implant did not conform to any system, but in the judgment of the oncologist covered the target
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volume. In the central plane, the prescription dose rate covered the target volume. This is also true at a distance of 5 mm in either direction from the central plane. However, at 10 mm in either direction, the prescription isodose no longer defined a contiguous volume. The hyperdose sleeves were within the Paris System criteria. The patient tolerated well and displayed minimum tissue changes one year later. The final implant was a permanent gold seed implant for unresectable rectal carcinoma. One hundred fifty-three seeds, at 1.81 mCi/seed, were implanted. CT-based dose distributions were calculated. The prescription isodose encompassed an approximately 6-cm-long volume, changing shape on each axial slice reflecting changes in the seed distribution. This implant did not follow a specific system. However, the oncologist was satisfied with this palliative procedure.
Summary The classical systems may not be widely used in current clinical practice. They do provide a reference for comparison of contemporary work. They also reflect significant clinical experience and suggest limits to fundamental parameters, such as separation between the sources. Interstitial brachytherapy still offers a unique therapeutic approach in selected circumstances. Knowledge of the classical systems is useful both in the planning of contemporary implants and in the evaluation of such implants. Interstitial brachytherapy has been an important element in the history of radiation oncology. It is important for contemporary medical physicists to know this history. As Robert Heinlein noted in The Notebooks of Lazarus Long, “A generation which ignores history has no past and no future.”
References Ballester, F., C. Hernandez, J. Perez-Calatayud, and F. Lliso. (1997). “Monte Carlo calculation of dose rate distributions around 192Ir wires.” Med Phys 24:1221–1228. Gibb, R., and J. B. Massey. (1980). “Radium dosage: SI units and the Manchester system.” Br J Radiol 53:1100–1101. Gillin, M. T., R. W. Kline, J. F. Wilson, and J. D. Cox. (1984). “Single and double plane implants: A comparison of the Manchester System with the Paris System.” Int J Radiat Oncol Biol Phys 10:921–925. Godden, T. J. The Physical Aspects of Brachytherapy. Medical Physics Handbook 19. Bristol: Adam Hilger, 1988. Johns, H. E., and J. R. Cunningham. The Physics of Radiology, 4th Edition. Springfield, IL: Charles C Thomas Publishers, 1983. Karaiskos, P., P. Papagiannis, A. Angelopoulos, L. Sakelliou, D. Baltas, P. Sandilos, and L. Vlachos. (2001). “Dosimetry of 192Ir wires for LDR interstitial brachytherapy following the AAPM TG-43 dosimetric formalism.” Med Phys 28:156–166. Major, T., C. Polgár, J. Fodor, A. Somogyi, and G. Németh. (2002). “Conformality and homogeneity of dose distributions in interstitial implants at idealized target volumes: A comparison between the Paris and dose-point optimized systems.” Radiother Oncol 62(1):103–111. Marinello, G., M. Valéro, S. Leung, and B. Pierquin. (1985). “Comparative dosimetry between iridium wires and seed ribbons.” Int J Radiat Oncol Biol Phys 11:1733–1739. Mazeron, J. J., G. Noël, and J. M. Simon. (2002). “Head and Neck Brachytherapy.” Semin Radiat Oncol 12(1):95–108. Nath, R., L. L. Anderson, G. Luxton, K. A. Weaver, J. F. Williamson, and A. S. Meigooni. (1995). “Dosimetry of interstitial brachytherapy sources: Recommendations of the AAPM Radiation Therapy Committee Task Group No. 43.” Med Phys 22(2):209–234. Also available as AAPM Report No. 51. Paterson, R., H. M. Parker, F. W. Spiers, M. C. Tod, S. K. Stephenson, and W. J. Meredith. Radium Dosage, The Manchester System. W. J. Meredith (ed). Baltimore, MD: The Williams and Wilkins Company, 1947. Perez, C. A., and L.W. Brady. (eds.). Principles and Practices of Radiation Oncology, Second Edition. Philadelphia: J. B. Lippincott Company, 1992.
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Pierquin, B., and G. Marinello. A Practical Manual of Brachytherapy. Madison, WI: Medical Physics Publishing, 1997. Shalek, R. J., and M. Stovall. Brachytherapy Dosimetry, The Dosimetry of Ionizing Radiation. Vol. III. K. R. Kase, B. E. Bjarngard, and F. H. Attix (eds). San Diego, CA: Academic Press, Inc., 1990. Shalek, R. J., and M. Stovall. Implant Dosimetry. The University of Texas M.D. Anderson Hospital and Tumor Institute at Houston, 1967. van der Laarse, R. (1994). “The stepping source dosimetry system as an extension of the Paris System.” Brachytherapy from Radium to Optimization 34:319–330.
Chapter 19
Quimby-Based Brachytherapy Systems Robert D. Zwicker, Ph.D. Department of Radiation Medicine University of Kentucky, Lexington, Kentucky Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 373 The Quimby System . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 374 Planar Source Distribution (226Ra) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 374 Dose Homogeneity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 374 Planar Exposure Tables . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 374 Volume Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 375 The Kwan System . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 375 Source Distribution (192Ir) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 375 Prescription Point Locations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 376 Dose Tables/Graphs . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 377 The Memorial System . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 379 Physical Optimization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 379 Prescription Point Location . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 380 192 Ir Nomograph . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 380 The Tufts/MCV System (Zwicker System) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 381 Source Distribution/Dose Uniformity in Seed Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 381 Prescription Surface Definition . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 381 Interplanar Spacing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 382 Dose Tables . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 382 The Saw Method . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 383 Optimization Indices . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 383 Optimal Reference Dose Rates . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 383 Generalized Optimization Indices . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 383 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 384
Introduction Systems of source distribution and dose determination for surface and interstitial brachytherapy were introduced decades ago by Quimby (1932) and by Paterson and Parker (1949), and these systems still influence the planning of brachytherapy treatments today. Low dose rate (LDR) interstitial implants most often have been carried out using uniform activity sources, yielding a Quimby-like dose distribution. With remote afterloading and variable dwell times, planar implants seeking dose uniformity in planes parallel to the implant planes will yield optimized activity distributions similar to those of the Manchester system. Even with more complex treatment target and source geometries, an understanding of the basic principles of source loading and consequent dose distribution, as illustrated by the early systems, is essential to effective interstitial brachytherapy treatment planning. The early implant systems, including those guiding intracavitary implants, were developed for use with 226 Ra, which was the primary isotope for use in brachytherapy from the earliest days of radiation therapy until its gradual replacement by 137Cs and 192Ir, completed in the United States about two decades ago. The use of active needles for interstitial brachytherapy was curtailed and generally replaced by the use of inserted catheters that could be afterloaded with 192Ir seeds or wires in the convenience of the patient’s room. The transition from radium needles to iridium seeds or wire did not, however, change the basic considerations of activity distribution and consequent dose distribution, and the early systems remained in use with minor modifications. In the following sections we will examine specifically the results of
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uniform planar source activity distribution, as examined first by Quimby (1932), and pursued in more detail in the ensuing years by a number of others (Anderson, Hilaris, and Wagner 1985; Anderson and Osian 1986; Kwan et al. 1983; Pierquin et al. 1978; Saw and Suntharalingam 1988; Zwicker, Schmidt-Ullrich, and Schiller 1985).
The Quimby System Planar Source Distribution (226Ra) The earliest planar brachytherapy system was described by E. H. Quimby in 1932. This in fact was not a true system in the sense of examining possible source distributions and providing consequent guidance to clinicians in the selection and placement of sources. It simply assumed an equally spaced array of uniform 226Ra source activities and presented the resulting exposure rates to be achieved at various distances from the source plane along a line perpendicular to and through the center of the implanted plane.
Dose Homogeneity It is of interest to examine the dose distributions resulting from a Quimby-like (uniform activity) planar implant, as shown for a single plane 192Ir implant in Figure 1. It is apparent and not surprising that in a parallel plane at a fixed distance from the implant plane, the dose is higher near the center of the implant than at the edges. This is expected because points near the center of the implant are in close proximity to more sources than are points near the edges. The isodose surfaces surrounding the implant are therefore thicker near the center than near the edges, as indicated in Figure 1, and this fact must be taken into consideration in planning a Quimby-type implant. Especially for large single-plane implants, it is important that appropriate planning should be carried out to ensure that doses delivered near the implant boundaries are sufficiently high to achieve the clinical intent of the procedure. The original Quimby system was applied only to the dosimetry of surface applicators, and did not address the problems encountered with implants of two or more planes. The characteristics of biplanar implants, however, are similar to those of single planes in that higher doses are typically delivered in the central portions of the target volume rather than near the edges, and the isodose surfaces of interest are thicker near the center of the implant. An example is given in Figure 2, which shows a 6×6 cm biplanar implant designed to treat a target volume of dimensions 6×6×2.5 cm along the symmetry axes. The reference isodose curve in this case is shown extending to the target boundaries along these axes, leaving the corners of the rectangular target volume at a lower dose. Points near the center of the implant, midway between the two implant planes, receive doses around 25% to 30% higher than the reference dose. If the chosen reference dose for this implant is extended outward to guarantee complete coverage of the target volume, the central dose will exceed the reference dose by an even greater amount. Elucidation and optimal exploitation of these characteristics of Quimby-like implants was a major impetus for the development of later implant systems specific to the use of 192Ir seeds or wires.
Planar Exposure Tables In the original Quimby work, tables were published which presented the exposure per milligram-hour to the center of the field at various distances from the 226Ra implant plane, for various implant areas. With the change to 192Ir and more recent changes in the basic calibration of these sources, the original tables of Quimby can no longer be regarded as directly useful for dose determination.
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Figure 1. Isodose curves for a single plane 192Ir seed implant with seed and ribbon spacings of 1.0 cm, and uniform source activities. (Reprinted from Saw, C. B., and N. Suntharalingam, “Reference dose rates for single- and double-plane Ir-192 implants,” Med Phys 15:391–396, © 1988 with permission of AAPM.)
Volume Implants For Quimby volume implants sources were expected to be uniformly distributed and spaced 1 to 2 cm apart. A table was provided for varying implant volumes, giving the source strengths required to deliver 1000 cGy minimum dose inside the implanted volume (Anderson and Presser 1995). Corrections for elongated volumes were the same as for the Manchester system. As with Quimby planar implants, tissues near the center of the implanted volume are expected to receive doses significantly higher than the minimum dose delivered at the volume boundaries.
The Kwan System Source Distribution (192Ir) Kwan and colleagues (1983) took note of the considerable dose inhomogeneity that is unavoidable when uniform activity seeds in a planar distribution are used to treat rectangular target volumes. These researchers carried out detailed studies of the doses delivered at two points of particular interest in Quimby-
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Figure 2. Isodose curves for a biplanar 192Ir implant with seed and ribbon spacings of 1.0 cm, and uniform source activities. (Reprinted from Zwicker, R. D., and R. Schmidt-Ullrich, “Dose uniformity in a planar interstitial implant system,” Int J Radiat Oncol Biol Phys 31:149–155, © 1995 with permission from Elsevier.)
like monoplane and biplanar implants. The studies incorporated square source planes, with sources spaced at 1.0 cm intervals along the source ribbons. The intraplanar ribbon spacing and interplanar separation were both set at either 1.0 cm or 1.5 cm, and the resulting dose homogeneity was reported and evaluated.
Prescription Point Locations The first point of interest in this study is located on the boundary of the rectangular target volume, near a corner of the volume, and is associated with the minimum dose to be allowed to the target tissues (Figure 3). The second point is near the geometric center of the target volume, positioned midway between seeds, and is identified as a target dose maximum. With the reference dose defined in this manner, the ratio of the minimum to maximum target doses can be quite low, depending on implant construction (sources inside, outside, or just at the target boundaries) and the lateral dimensions of the implant plane or planes. Based on a combination of dose uniformity and normal tissue sparing, Kwan et al. (1983) concluded that the best source configuration is that illustrated in Figure 3b, with the sources implanted exactly to the edge
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Figure 3. End-on view of biplanar implant showing interest points in Kwan system. (Reprinted from Kwan, D. K., A. R. Kagan, A. J. Olch, P. Y. M. Chan, B. Hintz, and M. Wollin,“Single and double-plane iridium-192 interstitial implants: Implantation guidelines and dosimetry,” Med Phys 10:456–461, © 1988 with permission of AAPM.)
of the target volume. With this configuration in place, the characteristics of implants of dimensions 2×2 cm to 10×10 cm were examined in this study.
Dose Tables/Graphs Figure 4 shows Kwan’s dose rate curves for the maximum and minimum target dose points, plotted as a function of implant area, for a 2.0 cm target thickness, with individual seed activities of 0.5 mCi. Results for the two possible ribbon and interplanar separations of 1.0 and 1.5 cm are shown. These curves can be used for preplanning of biplanar implants by scaling the seed activity to achieve a desired minimum dose
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rate for implants of thickness 2.0 cm constructed in this manner. For a 2.5 cm target thickness, the dose rates given in Figure 4 must be scaled downward by a factor of 0.87. Figure 5 shows the ratio of the minimum-to-maximum target dose for biplanar implants, again plotted as a function of implant area. Kwan et al. (1983) concluded from these curves that for biplanar implants treating target thicknesses of 2.0 or 2.5 cm, a ribbon and interplanar separation of 1.5 cm is preferred. Even then, however, the minimum-to-maximum dose ratio is seen to vary from 0.5 down to about 0.38, which suggests that if the prescription dose is identified as the minimum target dose for such an implant, a large fraction of the target volume near the center of the implant may receive doses at levels two or more times the prescription dose.
Figure 4. Dose rate curves from Kwan system showing maximum and minimum target dose points as a function of implant area. Target thickness is 2.0 cm and seed activities 0.5 mCi. (Reprinted from Kwan, D. K., A. R. Kagan, A. J. Olch, P. Y. M. Chan, B. Hintz, and M. Wollin,“Single and double-plane iridium-192 interstitial implants: Implantation guidelines and dosimetry,” Med Phys 10:456–461, © 1988 with permission of AAPM.)
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Figure 5. Ratio of minimum to maximum target dose for biplanar implants, according to Kwan system. (Reprinted from Kwan, D. K., A. R. Kagan, A. J. Olch, P. Y. M. Chan, B. Hintz, and M. Wollin, “Single and double-plane iridium-192 interstitial implants: Implantation guidelines and dosimetry,” Med Phys 10:456–461, © 1988 with permission of AAPM.)
For rectangular implant areas, Kwan et al. (1983) proposed the use of an equivalent square, determined as for external beam dosimetry, to extract data from the square-area curves for planning purposes. Their studies showed this technique to be accurate to within about 3% for single-plane and 5% for two-plane implants. Studies of actual implants in which the source positions sometimes deviated by ±2 mm from their ideal locations yielded agreement in the minimum target doses within 5% between the computer treatment planning calculations and the published dose rate curves.
The Memorial System Physical Optimization Workers at Memorial Sloan-Kettering Cancer Center (MSKCC) first published guidelines for an 192Ir implant design over four decades ago (Laughlin et al. 1963). The system later introduced at MSKCC by Anderson, Hilaris, and Wagner (1985) and by Anderson and Osian (1986) was intended to make use of existing in-house stocks of 192Ir seeds in ribbons, and a nomograph was developed for staggered biplanar implants to determine the number of ribbons needed in the base plane for a fixed-seed activity to achieve a dose rate of about 10 Gy/day at a reference point near the periphery of the implant volume. It was expected that this choice of activity distribution and reference dose point selection would represent
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a “physical optimization” in some sense, achieving the best possible approximation to the desired dose rate distribution with the sources available. No assumptions are made concerning the intraplanar seed spacings. For these implants the superficial plane has one fewer ribbon than the base plane, and the interplanar separation is 1.5 cm.
Prescription Point Location The reference point for this system is described as being located 1.5 spacing units (usually 1.5 cm) inside end sources in the ribbon direction, midway between edge ribbons, and midway between source planes. As the reference point is near the implant corner, it is expected that the reference isodose surface will cover most of the target tissues if the target thickness is 1.5 cm. 192
Ir Nomograph
A nomograph was developed at MSKCC to make use of in-house stocks of 192Ir of known seed strength (Figure 6). The purpose of the nomograph was to determine, for a given seed activity, the number of 192Ir ribbons required to deliver a specified dose rate to the reference isodose surface. Input data needed to use the nomograph include the active ribbon length, base plane width, and seed strengths of available sources. A similar approach was later developed for volume implants using 125I, and still later for 103Pd implants.
Figure 6. Nomograph used to determine source data from the Memorial system. (Reprinted from Anderson, L. L., and A. D. Osian, “Brachytherapy optimization and evaluation,” Endocuriether Hypertherm Oncol (now called Brachytherapy) 2:S25–S32. © 1986 with permission from American Brachytherapy Society)
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The Tufts/MCV System (Zwicker System) Source Distribution/Dose Uniformity in Seed Implants In the work of Zwicker, Schmidt-Ullrich, and Schiller (1985) it was recognized that the conditions of the Manchester system for higher activity at the implant periphery or extension of active sources outside the target volume could often not be realized. This was especially true of breast implants, in which anatomical constraints generally precluded the use of crossing needles or catheters, and extension of the implanted area beyond the target volume could leave active sources at the level of the skin, likely resulting in added discomfort for the patient and a poor cosmetic result. It was further recognized that the need for enclosing a rectangular volume in the prescription isodose surface was not clear, as tumors and surgical beds are not expected to assume a rectangular shape. In view of this, a system was developed for biplanar implants of target thickness 2.0 to 3.5 cm which would guarantee the extension of the reference isodose surface to the exact target boundaries along the three symmetry axes, but would allow the treated volume to be thinner than the stated target thickness (but no thinner than the interplanar source separation) near the implant corners. It was expected that this would lead to a more homogeneous dose distribution in terms of the central planar doses. This system also assumed intraplanar seed and ribbon spacings of 1.0 cm, but allowed the interplanar separation to vary in order to effect exact coverage of the target along the symmetry axes.
Prescription Surface Definition Figure 7 illustrates the concepts of this system for a 6×6×2.5 cm biplanar implant with an optimized interplanar separation of 1.3 cm. It is seen that the reference isodose provides exact coverage of the target volume along the implant symmetry axes. A greater interplanar separation for this implant would force the reference isodose curve out beyond the target boundary in the direction of the target thickness, while a smaller separation would increase unnecessarily the dose in the central plane relative to the reference dose.
Figure 7. Isodose curves for a biplanar 6×6×2.5 cm 192Ir implant planned by the Zwicker system, with seed and ribbon spacings of 1.0 cm, and optimized interplanar separation of 1.3 cm. (Reprinted from Zwicker, R. D., R. Schmidt-Ullrich, and B. Schiller, “Planning of Ir-192 seed implants for boost irradiation to the breast,” Int J Radiat Oncol Biol Phys 11:2163–2170, © 1985 with permission from Elsevier.)
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Interplanar Spacing Exact coverage of the target along all three symmetry axes is accomplished by finding the optimal interplanar separation that accomplishes this. Table 1 gives dose rates and optimized interplanar separations for biplanar implants with implant areas of 3×3 to 10×10 cm, target thicknesses of 2.0 to 3.5 cm, implanted with 1.0 mCi seeds. For implant planning purposes, the target dimensions are determined by any combination of direct examination, surgical report, films, and/or three-dimensional (3-D) imaging deemed appropriate. Once the target has been defined, the areas of the implant planes are known, and the interplanar separation needed for an optimal implant is taken from the table. As in the Kwan system, for rectangular implants an equivalent square is used to enter the data tables, but in this case the interplanar separation is reduced by 1 mm.
Dose Tables Table 1 shows the expected dose rates at the reference dose position for the range of biplanar implants described above, assuming seed activities of 1.0 mCi. For seed ordering purposes, these values can be scaled to find the required activity to achieve any desired dose rate. It is of interest that with the reference (prescription) dose defined in the manner described here and the interplanar separation varied for optimal target coverage, the dose near the geometric center of the implant, midway between planes, is only about 25% to 30% higher than the reference dose. A later work in this series (Zwicker and Schmidt-Ullrich 1995) investigated the effects of increasing the intraplanar ribbon spacing to 1.5 cm, leaving the seed spacing at 1.0 cm. It was shown that in an idealized point-source case, the small volume treated to very high dose in the vicinity of the seeds is increased by roughly the square root of the ratio of the number of seeds in the implant at 1.0 spacing to the number of seeds in the implant at 1.5 cm spacing. For typical cases, this translates to an increase of about 30% in the high dose volume if the intraplanar ribbon spacing is increased to 1.5 cm. In spite of this, circumstances sometimes dictate that the number of catheter insertions be minimized, so tables of dose rates and interplanar separations for 1.5 cm ribbon spacings were presented in this paper. This is reproduced in Table 2. Table 1. 192Ir Double-plane Implant Table. Ribbon Spacing = 1.0 cm. t(cm)
2.0
2.5
3.0
3.5
A(cm) 3×3
s 1.2
DR 61
s 1.5
DR 53
s 1.8
DR 46
s 2.1
DR 40
4×4
1.2
69
1.5
60
1.7
53
2.0
47
5×5
1.1
79
1.4
69
1.7
61
1.9
55
6×6
1.1
84
1.3
75
1.6
67
1.9
60
7×7
1.0
92
1.3
81
1.5
73
1.8
66
8×8
1.0
96
1.2
87
1.5
78
1,7
71
9×9
0.9
103
1.2
91
1.4
83
1.7
75
10×10
0.9
107
1.1
97
1.4
87
1.6
80
Double-plane implant, seed activity = 1.0 mCi; dose rate DR in cGy/h. Note: For rectangular implants, use equivalent square field size formula, then reduce s by 1.0 mm. (Reprinted from Zwicker, R. D., R. Schmidt-Ullrich, and B. Schiller, “Planning of Ir-192 seed implants for boost irradiation to the breast,” Int J Radiat Oncol Biol Phys 11:2163–2170, © 1985 with permission from Elsevier.)
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Table 2. 192Ir Double-plane Implant Table. Ribbon Spacing = 1.5 cm. t(cm)
2.0
2.5
3.0
3.5
A(cm) 3×3
s 1.2
DR 38.5
s 1.5
DR 34.0
s 1.8
DR 30.1
s 2.1
DR 26.8
4.5×4.5
1.1
46.7
1.4
41.7
1.7
37.3
2.0
33.5
6×6
0.9
55.2
1.3
49.0
1.6
44.3
1.9
40.2
7.5×7.5
0.8
60.8
1.2
54.5
1.5
49.5
1.8
45.1
9×9
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Double-plane implant, seed activity = 1.0 mCi; dose rate DR in cGy/h. (Reprinted from Zwicker, R. D., and R. Schmidt-Ullrich, “Dose uniformity in a planar interstitial implant system,” Int J Radiat Oncol Biol Phys 31:149–155, © 1995 with permission from Elsevier.)
The Saw Method Optimization Indices Saw and Suntharalingam (1988) recognized the influence of the choice of reference dose rate on the quality of an implant and carried out a detailed study of this effect using three volumetric indices to characterize implant quality. The coverage index (CI) was defined in terms of the percent of the target volume receiving dose rates equal to or greater than the reference dose rate. The relative dose homogeneity index (HI) gave the percent of the target volume receiving dose rates between 100% and 150% of the reference dose rate. The external volume index (EI) represented the percent (normalized to target volume) of volume outside the target volume receiving dose rates equal to or greater than the reference dose rate.
Optimal Reference Dose Rates Calculations were carried out for two implant geometries: a single plane treating a 1.0-cm-thick volume, and a biplanar implant treating a volume of thickness 2.5 cm. Implant areas varied from 3×3 cm to 12×12 cm. Intraplanar seed and ribbon spacings were 1.0 for the single-plane implants. For the two-plane implants, seed spacings were 1.0 cm within the ribbons, but both intraplanar ribbon spacing and interplanar separation were 1.5 cm. Seed activities were 1.0 mCi in both cases. The study examined the choice of reference dose rate for each implant geometry, and plotted the three indices as functions of reference dose. It was found in the cases presented that the HI reached a maximum, typically in the range where the external coverage index was at a low value. The plots for biplanar implants are shown in Figure 8. It was proposed, therefore, that the optimal dose rate for each geometry was that at which HI was maximum. Values of EI, CI, and HI were then plotted for optimized implants (Figure 9). The coverage index over the range of implant areas studied varied from 85% to 95%.
Generalized Optimization Indices Further studies of the effect of reference dose choice on dose uniformity in biplanar implants were carried out with variable interplanar separations and with the upper limit of the HI also allowed to vary (Zwicker
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Figure 8. External volume index (EI), coverage index (CI), and dose homogeneity index (HI) plotted as a function of reference dose, for a 6×6 cm biplanar implant, as described in the Saw system. Curves are for implants with seed spacing 1.0 cm, and intraplanar and interplanar ribbon spacings of 1.5 cm. Implant areas: (a) 3×3 cm; (b) 6×6 cm; (c) 9×9 cm; (d) 12×12 cm. (Reprinted from Saw, C. B., and N. Suntharalingam, “Reference dose rates for single- and double-plane Ir-192 implants,” Med Phys 15:391–396, © 1988 with permission of AAPM.)
and Schmidt-Ullrich 1995). It was shown here that for variable target thicknesses, optimization of the interplanar separation was an important element in achieving acceptable dose uniformity.
References Anderson, L. L., and A. D. Osian. (1986). “Brachytherapy optimization and evaluation.” Endocuriether Hypertherm Oncol 2:S25–S32. Anderson, L. L., and J. L. Presser. “Classical Systems I for Temporary Interstitial Implants: Manchester and Quimby Systems” in Brachytherapy Physics. J. F. Williamson, B. R. Thomadsen, and R. Nath (eds). Madison, WI: Medical Physics Publishing, pp. 301–322, 1995. Anderson, L. L., B. S. Hilaris, and L. K. Wagner. (1985). “A nomograph for planar implant planning.” Endocuriether Hypertherm Oncol 1:9–15.
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Figure 9. Optimized values of EI, CI, and HI for biplanar implants of implant area 3×3 to 12×12 cm. (Reprinted from Saw, C. B., and N. Suntharalingam, “Reference dose rates for single- and double-plane Ir-192 implants,” Med Phys 15:391–396, © 1988 with permission of AAPM.)
Kwan, D. K., A. R. Kagan, A. J. Olch, P. Y. M. Chan, B. Hintz, and M. Wollin. (1983). “Single and double-plane iridium-192 interstitial implants: Implantation guidelines and dosimetry.” Med Phys 10:456–461. Laughlin, J. S., W. M. Silver, E. I. Holodny, and F. W. Ritter. (1963). “A dose description system for interstitial radiation therapy.” Am J Roentgenol 89:470–490. Paterson, R., and H. M. Parker. “Interstitial Treatments” in Radium Dosage: The Manchester System. W. J. Meredith (ed.). Baltimore, MD: Williams and Wilkins Company, pp. 28–38, 1949. Pierquin, B., A. Dutreix, C. H. Paine, D. Chassagne, G. Marinello, and D. Ash. (1978). “The Paris system in interstitial radiation therapy.” Acta Radiol Oncol 17:33–48. Quimby, E. H. (1932). “The grouping of radium tubes in packs and plaques to produce the desired distribution of radiation.” Am J Roentgenol 27:18. Saw, C. B., and N. Suntharalingam. (1988). “Reference dose rates for single- and double-plane Ir-192 implants.” Med Phys 15:391–396. Zwicker, R. D., and R. Schmidt-Ullrich. (1995). “Dose uniformity in a planar interstitial implant system.” Int J Radiat Oncol Biol Phys 31:149–155. Zwicker, R. D., R. Schmidt-Ullrich, and B. Schiller. (1985). “Planning of Ir-192 seed implants for boost irradiation to the breast.” Int J Radiat Oncol Biol Phys 11:2163–2170.
Chapter 20
Implant Design and Execution Robert D. Zwicker, Ph.D. Department of Radiation Medicine University of Kentucky, Lexington, Kentucky Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 387 Implant Techniques . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 387 Steel Needle/Catheter Lead Insertion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 387 Catheter Seating: Buttons and Clamps . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 388 Blind-End Catheters . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 388 Freehand Versus Template Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 389 Dose Homogeneity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 389 Dose Homogeneity Indicators . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 389 Source Placement and Dose Uniformity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 390 Planar Implant Preplanning: Source Positioning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 391 Quimby-like Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 391 Manchester Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 391 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 392
Introduction Early interstitial implants were carried out using radium needles, generally inserted into the patient using special clamps, or, frequently, using the fingers to hold the sources. The difficulty of handling the sources and the speed with which the implantation had to be carried out were factors in reducing the quality of these implants. The introduction of afterloading techniques allowed a more relaxed atmosphere for positioning of source guides in the target tissues. For low dose rate (LDR) implants, the radioactive material, usually 192Ir, is then loaded into the carefully pre-positioned source guides, usually in the safety and convenience of the patient’s hospital room. For high dose rate (HDR) treatment, the source guides are connected to the treatment machine in the shielded HDR treatment room, and the remote-controlled source is stepped through the guides while the operators are in a shielded location outside the room. Interstitial implants as used in breast brachytherapy may involve the use of any number of source guides, typically 10 to 25 or more. Techniques for placement of source guides include the use of plastic catheters and/or steel needles, placed either freehand or with the help of a template. These techniques are described below
Implant Techniques Steel Needle/Catheter Lead Insertion The most common method of source guide placement (Perez, Zwicker, and Williamson 2003) involves the use of a hollow steel needle (a stylette, typically 17 gauge), inserted into and through the target tissues in the desired location and threaded with the solid-lead end of a catheter of the same gauge as the needle, as shown in Figure 1. Often the inserted needles are left in place while the next and remaining needles are inserted, as this greatly facilitates the relative positioning of the source guides (catheters). Once the needles are positioned and the catheter leads threaded through them, the needles, with the leads, are then pulled through the tissue and out the other side. The catheter follows into the needle’s previous position, and is pulled until the motion is stopped by a button on the end of the catheter (see Figure 1). The open end of the catheter is then usually secured by means of a friction washer or clamp, and the catheter may
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Figure 1. Insertion of guide needles and implant catheters. (Reprinted from Neblett (1995), Figure 1, p. 283, with permission from Medical Physics Publishing, ©1995 American Association of Physicists in Medicine.)
then be cut to any desired length. The friction washer or clamp may be sutured to the skin. If it is desired to treat tissues up to the skin, spacers may be positioned between the button and the skin to allow the source to protrude to the level of the skin. This can be especially helpful in the treatment of base of tongue disease (Schmidt-Ullrich et al. 1991). For implants involving a large number of catheters, later identification of the catheters may be facilitated by the use of catheters of different colors. In conventional brachytherapy planning, the catheters are filled with dummy ribbons containing radioopaque marker seeds at 1.0 cm intervals, along with some other coding seed markers for ease of identification. With the markers in place, orthogonal films are usually taken and these are used to digitize the seed locations into the treatment-planning computer. The catheter buttons and washers may also be radio-opaque for easy visualization on radiographs. More details are included in another chapter in this monograph. For LDR implants, the 192Ir seed ribbons are cut to the required length (i.e., the planned number of seeds) and are inserted and secured in the catheters. For HDR implants, the open catheter ends are covered with protective caps that can be easily removed at treatment time.
Catheter Seating: Buttons and Clamps A variety of button, washer, and clamp types have been introduced over the years, and a number are still available. These include various sizes, shapes, and materials, often designed for patient comfort and to minimize the intrusion of the button or clamp into the patient’s skin. Some examples are shown in Figure 2. For conventional radiographic localization it is often convenient that the skin position be made visible in radiographs by the use radio-opaque buttons, washers, or clamps. For CT-based treatment planning, however, such radiographic contrast may cause harmful artifacts.
Blind-End Catheters Situations sometimes arise in which the interstitial source guides cannot be passed completely through the tissue, so that one end of the catheter or needle must remain inside the patient. Examples of such cases include brain implants, interstitial gynecological implants, and some deep organ brachytherapy treatments.
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Figure 2. Button end (top) and blind end (bottom) catheters, washers, and buttons for securing catheters.
Special source guides have been developed to accommodate these circumstances. An example of these is shown in Figure 3.
Freehand Versus Template Implants The use of templates to guide the needles along the desired paths is expected to yield an implant in which the source locations are closely representative of the preplanned locations (Figure 3). This technique is common in gynecological and prostate implants, in which the path of the needles is easily predicted and fairly reliably followed. In some other treatment sites such as head and neck and breast, anatomical constraints may present problems that are more easily addressed with the flexibility of a freehand implant. Such constraints may be addressed to some extent with the use of a custom template, although this will lead to a higher cost and longer planning time prior to the implant. In some techniques the template is left in place for the duration of the dose delivery, while in others the template may be removed after the source guides are positioned. Some critics of template implants claim that they are associated with a greater degree of discomfort for the patient, as tissues placed under pressure by the needles will not be relaxed when the needles are released. The question of freehand versus template implants is still under debate at many centers.
Dose Homogeneity Dose Homogeneity Indicators Source guide placement is a critical step in the process of interstitial brachytherapy treatment. Even with the high degree of flexibility in dwell times available with HDR, a poor source guide placement can obviate the possibility of acceptable target coverage with good dose uniformity. In general, placement of source guides tightly inside the target volume will lead to unacceptably high doses in the central interior of the target, while placement well outside the target may lead to unnecessary treatment of surrounding healthy tissues to high doses. For characterization of dose uniformity, several dose homogeneity indicators have
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Figure 3. Template with blind-end catheters (Syed-Neblett design).
been introduced. The method of Paul (Paul et al. 1986) and others (Saw and Suntharalingam 1988; Wu, Ulin, and Sternick 1988) is to rate dose uniformity in terms of the relative volume of tissue treated to doses between the reference dose and some acceptable upper limit, usually 150% of the reference dose. The Radiation Therapy Oncology Group (RTOG) 95-17 breast brachytherapy protocol (Kuske, Bolton, and Hanson 1998) made use of a homogeneity indicator more closely related to the concepts of the Paris system (Pierquin et al. 1978), defining a mean central dose as an average minimum dose between the implant planes, in the central plane normal to and bisecting the source ribbons (Figure 4). The ratio of the prescription dose to the mean central dose was taken as the homogeneity index, and for an acceptable implant, this was expected to be greater than or equal to 75%. Some measure of dose uniformity should be adopted at all active clinics, and each implant evaluated against the in-house dose homogeneity standard. For further discussion on the topic, see chapter 25 in this monograph on post-procedure evaluation.
Source Placement and Dose Uniformity Some simple guidelines can help in effective source guide positioning. For example, the distance between seeds or dwell positions should be kept at a practical minimum. For simple point-source Quimby-like (uniform activity distribution) implants it has been shown that as the seed number in a planned implant decreases (and seed activity correspondingly increases by a factor f), the total volume enclosed in the highdose bubbles surrounding each source will increase by approximately the square root of the factor f (Zwicker and Schmidt-Ullrich 1995). Direct calculations have shown that for biplanar 192Ir implants, increasing the intraplanar ribbon spacing from 1.0 to 1.5 cm can increase the high dose volume by roughly 30%. Hence any increase in the intraplanar seed or ribbon spacing, as is often desired in order to reduce the number of needle insertions, must be considered in the light of the higher degree of dose inhomogeneity which follows. For biplanar implants the interplanar separation is also critical to the achievement of good target coverage and dose uniformity (Zwicker, Schmidt-Ullrich, and Schiller 1985). Guidelines for interplanar
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Figure 4. Illustration of Mean Central Dose as average dose at the center of triangles connecting sources, in a plane normal to and bisecting the sources. (Adapted and reprinted from Hanson and Graves (1995), Figure 6, p. 371, with permission from Medical Physics Publishing, ©1995 American Association of Physicists in Medicine.)
separation have been given in the literature for both LDR and HDR implants, as summarized in the following section.
Planar Implant Preplanning: Source Positioning Quimby-like Implants For two-plane Quimby-like (Quimby 1932) implants, interplanar separations have been given for a range of target thicknesses, with the aim of guaranteeing coverage of the target along the symmetry axes. It has been shown that these guidelines lead to implants with dose homogeneity that is optimal in some sense. The details of these guidelines are contained in the publications of Zwicker, SchmidtUllrich, and Schiller (1985) and Zwicker and Schmidt-Ullrich (1995) and are described in the previous chapter on Quimby-based brachytherapy systems.
Manchester Implants For simple biplanar HDR implants planned to give a Manchester-type (Paterson and Parker 1949) dose distribution (uniform dose in a plane at a fixed distance from an implant plane), the optimal interplanar separations has been shown (Zwicker and Schmidt-Ullrich 1995) to stand in a simple relationship to the target thickness T: S = T/1.414.
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An idealized implant constructed according to this guideline will show a highly uniform dose in the central plane parallel to the implant planes, and that dose will be the same as the peripheral prescription dose. One problem with the Manchester approach is that it leads to a prescription isodose surface that is highly rectangular in shape, which seems a somewhat unnatural fit to the basically amorphous volume of disease typically encountered. In view of this, a careful preplanning of each implant should be carried out, making use of all the information available to characterize the target size and shape. In such cases, implant guidelines such as that given above can be expected to aid in predicting the result of any proposed source guide positioning scheme, but would not in general be rigidly applied to every implant.
References Hanson, W. F., and M. Graves. “ICRU Recommendations on Dose Specification in Brachytherapy” in Brachytherapy Physics. J. F. Williamson, B. R. Thomadsen, and R. Nath (eds). Madison, WI: Medical Physics Publishing, pp. 361–378, 1995. Kuske, R. R., J. S. Bolton, and W. Hanson. (1998). RTOG 95-17: A Phase I/II Trial To Evaluate Brachytherapy as the Sole Method of Radiation Therapy for Stage I and II Breast Carcinoma. Philadelphia: Radiation Therapy Oncology Group 1998:1–34. Neblett, D. L. “Clinical Techniques and Applications Available for Interstitial Implantation” in Brachytherapy Physics. J. F. Williamson, B. R. Thomadsen, and R. Nath (eds). Madison, WI: Medical Physics Publishing, pp. 281–300, 1995. Paterson, R., and H. M. Parker in Radium Dosage: The Manchester System. W. J. Meredith (ed.). Baltimore, MD: Williams and Wilkins Company, pp. 6–38, 1949. Paul, J. M., R. F. Koch, P. C. Philip, and F. R. Khan. (1986). “Uniformity of dose distribution in interstitial implants.” Endocuriether Hypertherm Oncol 2:107–118. Perez, C. A., R. Zwicker, and J. Williamson. “Clinical Applications of Brachytherapy. I. LDR and PDR” in Principles and Practice of Radiation Oncology. C. A. Perez, L. W. Brady, E. C. Halperin, and R. K. Schmidt-Ullrich (eds.). Philadelphia: Lippincott, Williams and Wilkins, pp. 538–603, 2003. Pierquin, B., A. Dutreix, C. H. Paine, D. Chassagne, G. Marinello, and D. Ash. (1978). “The Paris system in interstitial radiation therapy.” Acta Radiol Oncol 17:33–48. Quimby, E. H. (1932). “The grouping of radium tubes in packs and plaques to produce the desired distribution of radiation.” Am J Roentgenol 27:18. Saw, C. B., and N. Suntharalingam. (1988). “Reference dose rates for single- and double-plane Ir-192 implants.” Med Phys 15:391–396. Schmidt-Ullrich, R., R. D. Zwicker, A. Wu, and K. Kelly. (1991). “Interstitial Ir-192 implants of the oral cavity: The planning and construction of volume implants.” Int J Radiat Oncol Biol Phys 20:1079–1085. Wu, A., K. Ulin, and E. S. Sternick. (1988). “A dose uniformity index for evaluating Ir-192 interstitial breast implants.” Med Phys 15:104–107. Zwicker, R. D., and R. Schmidt-Ullrich. (1995). “Dose uniformity in a planar interstitial implant system.” Int J Radiat Oncol Biol Phys 31:149–155. Zwicker, R. D., R. Schmidt-Ullrich, and B. Schiller. (1985). “Planning of Ir-192 seed implants for boost irradiation to the breast.” Int J Radiat Oncol Biol Phys 11:2163–2170.
Chapter 21
Advanced 3-D Planning Jean Pouliot, Ph.D., Étienne Lessard, Ph.D., and I-Chow Hsu, M.D. Department of Radiation Oncology, UCSF Comprehensive Cancer Center San Francisco, California Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 393 Conventional Brachytherapy Planning (Catheter-based Planning) . . . . . . . . . . . . . . . . . . . . . . . . . . 393 3-D Imaging in Brachytherapy (Anatomy-based Planning) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 395 Advanced 3-D Planning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 395 A Change of Paradigm (Inverse Planning) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 395 Intensity-Modulated Brachytherapy (IMBT) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 396 Treatment Inverse Planning Optimization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 396 Dose Distribution . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 397 Dose Constraints . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 397 Surface and Volume Dose Constraints . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 398 Flexible Definition of the Clinical Objectives . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 401 Applicable to All Anatomical Sites Accessible to Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . 401 Multiple Targets, Each with its Own Dose Prescription . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 401 Multiple Organs at Risk, Each with its Own Dose Prescription . . . . . . . . . . . . . . . . . . . . . . . . . . 402 Facilitate Adjustment Based in Specific Clinical Situation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 402 Class Solution . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 403 Clinical Implementation and Implications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 403 Prostate HDR Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 404 Comparison Between GEO and IPSA . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 404 Protection of the Penile Bulb . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 405 MRS-defined Boost . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 405 Number of Catheters . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 406 Gynecological Cancer . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 408 Interstitial Gynecological Implant . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 408 Vaginal Boost . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 408 Different Choice of Weighting Factors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 409 Permanent Prostate Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 409 Forward vs. Inverse Planning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 409 Variation of Seed Activity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 409 Boost of Positive Biopsy Areas . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 410 Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 412 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 412
Introduction Conventional Brachytherapy Planning (Catheter-based planning) The last decade has seen major changes in the way radiation treatments are delivered. The century-old objective of radiation therapy, to deliver a curative dose to the target while preserving normal tissues, can now be aimed at with a high degree of sophistication. However, in spite of major improvements achieved with three-dimensional (3-D) imaging modalities that allow the anatomy to be properly defined, brachytherapy is slowly taking advantage of these important new pieces of information. The insertion of the seeds for permanent implants or of catheters for afterloading treatment is now image guided, improving greatly the localization of the applicators. At the dose planning stage however, not so long ago the
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(a)
(b) Figure 1. Evolution of the type of information presented to the planning team over the years. (a) Traditional orthogonal pair of x-ray films showing the dummy catheters. (b) CT or MRI based planning where the anatomy and the catheters are defined in 3-D.
anatomy was only indirectly taken into account. For high dose rate (HDR) brachytherapy, catheters were placed in the target volume and it was assumed that if the dose distribution covers the catheters, it should also cover the anatomy. Imaging was commonly used to set the treatment margins, but optimized dose distributions were simply based on considerations such as catheter positions and desired dose to a few defined points. This necessarily resulted in an approximation of the shape of the anatomy. For the case of treatment of the prostate, volume optimization resulted in a dose distribution that was essentially cylindrical in shape. This may represent an “average” prostate, but each prostate has a very specific shape. Any approximation of its shape to a geometrical representation results in an overdosage of normal tissues surrounding the prostate. The organs at risk in the vicinity of the prostate may be included in this unnecessarily overdosed region. What do you see? The information provided to the team performing the dose planning has dramatically changed over time, both in nature and in quality. This evolution is illustrated in Figure 1. In Figure 1a, a conventional pair of orthogonal x-ray films is used for the planning of an interstitial HDR prostate implant. Dummy ribbons with radio-opaque markers are inserted in each catheter. The dosimetrist must digitize the dwell positions (after identifying which dummy sources belong to which catheter), and determine which of these dwell positions will be activated to cover the treatment length specified by the physician. With the exception of few patient dose points, a Foley catheter for the bladder, a rectum radioopaque marker, etc., the emphasis is placed on the catheters. Where are they, how many, how long should they be activated? This is why most, if not all, of the optimization tools developed were catheter based; i.e., they would produce a dose distribution closely related to the distribution of catheters. Because most of the anatomical information is unknown to the optimization tool, manual adjustments always follow to customize the plan.
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3-D Imaging in Brachytherapy (Anatomy-based Planning) 3-D imaging modalities, mainly ultrasound and computed tomography (CT), but also magnetic resonance imaging (MRI) (Ménard et al. 2004; Pouliot, Lessard, and Hsu 2004) and functional imaging such as magnetic resonance spectroscopy (MRS) (Pouliot, Lessard, and Hsu 2004) or positron emission tomography (PET), are rapidly being adopted as a standard for dose planning in brachytherapy. Similarly to external beam radiation therapy, the anatomical structures of interest are contoured and used to define planning target volumes (PTVs) and organs at risk (OARs). Here, PTVs and CTVs (clinical target volumes) are considered the same in brachytherapy. A discussion to differentiate them is beyond the scope of this chapter. Equipped with these new imaging modalities, the dosimetrist is now provided with both anatomical and catheter position information. This is illustrated in Figure 1b. The anatomical structures (prostate, rectum, urethra, bladder, etc.) are visible, in addition to the catheters. Therefore, when sources are placed, or dwell times adjusted, the dose distribution can be calculated and the user can assess the amount of dose received by the organs, also opening the possibilities to compute organ-specific dose volume histograms (DVHs). Most of the optimization tools are catheter based and therefore do not take into account the relation of the dose coverage and the protection of OARs during the optimization process. Yet, a number of modern tools have made the final dose distribution adjustment a lot easier than before. Examples of such tools are the ability to drag an isodose with the mouse while dwell times are adjusted or source positions are modified on the fly. A variety of dose-point optimization tools are also available, where dose can be optimized at certain points. All those tools however can only aim at one objective at a time, therefore relying on the user to judge what is an appropriate compromise simply based on visual assessment of the dose distribution and DVH. This approach is time consuming and the final treatment planning quality is often limited by the tight schedule of the clinical staff.
Forward or Conventional Planning A method of radiation treatment planning where the treatment parameters are first chosen and then the resulting dose distribution is calculated and evaluated.
Advanced 3-D Planning A Change of Paradigm (Inverse Planning) The inverse planning approach can be defined as a method of radiation treatment planning where one starts with the desired dose distribution, or clinical objectives, and then determines the treatment parameters that will achieve it. This is opposed to the conventional forward planning approach where the treatment parameters are first chosen and then the resulting dose distribution is calculated and evaluated. Since inverse treatment planning begins with the description of the desired dose distribution, it represents a change of paradigm in the planning process. CT or MRI contours of the CTV and OAR are used not only to define the anatomy for visual assessment and DVH calculation, but also to optimize the dose distribution. Therefore, they provide the physician with added flexibility and control to shape the dose distribution. In inverse planning, the anatomical features together with the dose constraints constitute the starting point of the dose optimization process. This requires that the multiple targets and OARs are known, and that the admissible and required dose coverage is specified. The computer knows the possible source positions because they need to be previously digitized, but from the user’s perspective, the emphasis is on the anatomy and the dose constraints. At the planning stage, the dwell times (for afterloader) or source
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positions (for permanent implants) become irrelevant to the user. This is different from forward planning where the dose distribution is iteratively adjusted by modifying the dwell times or the source positions until an acceptable dose distribution is produced. With the inverse planning approach, the dosimetrist works directly on the dose distribution and the compromise between target coverage, dose homogeneity, and OAR protection. This change of perspective brings the planning process closer to the real concern of the clinician, the dose delivered to the anatomy. The main benefit of the inverse planning approach is that all clinical requirements (dose coverage, dose homogeneity, OAR protection, etc.) are simultaneously and automatically taken into account in the planning process. As we move from the old two-dimensional (2-D) catheter-based planning to the new 3-D anatomy-based planning, the necessity of an inverse-planning approach becomes clear. With hundreds of source positions, irregularly shaped volumes, multiple target and organ volumes, multiple organ sensitivities, the chance of finding a treatment plan that would optimally satisfy all the requirements with a trial and error strategy becomes more and more remote. While the conventional forward planning approach may be adequate for uncomplicated cases, the adjustment of a dose distribution to respect different dose
Inverse Planning A method of radiation treatment planning where one starts with the desired dose distribution, or objectives, and then determines the treatment parameters that will fulfill it. Opposed to forward planning.
constraints on various targets and organs at risk in a reasonable time for clinical application is often beyond human capabilities. As the clinical expectation from brachytherapy escalates, treatment-planning requirements can only be met with an efficient approach such as inverse planning.
Intensity-Modulated Brachytherapy (IMBT) The concept of inverse planning for brachytherapy is the same as in external beam radiation therapy. The same three steps are required. First, imaging modalities must be used to provide 3-D anatomical structures. Then an inverse planning optimization tool, along with defined dose constraints, is used to determine the optimal dose distribution. Finally, a computerized delivery unit is used to deliver the complex intensity patterns, by controlling the leaf positions in external beam radiation therapy (EBRT) or the source positions and the dwell times in brachytherapy. Because of the close analogy between intensity-modulated radiation therapy (IMRT) and inverse-planned brachytherapy, the latter can be referred to as intensity-modulated brachytherapy (IMBT).
Treatment Inverse Planning Optimization The current treatment-planning optimization problem can be formalized as a combinatorial optimization problem. A combinatorial optimization problem is either a minimization or a maximization problem and can be described as a pair: a set of solutions and an objective function that assigns a quantitative value to each solution based on the optimization objectives. The solution space and objective function carry together a concise description of the problem. The goal is to find the optimal solution included in the solution space that possesses the optimal quantitative value from the objective function. The problem considered in the present case is the selection of the optimal treatment plan. The group of all possible treatment plans forms the solution space and the objective functions mathematically describe the ideal treatment
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plan based on clinical objectives. Following the inverse planning approach, the clinical objectives are described by means of dose constraints specified to each anatomical structure. The anatomical structures are represented by planar contours digitized by the physician on parallel cross-sectional images (CT or MRI) using a treatment-planning system. The algorithm uses these digitized anatomical structures to define in the 3-D space a set of dose points from which the dose distribution is evaluated based on the clinical objectives. The multiple objectives sought are to deliver the curative dose to all the targets, reduce the high dose regions inside the targets, and reduce the dose to the OARs and the normal tissues. The algorithm also required a set of 3-D positions from which the radioactive source may deliver the dose. The source dwell positions for the afterloader and the source template positions for permanent implant are also defined on parallel cross-sectional images using the treatment-planning system. Dose Distribution Based on the information provided by the treatment-planning system, the dose rate contribution from a given source position to a given dose point is calculated according to the AAPM task group 43 (TG 43) formalism (Rivard et al. 2004). The dose rate contribution is saturated for the source positions closer than 2 mm from the dose points to eliminate excessive contributions. The choice of this saturation distance defines the size of the high-dose bubbles within the target that are considered clinically relevant. This distance also takes into account the outer diameter of a typical catheter. The group of all dose rates from every source position to every dose point forms a dose rate matrix. In order to reduce the calculation time within the optimization iteration loop, this dose rate matrix is stored in a look-up table. The total dose Di delivered to a dose point i can be calculated from this look-up table and the set of dwell time values for the afterloader or the set of source positions for permanent implant. Dose Constraints The calculated dose distribution Di has to be translated into a quantitative value in order to measure its degree of fulfillment with the clinical objectives. The first step toward the evaluation of the dose distribution is to mathematically describe the physician’s requests by means of dose constraints. A dose constraint is a set of numerical values that defines the acceptable dose limits for a specific organ and their relative importance over the other clinical criteria. They convert the dose delivered to a dose point i into a penalty value Wi. This conversion is defined by the following penalty relation and illustrated in Figure 2.
m min ( Di − D min ) if Di < D min Wi = m max ( Di − D max ) if Di > D max if D min ≤ Di ≤ D min 0
(1)
The coefficients Dmin and Dmax represent the lower and the upper range of acceptable doses. If the dose is within the permissible dose range, the penalty is null. If the dose goes below or above the range, the penalty increases at rates mmin and mmax, respectively. Adjustment of the weights mmin and mmax sets the relative importance between the clinical criteria. The bigger the weight, the stronger is the penalty. The standard notation used to abridge the penalty relation is given below. mmin [Dmin Dmax] mmax
standard dose constraint notation
The advantage of using this type of penalty relation is that the translation from the physician’s expectations of the ideal dose distribution to a mathematical form is straightforward. The relation defines a clear border between acceptable and unacceptable doses and the physical meaning of dose constraints is clear.
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Figure 2. General form of a dose constraint.
The physicians have much clinical experience that can be translated into such dose constraints. In addition, this approach of defining the objectives induces an intuitive understanding of the optimization results. Moreover, this method is flexible and can describe different clinical objectives such as target minimal dose coverage and organ at risk protection. A set of dose constraints for HDR prostate brachytherapy is illustrated at Figure 3. In that case, dose constraints of two targets are defined, the prostate and a dominant intraprostatic lesion (DIL), and three OARs: the urethra, the rectum, and the bladder.
Dose Constraint A set of numerical values that defines the acceptable dose limits for a specific organ and their relative importance over the other clinical criteria.
Surface and Volume Dose Constraints In order to represent a wide variety of clinical objectives, the dose constraints applied to each digitized volume are divided into two types: one to evaluate the dose at the surface of the volume, the other to evaluate the dose inside the volume. In the case of a target, the first acts on dose points generated on the surface of the volume, thus forcing the dose distribution to be conformal to the volume. The second acts on dose points generated inside the volume to control the dose homogeneity. The following notation is used to describe the dose constraints of the each volume: on in
mmin [Dmin Dmax] mmax mmin [Dmin Dmax] mmax
surface dose constraints volume dose constraints
This set of dose constraints provides the ability to constrain the dose delivered to each volume independently and also permits defining different Dmin and Dmax limits on the surface and inside each volume.
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Figure 3. Set of dose constraints for HDR prostate brachytherapy including two targets, the prostate and a dominant intraprostatic lesion (DIL), and three organs at risk: the urethra, the rectum, and the bladder
Moreover, the different weights open the possibility to balance the clinical importance between low and high dose on the surface and inside each volume. All the weights defined over all volumes taken into account are relative to each other and correspond to relative clinical priorities. The convention is to define the mmin of the principal target surface as the reference weight and to set it at the arbitrary value of 100. From Local to Global Dose Prescription The concept of prescribing at a given point no longer applies for inverse planning; the prescription is global and based on the anatomy. Hence, the set of all dose constraints over all the digitized volumes taken into account is considered as the definition of this global prescription. Objective Function Once the dose constraints are set, the penalty values Wi are evaluated for each dose points i generated by the algorithm. Each dose point is associated with a specific anatomical structure (prostate, urethra, bladder, rectum, etc.) and a specific region of this anatomical structure (surface, volume) and so to a specific dose constraint. Finally, the sum of the penalty values Wi over all dose points i is performed to obtain the global penalty also known as the “cost function” or the “objective function.” This objective function mathematically describes the clinical objectives and is used to evaluate the quality of a given dose distribution. The closer the dose distribution is from the ideal dose distribution, the smaller the objective function value.
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The minimum value of this objective function corresponds to the optimal treatment plan as defined by the set of dose constraints. This objective function that simply considers the dose distribution is adequate for the afterloader for which the needles are implanted prior to the treatment planning being performed. At the planning stage, only the dwell times can be adjusted since the needles are already in place. However, for permanent implant the objective function should also include additional characteristics such as the number of needles that need to be implanted. Effectively, many plans may be clinically equivalent for a given dose distribution, but the one with the least number of needles should be favored. A previous study reported that minimizing the number of needles in the implant minimizes the edema associated with post-implant trauma (Speight et al. 2000). Hence, in order to complete the objective function, an additional term is included that constraints the number of needles. Optimization Engine Once the set of dose constraints is defined, an optimization engine is needed to minimize the objective function and converge toward the ideal treatment plan. The optimization engine is conceived to find the optimal dose distribution related to the clinical objectives among the extremely large number of possible solutions included in the solution space. Because the objective function is nonlinear and therefore may present multiple minima, an optimization technique able to avoid these local minima is required (Deasy 1997). The choice of this nonlinear objective function instead of a more simple linear objective function offers more flexibility in facilitating the description of sophisticated dose distribution. Different mathematical methods can be applied to minimize the objective function. They can be divided in two groups: deterministic methods and stochastic methods. Deterministic optimization algorithms travel downhill continuously on the objective function surface until a minimum is reached. Deterministic optimization algorithms have been applied for HDR brachytherapy planning optimization (Milickovic et al. 2002; Gabor, Streeter, and Astrahan 2003) and permanent implant planning optimization (Roy et al. 1991; Chen, Boyer, and Xing 2000). Although deterministic optimization algorithms are faster, they yield a solution at the nearest local minimum depending on the starting conditions and the best starting condition of the nonlinear objective function is unknown. Consequently, the solution may get trapped in a local minimum that is not the global or near the global minimum except by pure chance. An option to overcome this generic problem is to repeatedly execute the optimization several times using randomly selected starting conditions. This process generates a pool of solutions from which the final solution can be selected. However, this method does not guarantee that the pool of solutions contains the global or a near the global minimum. Moreover, the physician still has to select one treatment plan from a pool that may contain hundreds of treatment plans. Alternatively, stochastic optimization algorithms such as genetic algorithms (GA) or simulated annealing algorithms (SA) can process any form of objective functions, including nonlinear objective functions. Stochastic optimization algorithms apply a random search and therefore have the ability to overpass the local minima. A GA class algorithm applies the principles of natural selection for the computation of complex problems. A GA class algorithm is interesting in its own right and has been previously applied for HDR brachytherapy planning optimization (Lahanas, Baltas, and Zamboglou 1999, 2003; Yu et al. 2000) and permanent implant planning optimization (Yu and Schell 1996; Yang et al. 1998; Yamada et al. 2003). However, GA was not originally developed as an optimization algorithm (DeJong 1993), and GA does not offer any statistical guarantee of global convergence to an optimal solution (Forrest 1993). Nevertheless, it should be expected that GA might be better suited for some problems than SA (Ingber 1996). On the other hand, SA class algorithm does offer a statistical guarantee of global convergence to an optimal solution (Geman and Geman 1984; van Laarhoven and Aarts 1987; Aarts and Korst 1989). For that reason, an SA approach seems also to be a reasonable choice for the present combinational optimization problem. Previous investigations have shown that SA can be used to govern an optimization
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process to automatically and rapidly produce plans for prostate permanent implant treatments (Sloboda 1992; Pouliot et al. 1996, 1999; Taschereau, Roy and Pouliot 1999; Redpath 2002). The same concept has also been applied for prostate HDR brachytherapy (Lessard and Pouliot 2001). Because it takes some time and effort to become familiar with a given code, the ability to tune a given SA algorithm for use in more than one problem should be considered an important feature of an SA algorithm (Ingber 1996). Therefore, the SA algorithm previously developed for prostate HDR brachytherapy was extended to optimize any type of permanent and temporary brachytherapy (Lessard 2004). This thesis concluded that SA is suitable for treatment-planning optimization and produced reliable high-quality treatment planning. Treatment-planning optimization does not aim to find the mathematically optimal treatment plan, but to reach a clinically acceptable treatment plan in a clinically acceptable computation time. This work also demonstrates that an SA optimization engine carefully designed for a specific task can find an adequate solution in a very short time, opening the door to intraoperative treatment-planning optimization. The algorithm produced by this research is a comprehensive inverse planning algorithm for brachytherapy planning based on simulated annealing (IPSA). The next sections will outline the clinical and technical advantages brought along with the introduction of this new technology in the clinic.
IPSA An automated treatment-planning tool based on a constraint satisfaction approach (inverse planning) and solved by a stochastic optimization engine (simulated annealing).
Flexible Definition of the Clinical Objectives Applicable to All Anatomical Sites Accessible to Brachytherapy The sophisticated nonlinear objective function described in the previous section offers sufficient flexibility to describe multiple types of clinical objectives for multiple types of anatomical structures. This approach has provided an efficient way of defining a wide variety of dose distributions from a standard implant procedure to new sophisticated implant procedure ever considered before. Therefore, a supple objective function simplified the clinical introduction of the inverse planning in the clinic starting with standard treatment procedure such as interstitial prostate temporary implant and facilitated the creation of a more sophisticated treatment procedure such as interstitial gynecological temporary implant. Hence, a wide variety of anatomical sites, some illustrated at Figure 4, have been treated since the clinical introduction of IPSA: prostate (Lachance et al. 2002; Hsu et al. 2004); penis, gynecological (uterus, cervix, vaginal wall, and vulva) (Lessard, Hsu, and Pouliot 2002; DeWitt et al. 2004); rectum, soft-tissues (sarcomas), breast, neck, nasopharynx, base of tongue and for surface applicators (Taschereau et al. 2004). Independently of the clinical site and the complexity of the case, the algorithm produces a plan in a short time for clinical application, generally in less than 30 seconds. Moreover, these plans are equal (for simple cases), even better (for complex cases), than any plans produced manually by an experienced medical physicist. Henceforth, this inverse planning tool is used for every single HDR brachytherapy case at University of California San Francisco (UCSF), from very simple vaginal cylinders to complex interstitial implants. Multiple Targets, Each with its Own Dose Prescription The ability to define multiple targets, each with its own dose prescription, permits the definition of complex dose gradient within one tumor and/or between different tumors. Examples of targets within target would
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Figure 4. Examples of clinical sites optimized with IPSA.
be the boost (delivery of higher dose than the prescription dose) of cancer-validated areas in prostate defined by MRS (Pouliot et al. 2004) or of the positive biopsy areas. Multiple Organs at Risk, Each with its Own Dose Prescription The inverse planning approach always benefits from additional anatomical information without increasing the complexity of the planning process. One more advantage of inverse planning is that the burden on the physicist does not increase with the number of targets and OARs. This facilitates and therefore encourages the inclusion in the planning process of OARs that were neglected before. Examples of additional OAR are the penile bulb and the neurovascular bundles in the treatment of prostate cancer. Facilitate Adjustment Based in Specific Clinical Situation Another feature of this objective function is its ability to prioritize its objectives. For example, by changing the penalty weight, the inverse planning can produce a plan with a slightly more dose heterogeneity but improved target coverage or produce a plan with improved OAR protection but reduced target coverage. Because the optimization process is efficient and fast, the physician can review multiple plans before delivering the treatment. This facilitates the adjustment of the dose distribution based on specific clinical
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situations. Note that no manual adjustment of the source positions and dwell times is needed; a simple change in the relative priorities of the clinical objective will produce a new treatment plan that respects the new requirements. This approach is straightforward and much more intuitive and efficient than the conventional forward planning approach. Class Solution Standard treatment planning procedures for standard treatment cases make the planning process more efficient and promote consistency between plans produced for individual patients and by different planning staff. In the context of inverse treatment planning, such treatment-planning standards are known as “class solutions.” A class solution is defined as a set of dose constraints, or clinical objectives, relevant to one anatomical site that has been tuned to cover variations over a wide range of patients. The aim of a class solution is not to produce the optimal treatment plan per patient but to offer a set of dose constraints that would lead to the best outcome for a cohort of patients. The class solution can be used as a starting point for every patient, reducing dramatically the time needed to plan individual patients. The acceptance of the final treatment plan must be based on clinical information and the physician’s judgment. A class solution is not required for treatment inverse planning but recommended. The advantage of such an approach is that an acceptable plan can be produced in a reasonable time frame and ensures that all patients are treated in a very similar way, making their comparison more straightforward. The concept of class solution is commonly used in external beam inverse planned IMRT. The establishment of an IMRT class solution has been demonstrated to significantly improve planning efficiency while keeping high-quality treatment plans (Khoo et al. 2003; Xia et al. 2004). In addition, carefully developed IMRT class solutions have been found to be robust, requiring minimal fine-tuning on a patient-to-patient basis (Mott, Livsey, and Logue 2004). Class solutions need to be robust enough to produce dose distributions within suitable clinical tolerances regardless of individual patient anatomy. In inversely planned brachytherapy, such a set of dose constraints for a particular type of implant that will produce a clinically acceptable plan can also be identified. The same set of dose constraints has been used over the last 3 years at UCSF for inversely planned prostate HDR brachytherapy, and it has consistently produced very good plans. Other institutions have also experienced similar conclusions with HDR prostate brachytherapy (Lachance et al. 2002; Ménard et al. 2004). Despite the wide variety in organ volumes and shapes between patients, this class solution produced acceptable plans in all patients with excellent consistency in the dose distribution among patients. Rarely do the dose constraints need to be adjusted. This is very encouraging for the use of a class solution approach in brachytherapy.
Class Solution A set of dose constraints, or clinical objectives, relevant to one particular anatomical site but tuned to cover variations over a wide range of patients.
Clinical Implementation and Implications IPSA was included as a beta version in PLATO-BPS (Nucletron, Veenendaal, The Netherlands) in 2000 and has been clinically evaluated at UCSF1 (Hsu et al. 2004) and CHUQ2 (Lachance et al. 2002) for CT1 2
University of California, San Francisco, UCSF Comprehensive Cancer Center. Centre hospitalier universitaire de Québec, Hôtel-Dieu de Québec.
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guided HDR brachytherapy planning. This beta version was also introduced at NIH3 in 2002 for their MRIguided prostate HDR brachytherapy program (Ménard et al. 2004). This new technology has been successfully introduced in the clinic at these three institutions where several hundred patients have been treated with inversely planned HDR brachytherapy in the past 4 years. These independent clinical trials indicated that dosimetric indices and overall procedure time were improved with the introduction of IPSA for treatment planning. In addition to these three first North American sites, a first European institution (CAV4) specializing in pulse dose rate (PDR) brachytherapy is evaluating IPSA for this special purpose (Poupon, Castelain, and Lartigau 2003). More recently IPSA was also included in SPOT-PRO (Nucletron, Veenendaal, The Netherlands) and clinically evaluated at CHUQ in 2004 for intraoperative permanent prostate implant planning. In both cases (PLATO-BPS and SPOT-PRO), once the anatomical structures and the needle positions are digitized, the physician selects the class solution reflecting the clinical objectives and calls IPSA from a special button implemented in the treatment-planning system for the purpose of the clinical evaluation. The treatment-planning system gives the digitized anatomical structure and needle positions to IPSA, which performed the optimization and gives back the optimal treatment plan to the treatment-planning system. The physician finally evaluates the quality of the plan through visual assessment of the dose distribution and the DVH generated by the treatment-planning system. It is hoped that IPSA will be officially released in these treatment-planning systems and be accessible to the community in the near future. The following sections present several clinical applications and investigations of IPSA capabilities performed during the past 4 years. These clinical examples should clarify the new avenues available with this innovative treatment-planning tool and their clinical implications.
Prostate HDR Brachytherapy Comparison Between GEO and IPSA Geometric optimization (GEO) is a popular optimization tool available in commercial planning systems for HDR brachytherapy (Edmundson 1990). GEO determines the dwell time of a particular dwell position proportional to the sum of the inverse squared distances to other dwell positions. In a geometrically optimized implant, the dwell positions near the peripheral of the implant tend to get larger dwell times. The opposite is true for dwell positions located in the center of the implant. With this approach the dwell time values (or dwell weights in this case) are relative to each other. Afterward, the dwell weight values are normalized so that the prescribed dose covers the whole tumor volume. In general, GEO improves the coverage of the implanted volume and decreases hot spots within the implant. However, this optimization tool fails to use the anatomical information because it is only based on catheter positions. Therefore, time-consuming manual adjustments are required prior and after GEO to moderate the dose delivered to the OARs while keeping an acceptable coverage of the prostate. First, the final dose distribution obtained with GEO strongly depends on the selection of the dwell positions that contribute to deliver the dose. These dwell positions, known as the “active” dwell positions, are chosen from the available dwell positions within the implant prior to the GEO computation being performed. This is traditionally performed manually, based on visual assessment, keeping the dwell positions located inside or close to the tumor volume while excluding the dwell positions in the surrounding area of the OARs. The effectiveness of this manual method strongly depends on the physicist’s experience (Giannouli et al. 2000). In addition, even a meticulous selection of the active dwell positions may not be sufficient to produce the dose distribution wanted. Because GEO does not take into account the
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National Institute of Health hospital, Bethesda, MD. Centre Alexis Vautrin, Nancy, France.
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anatomy during the computation, it is frequent that manual dwell times are required after the GEO computation to reduce the dose delivered to the OARs and ensure an acceptable coverage of the prostate. Two independent clinical studies have concluded that IPSA produces superior treatment plans to GEO by improving coverage of the target while minimizing dose to normal structures (Lachance et al. 2002; Hsu et al. 2004). These studies also demonstrated that IPSA provides consistent results from patient to patient. Protection of the Penile Bulb It has been demonstrated that the dose of radiation received by the bulb of the penis correlates with the risk of impotence after radiotherapy for prostate cancer (Fisch et al. 2001; Merrick et al. 2001). In order to investigate if the dose received by the penile bulb could be significantly reduced, this anatomical structure has been contoured for several patients and included in the inverse planning process. For each case, IPSA was called twice; with and without taking into account this new volume. On average, the V50 of the penile bulb is reduced by 14% for a small penalty of 1% on the prostate V100. Moreover, the inclusion of this OAR eliminates all cases where the prescription isodose reached the penile bulb. This shows how the explicit inclusion of an additional OAR in the dose constraints can result in a substantial reduction of dose delivered to this organ at risk without a significant change of the target dose coverage. With MRI-based planning, the neurovascular bundles could be identified and contoured for inclusion as an organ to be protected. The irradiation of the penile bulb and the neurovascular bundles play a role in the loss of potency. Because of the proximity of the neurovascular bundles to the prostate, the level of protection might not be as significant as for the bulb. Constraints may nevertheless be defined to ensure prostatic coverage while avoiding undesirable hotspots to the neurovascular bundles. MRS-defined Boost The advent of functional imaging modalities such as PET or combined MRI/MRS imaging is an important step toward better cancer-validated tumor targeting in radiation therapy. MRS provides a non-invasive method of detecting small molecular markers (historically the metabolite choline, citrate, and polyamines) within the cytosol and extracellular spaces of the prostate. It has been demonstrated that the high specificity of MRS to metabolically identify cancer can also be used to improve the ability of MRI to identify the location and extent of cancer within the prostate (Kurhanewicz, Vigneron, and Nelson 2000). A study of 53 biopsy-proven prostate cancer patients prior to radical prostatectomy and step-section pathologic examination demonstrated a significant improvement in cancer localization to a prostatic sextant (left and right—base, mid-gland, and apex) using combined MRI/MRS versus MRI alone (Scheidler et al. 1999; Wefer et al. 2000). This information can be used to dose escalate only the cancer-validated area within the prostate also known as dominant intraprostatic lesions (DIL). Pouliot conducted a feasibility study using MRS guidance on 10 patients treated with HDR brachytherapy (Pouliot et al. 2004). MRI and MRS scans were obtained on a 1.5 Tesla system to determine the location of the DIL. Eight of the ten patients showed two distinct DIL. In all cases, the DIL were located in the peripheral zone of the gland. The population of patients included prostate volumes ranging from 30 to 45 cc. The DIL volumes ranged from 0.4 cc to 6.0 cc, representing 5% to 18 % of the prostate volumes. CT or MRI scans (five patients each) with the catheters in place were performed and transferred to the treatment planning system. The transversal CT/MRI slices were aligned manually to match the corresponding anatomical structures. In addition to the prostate, urethra, bladder, and rectum, the DIL were also contoured based on MRI and MRS. No restrictions on the relative distance between the DIL and the rectum or urethra were imposed in delineating the DIL contours. Therefore, on several occasions, the DIL was very close to the rectum. For each case, IPSA was used to optimize the dose distribution and to dose escalate different levels of DIL boosts. DVHs of the target and each OAR were computed and the results
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compared with optimized plans without DIL boost. The D1cc of both the bladder and the rectum show a small increase of 5% from no boost to the 150% boosts. The V120 of the urethra shows no increase for the 120% boost compared to no boost. However, increases of up to 50% of the urethra volume are observed for the 150% DIL boost. This indicates that DIL boosts of 120% can be safely delivered. Larger boost values up to 150% need more investigation. Besides, it could be argued that only the DIL plus a margin should receive the full-prescribed dose, allowing one to reduce the dose delivered to the rest of the prostate. This would result in a strong reduction of dose delivered to the surrounding tissues and OARs. However, better radiobiological data, perhaps gathered from functional imaging, will need to be obtained before such dose reduction could be considered. Number of Catheters A study focused on the number of catheters needed for a proper implantation of prostate using remote afterloader and GEO was recently performed (Charra-Brunaud et al. 2003). One could infer that it would be of value to minimize the number of implant catheters, both to minimize trauma to the patient and to simplify the implant procedure, while still obtaining the desired dose distribution. The extension of this study to compare the impact of the number of catheters on the dose distribution in prostate HDR brachytherapy for GEO and IPSA optimizations was recently presented (Pouliot, Lessard, and Hsu 2004). Transrectal ultrasound images of the prostate from 24 patients were transferred into the planning system. Urethra and prostate contours were digitized on each axial slice, as well as hypothetical locations of the catheters (2/3 of the catheters along the prostate contour, 1/3 around the urethra). Each patient data set was implanted with 9, 12, 15, 18, and 21 catheters (examples for 9, 15, and 21 are shown in Figure 5). For each case, two dose plans were generated: A GEO with local prescription to ensure a PTV V100 = 95%, and an IPSA with global prescription defined by a class solution. The dosimetrical indices were analyzed as a function of the number of catheters and categorized by prostate volumes. The dosimetrical indices variations relative to the number of catheters are presented in Figure 6. IPSA V100 are superior to GEO V100 for all number of catheters except for nine catheters (V100, IPSA = 92%, GEO = 95%). With GEO, an increase of the urethra V150 is observed (from 4 to 40%) when the number of catheters decreases from 21 to 9. With IPSA, the increase is only from 0 to 2%. For plans optimized with GEO, V150 gradually increases when the number of catheters decreases from 15 to 9 (p < 0.0001). The same trend is observed for small and large prostates. For IPSA plans, V150 is independent of the number of catheters. As one might expect from a dose distribution optimized on dose constraints, the constant target coverage and high dose level observed with IPSA also correspond to constant dose homogeneity. For GEO, the homogeneity index (HI) decreases from 0.69 to 0.49 when the number of catheters decreases from 21 to 9. Only a small HI decrease, from 0.69 to 0.65, is observed with IPSA. The average IPSA HI for 9 catheters is larger (better) than with GEO for 15 catheters. Therefore, IPSA can achieve dosimetrically equivalent treatment plans using a smaller number of catheters than when using GEO. The HI value is relatively constant with IPSA because the dose distribution homogeneity is controlled by the same set of dose constraints independently of catheter locations and their quantity. On the other hand, GEO is closely dependent on the catheter locations and therefore on their quantity. IPSA can also achieve a better dose distribution homogeneity with 21 catheters by changing the dose constraints. However, this is at the cost of the target coverage and OAR protection. One must clearly understand that the compromise between dose homogeneity and the other dosimetric indices is now adjustable with IPSA and therefore needs to be defined by the physician through the dose constraints. This new degree of freedom brings added flexibility for shaping the dose distribution but also poses questions about what should be the ideal shape of the dose distribution.
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Figure 5. Simulation of a prostate implanted with various numbers of catheters, from left to right: 9, 15, and 21 catheters.
Figure 6. Variation of the dosimetric indices of the prostate (V150, HI, COIN) and the urethra (V150) with the reduction of the number of catheters.
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For plans optimized with GEO, V150 increased when fewer catheters were used. Fifteen to twentyone catheters cover the prostate adequately without creating excess hot spots. For plans optimized with IPSA, for the same prostate coverage as GEO plans, smaller volumes of hot spots were observed, independently of the number of catheters between 9 and 21, and better urethra protection was achieved. Plans obtained with IPSA show reduced dosimetrical sensitivity variability due to changes in the number of catheters, providing additional flexibility to the physician for the implantation technique. IPSA appears to compensate for variation in catheter position tested in the study and makes brachytherapy less operator dependent.
Gynecological Cancer Interstitial Gynecological Implant The tandem-ovoid intracavitary system is an integral part of radiation therapy for brachytherapy of gynecological malignancies. The classic Fletcher-Suit intracavitary applicator system includes an intrauterine tandem and two intravaginal ovoids. Using the standard loading, this system produces a pear-shaped highdose region centered on the cervix. This brachytherapy system allows a very high dose to be delivered to the cervix while sparing adjacent bladder and bowel. The flexibility of this applicator allows it to be tailored to a variety of different patient anatomy types. However, for cases where the tumor extends beyond the range of the tandem-ovoid applicator, the insertion of additional catheters has the potential to deliver a curative dose to tumors located away from an accessible anatomical cavity (Hsu et al. 2002). The catheters are inserted around the vaginal wall, into the parametrial, paravaginal, or paraurethral regions, in addition to the vaginal tandem in an attempt to increase the dose to the region that is outside the standard pear-shaped dose distribution. This interstitial gynecological implant technique is promising but implies complex treatment planning. The number of catheters implanted and the implant geometry varies widely with the different tumor volumes and shapes. Conventional forward planning starting with geometrical optimization requires intensive adjustments to ensure tumor coverage and OAR protection. Moreover, because the implant is very different for each patient, it is impossible to develop a forward treatment-planning methodology applicable to each patient that will accelerate the planning process. Alternatively, the inverse planning approach applies the same dose constraints to each patient and produces consistent treatment plans independently of the number of catheter implanted and the implant geometry. The introduction of the inverse planning in our clinic facilitated the development of this new treatment procedure because it gives enough flexibility to the physician during the implantation and enough flexibility to the physicist to manipulate the dose distribution as wanted. Complex dose distribution based on complex implant geometry is achievable in a clinically acceptable time only with an inverse planning approach (Lessard, Hsu, and Pouliot 2002). Vaginal Boost In addition to the tumor volume and the OARs, a region around the vaginal tandem is contoured to specify to the algorithm to ignore the dose within the vaginal cylinder, as doses delivered to the nylon composing the cylinder are clinically irrelevant. The dose constraint defining the dose delivered within that volume is simply neglected using null weight penalties. In addition, it is the author’s belief that it is advantageous to maintain a high-dose region around the tandem in a fashion similar to standard tandemovoid applications. To ensure the high dose wanted at the surface of the cylinder, the minimal dose constraint at the surface is usually boosted to 120% of the prescription dose.
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Different Choice of Weighting Factors For the first gynecological implant performed at UCSF, the dose distribution was obtained with conventional geometric optimization followed by manual adjustment of the dwell times. In parallel, an inverse planning was performed with the same anatomical information. The set of dose constraints was determined such that the inverse planning would generate a dose distribution similar to the one produced with forward planning method. Afterward, several inverse plans were obtained with different emphasis given to the PTV coverage by varying the surface and volume minimal dose weight penalties from 70 to 500 (70 ≤ mmin ≤ 500) while the other dose constraint values were kept constant. DVHs were generated for the PTV and the OARs for all plans (Lessard, Hsu, and Pouliot 2002). The improvement in PTV coverage comes at the expense of an increased dose to the OARs. Also, while the coverage approaches 100% of the PTV, each OAR receives more doses. This is the old compromise of radiotherapy, to deliver a curative dose to the tumor while preserving surrounding normal tissues. Some inverse plans show improved dose coverage for an equivalent protection to OARs compared to the forward plan. The forward plan was performed by an experienced physician and took over an hour, while a graduate student produced all the inverse plans together in about 15 minutes. This illustrates the capacity of the inverse planning method to produce better treatment plans in a short time.
Permanent Prostate Implants Forward vs. Inverse Planning Radioactive seed implantation has become a popular alternative treatment option for men with localized prostate cancer. The success of this procedure depends on the proper selection of the seed distribution followed by the accurate placement of radioactive isotopes within a known volume of prostate cancer. Planning a 3-D seed array before implantation involves choosing a radioactive seed type and an appropriate activity that will adequately encompass the prostate gland while minimizing the dose to the surrounding adjacent structures. At UCSF, more than 750 prostate patients have been implanted with radioactive seeds since 1996. Over time, the clinical experience and post-implant analysis led to the establishment of a number of dosimetrical and technical rules that are followed to determine, using forward planning, the number of seeds and their localization within the prostate to rapidly produce a good dose distribution. Rules such as minimum target coverage with the prescribed dose, maximum dose to urethra, proximity of the seeds from the rectum wall, number of needles, etc., ensure that the prostate will be well treated and that the plan will remain effective even if seed misplacement or edema occur. With the advent of inverse planning, it was possible to generate a set of dose constraints and use IPSA to automatically generate plans that fulfill the rules established before. On Figure 7a, the experienced dosimetrist forward-planned a dose distribution following the rules. The dose distribution obtained with the inverse planning for the same prostate is shown on Figure 7b. The anatomical structures visible on this plane are the prostate and the urethra (in the center). The three isodoses displayed are the 100%, 150%, and 200%. IPSA consistently reproduces equivalent implants in terms of dosimetric indices and seed placement rules as the forwardplanned implants produced by an experienced physicist. In addition, the optimization is performed in a very short time, about 15 seconds for 400,000 iterations (2.5 GHz PC), opening the door to intraoperative treatment-planning optimization. Variation of Seed Activity In the case of a forward planning approach, changing the seed activity would force the dosimetrist to start over, just as if doing the plan for the first time. However, with inverse planning, once the dose constraints are specified, one can rapidly generate a new plan using a different seed activity. Figure 8 shows isodose
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Figure 7. Implant performed by an experienced physicist (a) and obtained with the inverse planning for the same prostate (b). The anatomical structures visible on this plane are the prostate and the urethra (in the center). The three isodoses displayed are the 100%, 150%, and 200%.
distributions for seed activities ranging from 0.37 to 0.9 mCi. The dosimetric indices are equivalent independently of the seed activities. The seed activity is directly taken into account in the objective function through the dose distribution calculation. As a result, the inverse planning distributes the seeds and determines automatically the number of needles and seeds needed with regard to the seed activity. Other studies indicate that good coverage can be obtained with seed activities ranging from 0.3 mCi to 0.7 mCi (Wuu et al. 2000; D’Souza and Meyer 2001; Sloboda, Pederson, and Halperin 2003; Sloboda et al. 2003). Furthermore, Taschereau showed that forward planning of the seed positions, even with low activity seeds, tends to be less robust to misplacement than plans generated by an inverse planning method with higher activity seeds (Taschereau, Roy, and Pouliot 2000). However, those authors also raised the question of dose homogeneity. The interplay between dose coverage, OAR protection, and dose homogeneity in the target volume has been an important question in external beam IMRT. In a number of clinical situations, the requirement of dose homogeneity can be relaxed to obtain better dose coverage and OAR protection. The same issue is now playing with the use of inverse planning in permanent implants. In a recent study (Beaulieu et al. 2004), plans were shown to be robust to misplacement and migration of seeds over a wide range of seed activity and for various seed models. With a properly tuned inverse planning algorithm able to ensure the dose coverage and OAR protection in the presence of placement errors, the choice of a preferred seed activity is open. The upper part of this range offers the opportunity to reduce significantly the number of seeds and needles, thus reducing the trauma to the patient, saving time in an operating room planning setting, and reducing the cost of the procedure. Boost of Positive Biopsy Areas The combination of an optimization algorithm with MRS information has proven to be a safe approach to boost positive biopsy areas with permanent implant (Zaider et al. 2000). We have demonstrated the feasibility of using IPSA to boost DIL with HDR prostate brachytherapy. To evaluate its ability to boost a region within the target for permanent prostate implants, small volumes were defined in addition to the prostate
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Figure 8. Implant obtained with the inverse planning for seed activities ranging from 0.37 to 0.9 mCi (a to f). The anatomical structures visible on this plane are the prostate and the urethra (in the center). The three isodoses displayed are the 100%, 150%, and 200%. The number of seeds counted in each implant is also displayed.
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Figure 9. Implant obtained with the inverse planning without dose constraints applied on these new volumes (a), with a boost of 150% with (b) and without (c) maximal dose limit. The anatomical structures visible on this plane are the prostate, the urethra (in the center), and three small volumes that simulate positive biopsy areas. The three isodose levels displayed are the 100%, 150%, and 200%.
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and the urethra already digitized by the physician over the transversal ultrasound images. Figure 9a presents the dose distribution for the optimization without dose constraints applied on these new volumes. Figures 9b and 9c correspond to two boosts of 150% with and without maximal dose limit. The anatomical structures visible on this plane are the prostate, the urethra (in the center), and three small volumes that simulate positive biopsy areas. The three isodose levels displayed are the 100%, 150%, and 200%. One can observe that the dose distributions are adequate for all optimizations. The inclusion of the boost to the positive biopsy areas clearly forces the 150% isodose to cover the regions. Further investigations are needed to establish a safe dose level for boost of positive biopsy areas in clinical applications.
Conclusion Modern planning tools, such as the anatomy-based IPSA presented in this chapter, efficiently use the information from the 3-D imaging modalities at all stages of the treatment procedure. IPSA automatically and rapidly produces conformal dose coverage of the target volume while limiting the dose to OARs in the delivery of HDR brachytherapy or permanent implants. With this inverse planning approach, the focus is on the physician’s prescription and dose constraints instead of on the technical limitations. Consequently, the physician’s control of the treatment is improved. With clear anatomical images, functional images to locate the cancerous tissues and rapid optimization to take all this information into account within the clinical time frame, plans can be generated with dose distribution that fulfills multi-objective dose constraints, sparing organs often not considered before, and personalized for each patient, a task inherently difficult to achieve with forward planning.
References Aarts, E. H. L., and J. H. M. Korst. Simulated Annealing and Boltzmann Machine. New York: John Wiley & Sons, 1989. Beaulieu, L., L. Archambault, S. Aubin, E. Oral, R. Taschereau, and J. Pouliot. (2004). “The robustness of dose distributions to displacement and migration of 125I permanent seed implants over a wide range of seed number, activity, and designs.” Int J Radiat Oncol Biol Phys 58:1298–1308. Charra-Brunaud, C., I.-C. Hsu, V. Weinberg, and J. Pouliot. (2003). “Analysis of the interaction between number of implant catheters and dose-volume histograms in prostate high-dose-rate brachytherapy using a computer model.” Int J Radiat Oncol Biol Phys 56(2):586–591. Chen, Y., A. L. Boyer, and L. Xing. (2000). “A dose-volume histogram based optimization algorithm for ultrasound guided prostate implants.” Med Phys 27:2286–2292. Deasy, J. O. (1997). “Multiple local minima in radiotherapy optimization problems with dose-volume constraints.” Med Phys 24:1157–1161. DeJong, K. A. “Genetic Algorithms Are NOT Function Optimizers” in Foundations of Genetic Algorithms: 2. E. D. Whitley (ed.). FOGA Proceedings July 24–29, 1992, Vail, Colorado. San Mateo, CA: Morgan Kaufman, pp. 5–18, 1993. DeWitt, K. D., I.-C. Hsu, V. K. Weinberg, É. Lessard, and J. Pouliot. (2004). “3-D inverse treatment planning for the tandem and ovoid applicator in cervical cancer.” Int J Radiat Oncol Biol Phys, In press. D’Souza, W. D., and R. R. Meyer. (2001). “Dose homogeneity as a function of source activity in optimized I-125 prostate implant treatment plans.” Int J Radiat Oncol Biol Phys 51:1120–1130. Edmundson, G. K. “Geometry-based Optimization for Stepping Source Implants” in Brachytherapy HDR and LDR. A. A. Martinez, C. G. Orton, and R. F. Mould, (eds.). Columbia, MD: Nucletron, pp. 184–192, 1990. Fisch, B. M., B. Pickett, V. Weinberg, and M. Roach. (2001). “Dose of radiation received by the bulb of the penis correlates with risk of impotence after three-dimensional conformal radiotherapy for prostate cancer.” Urology 57(5):955–959. Forrest, S. (1993). “Genetic algorithms: Principles of natural selection applied to computation.” Science 261:872–878.
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Gabor, J., O. E. Streeter, and M. A. Astrahan. (2003). “The use of linear programming in optimization of HDR implant dose distributions.” Med Phys 30(5):751–760. Geman, D., and S. Geman. (1984). “Stochastic relaxation, Gibbs distribution and Bayesian restoration images.” IEEE Trans Pattern Anal. Mac Intell (PAMI) 6:721–741. Giannouli, S., N. Milickovic, D. Baltas, M. Lahanas, N. Uzunoglu, C. Kolotas, and N. Zamboglou. (2000). “Autoactivation of source dwell positions for HDR brachytherapy treatment planning.” Med Phys 27(11):2517–2520. Hsu, I.-C., É. Lessard, V. Weinberg, and J. Pouliot. (2004). ”Comparison of inverse planning simulated annealing and geometrical optimization for prostate high-dose-rate brachytherapy.” Brachytherapy 3:147–152. Hsu, I.-C., J. Speight, J. Hai, É. Vigneault, T. Phillips, and J. Pouliot. (2002). “A comparison between tandem and ovoids and interstitial gynecologic template brachytherapy dosimetry using a hypothetical computer model.” Int J Radiat Oncol Biol Phys 52(2):538–543. Ingber, L. (1996). “Adaptive simulated annealing (ASA): Lessons learned.” Control and Cybernetics 25(1):33–54. Khoo, V. S., J. L. Bedford, S. Webb, and D. P. Dearnaley. (2003). “Class solutions for conformal external beam prostate radiotherapy.” Int J Radiat Oncol Biol Phys 55(4):1109–1120. Kurhanewicz, J., D. Vigneron, and S. J. Nelson. (2000). “Three-dimensional magnetic resonance spectroscopic imaging of brain and prostate cancer.” Neoplasia 2(1-2):166–189. Lachance, B., D. Beliveau-Nadeau, É. Lessard, M. Chretien, I.-C. Hsu, J. Pouliot, L. Beaulieu, and É. Vigneault. (2002). “Early clinical experience with anatomy-based inverse planning dose optimization for high-dose-rate boost of the prostate.” Int J Radiat Oncol Biol Phys 54:86–100. Lahanas, M., D. Baltas, and N. Zamboglou. (1999). “Anatomy-based three-dimensional dose optimization in brachytherapy using multiobjective genetic algorithms.” Med Phys 26(9):1904–1918. Lahanas, M., D. Baltas, and N. Zamboglou. (2003). “A hybrid evolutionary algorithm for multi-objective anatomybased dose optimization in high-dose-rate brachytherapy.” Phys Med Biol 48:399–415. Lessard, É. (2004). “Development and clinical introduction of an inverse planning dose optimization by simulated annealing (IPSA) for high dose rate brachytherapy.” Thesis abstract. Med Phys 31(10):2935. Lessard, É., and J. Pouliot. (2001). “Inverse planning anatomy-based dose optimization for HDR-brachytherapy of the prostate using fast simulated annealing algorithm and dedicated objective function.” Med Phys 28:773–779. Lessard, É., I.-C. Hsu, and J. Pouliot. (2002). “Inverse planning for interstitial gynecological template brachytherapy: Truly anatomy-based planning.” Int J Radiat Oncol Biol Phys 54(5): 1243–1250. Ménard, C., R. C. Susil, P. Choyke, G. S. Gustafson, W. Kammerer, H. Ning, R. W. Miller, K. L. Ullman, N. S. Crouse, S. Smith, É. Lessard, J. Pouliot, V. Wright, E. McVeigh, C. N. Coleman, and K. Camphausen. (2004). ”MRI-guided HDR prostate brachytherapy in a standard 1.5T scanner.” Int J Radiat Oncol Biol Phys 59(5):1414–1423. Merrick, G. S., K. Wallner, W. M. Butler, R. W. Galbreath, J. H. Lief, and M. L. Benson (2001). “A comparison of radiation dose to the bulb of the penis in men with and without prostate brachytherapy-induced erectile dysfunction.” Int J Radiat Oncol Biol Phys 50(3):597–604. Milickovic, N., M. Lahanas, M. Papagiannopoulou, N. Zamboglou, and D. Baltas. (2002). “Multiobjective anatomybased dose optimization for HDR-brachytherapy with constraint free deterministic algorithms.” Phys Med Biol 47:2263–2280. Mott, J. H., J. E. Livsey, and J. P. Logue. (2004). “Development of a simultaneous boost IMRT class solution for a hypofractionated prostate cancer protocol.” Br J Radiol 77:377–386. Pouliot, J., É. Lessard, and I.-C. Hsu. “Number of catheters in prostate high dose rate brachytherapy: The role of inverse planning.” ESTRO Joint Brachytherapy Meeting GEC/ESTRO-ABS-GLAC, Barcelona, Spain, May 13–15, 2004. Pouliot, J., D. Tremblay, J. Roy, and S. Filice. (1996). “Optimization of permanent 125I prostate implants using fast simulated annealing.” Int J Radiat Oncol Biol Phys 36(3):711–720. Pouliot, J., R. Taschereau, C. Coté, J. Roy, and D. Tremblay. (1999). “Dosimetric aspects of permanents radioactive implants for the treatment of prostate cancer.” Physics in Canada 55(2):61–68. Pouliot, J., Y. Kim, É. Lessard, I.-C. Hsu, D. B. Vigneron, and J. Kurhanewicz. (2004). “Inverse planning for HDR prostate brachytherapy used to boost dominant intraprostatic lesions defined by magnetic resonance spectroscopy imaging.” Int J Radiat Oncol Biol Phys 59(4):1196–1207. Poupon, L., B. Castelain, E. Lartigau. (2003). “Pulse dose rate brachytherapy: Optimization and place of imaging.” Cancer Radiather 7:136–146 (French).
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Redpath, A. T. (2002). “Automatic determination of needle and source positions for brachytherapy of the prostate using 125Iodine rapid strand.” Radiother Oncol 64:215–227. Rivard, M. J., B. M. Coursey, L. A. DeWerd, W. F. Hanson, M. S. Huq, G. S. Ibbott, M. G. Mitch, R. Nath, and J. F. Williamson (2004). “Update of AAPM Task Group No. 43 Report: A revised AAPM protocol for brachytherapy dose calculations.” Med Phys 31:633–674. Also available as AAPM Report No. 84. Roy, J. N., K. E. Wallner, S. T. Chiu-Tsao, L. L. Anderson, and C. C. Ling. (1991). “CT-based optimized planning for transperineal prostate implant with customized template.” Int J Radiat Oncol Biol Phys 21:483–489. Scheidler, J., H. Hricak, D. B. Vigneron, K. K. Yu, D. L. Sokolov, L. R. Huang, C. J. Zaloudek, S. J. Nelson, P. R. Carroll, and J. Kurhanewicz. “Prostate cancer: Localization with three-dimensional proton MR spectroscopic imaging—clinicopathologic study.” Radiology 213(2):473–480. Sloboda, R. S. (1992). “Optimization of brachytherapy dose distribution by simulated annealing.” Med Phys 19:955–964. Sloboda, R., J. E. Pedersen, and R. M. Halperin. (2003). “Is there a preferred strength for regularly spaced 125I seeds in inverse-planned prostate implants?” Int J Radiat Oncol Biol Phys 55:234–244. Sloboda, R., J. E. Pedersen, J. Hanson, and R. M. Halperin. (2003). “Dosimetric consequences of increased seed strength for I-125 prostate implants.” Radiother Oncol 68:295–297. Speight, J., K. Shinohara, B. Pickett, V. Weinberg, I.-C. Hsu, and M. Roach III. (2000). “Prostate volume change after radioactive seed implantation: Possible benefit of improved dose volume histogram with perioperative steroid.” Int J Radiat Oncol Biol Phys 48(5):1461–1467. Taschereau, R., J. Roy, and J. Pouliot. (1999). “Monte Carlo simulation of prostate implants to improve dosimetry and compare planning methods.” Med Phys 26(9):1952–1959. Taschereau, R., J. Roy, J. Pouliot, and D. Tremblay. (2000). “Seed misplacement and stabilizing needles in transperineal permanent prostate implants.” Radiother Oncol 55:59–63. Taschereau, R., P. Stauffer, I.-C. Hsu, J. L. Schlorff, A. J. Milligan, and J. Pouliot. (2004). “Radiation dosimetry of a conformal heat-brachytherapy applicator.” Technol Cancer Res Treat 3(4):347–358. van Laarhoven, P. J. M., and E. H. L. Aarts. Simulated Annealing: Theory and Applications. Dordrecht, The Netherlands: D. Reidel Publishing, 1987. Wefer, A. E., H. Hricak, D. B. Vigneron, F. V. Coakley, Y. Lu, J. Wefer, U. Mueller-Lisse, P. R. Carroll, and J. Kurhanewicz. (2000). “Sextant localization of prostate cancer: Comparison of sextant biopsy, magnetic resonance imaging and magnetic resonance spectroscopic imaging with step section histology.” J Urol 164:400–404. Wuu, C. S., R. D. Ennis, P. B. Schiff, E. K. Lee, and M. Zaider. (2000). “Dosimetric and volumetric criteria for selecting a source activity and a source type (125I or 103Pd) in the presence of irregular seed placement in permanent prostate implants.” Int J Radiat Oncol Biol Phys 47:815–820. Xia, P., N. Lee, Y.-M. Liu, I. Poon, V. Weinberg, E. Shin, J. M. Quivey, and L. J. Verhey. (2004). “A study of planning dose constraints for treatment of nasopharyngeal carcinoma using a commercial inverse treatment planning system.” Int J Radiat Oncol Biol Phys 59(3):886–896. Yamada, Y., L. Potters, M. Zaider, G. Cohen, E. Venkatraman, and M. Zelefsky. (2003). “Impact of intraoperative edema during transperineal permanent prostate brachytherapy on computer-optimized and preimplant planning techniques.” Am J Clin Oncol 26(5):130–135. Yang, G., L. E. Reinstein, S. Pai, Z. Xu, and D. L. Carroll. (1998). “A new genetic algorithm technique in optimization of permanent 125I prostate implants.” Med Phys 25(12):2308–2315. Yu, Y., and M. C. Schell. (1996). “A genetic algorithm for the optimization of prostate implants.” Med Phys 23:2085–2091. Yu, Y., J. B. Zhang, G. Cheng, M. C. Schell, and P. Okunieff. (2000). “Multi-objective optimization in radiotherapy: Applications to stereotactic radiosurgery and prostate brachytherapy” Artif Intell Med 19:39–51. Zaider, M., M. J. Zelefsky, E. K. Lee, K. L. Zakian, H. I. Amols, J. Dyke, G. Cohen, Y. C. Hu, A. K. Endi, C.-S. Chui, and J. A. Koutcher. (2000). “Treatment planning for prostate implants using magnetic-resonance spectroscopy imaging.” Int J Radiat Oncol Biol Phys 47(4):1085–1096.
Chapter 22
Optimization in Brachytherapy Gary A. Ezzell, Ph.D. Mayo Clinic Scottsdale Scottsdale, Arizona Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 415 General Concepts of Optimization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 415 Classes of Optimization Techniques . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 425 Concepts Underlying Two Stochastic Methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 426 Simulated Annealing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 426 Genetic Algorithms . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 427 Specific Brachytherapy Applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 428 Permanent Prostate Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 428 High Dose Rate Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 431 Summary . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 432 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 432
Introduction Mathematical optimization is a very wide field with a huge literature. A number of different techniques have been applied to brachytherapy, and to describe them all is beyond both my competence and time allotment. So, my plan here is to first discuss general concepts and principles of optimization, and then show examples of how some have been applied to brachytherapy. The discussion will not be particularly mathematical, but it will introduce some of the terminology used in papers related to optimizing radiation treatments. My goal is primarily to explain the most important concepts so that users of brachytherapy optimization software will find it less mysterious. Secondarily, readers with a deeper interest can use these concepts and terms to make the literature more accessible.
General Concepts of Optimization First, let us start with some specific applications to set the context. Consider prostate implants with permanent seeds. A common optimization problem is to decide which seed locations are to be loaded in order to meet some objectives related to target coverage, dose uniformity, and rectal/urethral sparing. Or, consider high dose rate (HDR) brachytherapy with variable dwell times. One might want to decide which dwell positions and times to use for a two-catheter endobrochial implant in order to deliver a sufficiently uniform dose at 1 cm from each of the catheters. Generalizing from these examples, the problem is always to design a distribution of source terms such that the resultant dose distribution satisfies certain constraints and meets certain objectives as well as possible. Let us discuss this in more detail and define some terms that are commonly used. In brachytherapy, the free variables are those elements of the problem that the planner can control. These may be source locations, source strengths, and/or dwell times. For low dose rate (LDR) prostate implants, the seed locations, as dictated by the template, can be selected, while the source strength may be fixed, as is the time, since the implant is permanent. For HDR treatment with implanted applicators, as for bronchus, breast, or cervix, the dwell locations within the applicator may be chosen as well as the dwell times, but the source strength is fixed. For some stereotactic brain implants, the location of each catheter may be optimized before the procedure. In this case the catheter locations are continuously variable, while the seed strengths may be chosen from an inventory. Each particular application has its own range of variables.
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Every problem also has its constraints. These may be hard constraints, meaning that they cannot be violated. Or they may be soft constraints, meaning that they may be violated but at some loss of plan quality. Constraints may be physical (e.g., dwell times cannot be negative) or clinical (e.g., “at least 98% of the target must be covered by the prescription dose”). Any solution that satisfies all of the constraints is feasible. Sometimes, finding a feasible solution is sufficient to meet the clinical need. More commonly, the goal is not just to satisfy the constraints, but also to find a solution that optimizes some objective. For example, “the dose to the surface of the prostate planning target volume (PTV) is to match the prescription dose as closely as possible.” This is generally stated as a minimization problem: minimize the variance of the doses Di at points i on the PTV surface from the prescription dose Dp. Minimize f = ∑( Di – Dp)2
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The optimization problem is to find the feasible solution that is optimal, that produces the lowest value of the objective function, f. Frequently, there is more than one clinical objective, and typically they compete with each other: “minimize the dose variation on the surface of the PTV” and “minimize the dose to the adjacent rectum.” Multiple objectives cannot, in general, be simultaneously optimized, but must in some fashion be balanced against each other. Although there are techniques for multi-objective optimization, most commonly the different objectives are combined into a single mathematical function, or objective function, that is to be minimized (or maximized). The simplest way to combine them is to sum the individual terms, each multiplied by a weighting factor, or importance factor. For example, the previous two objectives could be combined into a single objective function: Minimize f = w1∑(Di – Dp)2 + w2∑(Dj – Lr , if >0, 0 otherwise) .
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Here, the first sum is over the PTV surface points i and the second is over the rectal points j. (Note that the first term is squared, to penalize over- and underdosage of the target, while the second penalizes doses only if they exceed the rectal limit Lr .) Soft constraints and objectives are conceptually very similar and can be handled identically by incorporating appropriate terms in the objective function. A penalty based on the number of needles, N, might be added: Minimize f = w1∑(Di – Dp)2 + w2∑(Dj – Lr , if >0, 0 otherwise) + w3N .
(3)
As is perhaps becoming clear, the objective function is a mathematical model of the clinical goals. Whether or not the optimization process produces useful results depends, to a large extent, on how well that model captures the essence of the clinical thinking. Since no model is perfect, and no two clinicians think alike, the optimization system needs to allow the planner to influence the outcome of the optimization. In equation (3), the parameters that the planner could vary include the relative values of the weight factors, wi , and the dose objectives, Dp and Lr . The reader might object that the dose objectives are set by the physician, and so are not actually available as planning parameters. Nevertheless, it is useful to recognize that the dose objectives that the physician really wants may not be the best values to use as input to the optimization software. There are:
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(a) the dose objectives you actually want to achieve, (b) the dose parameters you give to the optimizer, and (c) the dose results you eventually accept. Ideally, these would all be the same. In some implementations, however, the user may not have tools like relative weights to adjust. In that case, the dose parameters become the tools used to steer the results in the desired direction. In order to illustrate these and some other concepts, let us look at an artificial problem, as shown in Figure 1. There are points 0–6 on the inside and outside of a target region, two sources, A and B, within the target, and a sensitive region above the target with a dose point R. To simplify the analysis, the “dose” is assumed to depend only on the source strength and the inverse square of the distance, so that a source strength of one unit produces one unit of dose at a unit distance. We have these simple relationships for the dose to the various points: D0 = A + B D1 = D5 = A + B/5 D2 = D4 = A/5 + B
(4)
D3 = A/9 + B D6 = A + B/9 DR = A/2 + B/2 . Let us take as our first dose objective that the target is to receive at least 5 units of dose. Our goal is to choose values of A and B that will satisfy that objective. It is clear from the geometry of the problem that point 0 will have the highest dose of the seven dose points specified and points 3 and 6 will have the least. If we plot on the [A,B] plane the equations defining these limiting doses, we can identify the region of solution space that meets this simple objective (Figure 2). Any combination of A and B in the shaded region will produce at least 5 units of dose in the target points 0–6.
Figure 1. Simple brachytherapy problem.
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Let us also assume that the dose to point R is not to exceed 5 units. Figure 3 shows that constraint plot. Any combination of A and B in the shaded region will produce no more than 5 units of dose to point R. Combining these target and sensitive region constraints identifies that region of solution space that would satisfy both types of objectives (Figure 4). The cross-hatched region shows the solutions that are feasible, giving at least 5 units of dose to all the target points and no more than 5 units to the sensitive structure point. (The scale of the axes has been changed to enlarge the feasible region on the plot.)
Figure 2. Solution space plot for the target dose constraints. The shaded region satisfies the constraints requiring all the target doses to be 5 units or more.
Figure 3. Solution space plot for sensitive region constraint. The shaded region satisfies the constraint that the dose to R be 5 units or less.
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Let us now make this an optimization problem and decide to minimize the dose to point R while still satisfying the target dose constraints. Graphically, reducing the acceptable value of DR shifts that line nearer the origin, leading to an optimal solution of A=B=4.5 units (Figure 5). It is possible, of course, to set up a problem with no feasible solutions. We cannot achieve a dose to point R less than 4.5 if we insist on the target points being at least 5. This simple example illustrates some other generalities. 1. Feasible solutions to sets of linear constraints are bounded by a convex multidimensional polyhedron. (This simple example is two-dimensional.)
Figure 4. Solution space plot for all the constraints showing the region of feasible solutions, those that give at least 5 units of dose to all the target points and no more than 5 units to the sensitive structure point.
Figure 5. Reducing the acceptable level for DR leads to an optimal solution at a vertex.
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Gary A. Ezzell 2. Inequalities define regions of the space. Equalities define surfaces of lower dimensionality. (For example, in Figure 3, DR4 cm), limited parametrial extension
20
30
8000
150
84–90
61–67
5600–9100
D
IIB/IIB bulky, extensive parametrial extension
20
40
8000
150
84–90
71–77
5600–9100
E
IIB, IIIB, IV, poor anatomy, poor regression
40
20
6500
150
92–94
69–74
4610–7410
(Reprinted from Williamson, J. F. “Clinical Brachytherapy Physics” in Principles and Practice of Radiation Oncology, 3rd ed. C. A. Perez, L. W. Brady, E. C. Halperin, and R. K. Schmidt-Ullrich (eds.). © 1998 with permission from Lippincott Williams & Wilkins.)
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Zuofeng Li insertion of sources into the patient. While not inherently a part of a treatment plan review, it is often convenient for the physicist to review the prescription, if used as the written directive for the treatment, to ensure that all required information is entered correctly.
Review of Technical Aspects of Treatment Plans A careful review of technical aspects of brachytherapy treatment plans typically constitutes the most timeconsuming part of a comprehensive treatment plan QA review. It is also the part of a treatment plan QA review that requires the expertise of a qualified physicist. Once satisfied that the treatment prescription is appropriate for the treatment under review, the physicist can now concentrate on the treatment plan itself. This review should address all milestones in the process of treatment plan creation, starting from patient data acquisition, be they volumetric or planar images, to the adequacy of final isodose plots. A checklist is often helpful for review of complex treatment plans, especially for HDR treatments, where the plan review must be completed in a compressed timeframe.
Patient Image Review A review of patient imaging serves multiple purposes in the process of treatment plan QA review. It assures that the applicator insertion is adequate for the treatment, that the images were acquired with minimal artifacts, and that the images are optimized for accurate source/applicator localization and dose calculation. The very first step in patient image review needs to confirm that the images belong to the correct patient. While this may seem a trivial check, the consequences of using the wrong patient image for treatment planning can be significant. The patient name, ID, and date of acquisition on the images must be verified to ensure that the correct image data set is used for treatment planning. Once that is done, the following aspects of image review can proceed. Applicator Insertion The physicist should review the placement of applicators and catheters relative to the patient’s anatomy to ensure that their locations are adequate for the desired treatment. The applicator insertion should allow adequate dose coverage of the treatment target, and minimize critical organ doses. For cervical cancer treatments, cervical markers are inserted into the cervix for identification on planar x-ray films. The optimal insertion of tandem and ovoid applicators requires that the ovoid surfaces to be at close proximity to these markers. As Katz and Eifel (2000) reported, a typical distance between the surfaces of the ovoids to the markers at M. D. Anderson Cancer Center is approximately 7 mm. Figure 1 shows a diagram of applicator geometry. Additional elements of a “good” tandem and ovoids implant include the symmetric placement of ovoids relative to the tandem both on the antero-posterior (AP) and on the lateral films, adequate distance of tandem to the sacrum and the pubis on the lateral film, and adequate packing to push the bladder and rectum away from the applicators. An inverted tandem insertion, indicated by the tandem curving toward the sacrum, may indicate perforation of the uterus (Jhingran and Eifel 2000). While patient-to-patient variations will occur, significant deviations from such typical “good geometry” should be brought to the attention of the treating physician. Such deviations may be due to the inadequate insertion of the applicators, or more likely due to the “slipping” of the applicator system caudally following insertion, and may indicate a need to re-insert and repack the applicators before patient films are acquired again for proper treatment planning. The dosimetry of interstitial implants similarly is dependent on the adequate insertion of treatment catheters, for both coverage of target volume and minimization of critical organ doses. Clips and markers are often placed in the tumor resection margin during the surgical removal of the tumor. They should
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Figure 1. Tandem and ovoids applicator system geometry. (Reprinted from Katz, A., and P. J. Eifel. (2000). “Quantification of intracavitary brachytherapy parameters and correlation with outcome in patients with carcinoma of the cervix.” Int J Radiat Oncol Biol Phys 48(5):1417–1425. © 2000 with permission from Elsevier.)
be used to help determine the adequacy of catheter insertion. Catheter separations or the distance between peripheral catheters and the target periphery should be evaluated, as they directly impact the quality of the implant in terms of both target coverage and dose inhomogeneity. Inadequate catheter insertion should be brought to the attention of the treating physician at the earliest possible time, so that remedial actions, such as insertion of additional catheters, may be taken. Catheters near critical organs, such as the rectum in gynecological interstitial implants, may need to be left unused to avoid excessive dose to this critical organ. Alternatively, “spacers” may need to be used in these catheters to allow adequate distances between the sources and the critical organs. The physicists performing treatment plan review should be familiar with all the applicator systems used in their institutions, especially in their physical dimensions and limitations. This is of particular relevance in HDR intraluminal treatments, such as endobronchial, esophageal, and bile duct treatments. As the HDR units typically have a limited range of distal-most and proximal-most treatable distances, the physicist should review the treatment planning films to ensure that the treatment target, as identified by the surgical markers, falls within this range, as can be identified by the x-ray markers inserted into the treatment catheters. Table 2 shows the maximum and minimum indexer lengths available in the Nucletron V2 HDR unit, together with the treatable target lengths and step sizes. Such mechanical limits are dependent on the HDR unit’s design and should be reviewed for each implant during patient imaging review. Quality of Images and Accuracy of Imaging Parameters Patient images, either volumetric or planar, need to be reviewed for their adequacy in allowing accurate applicator and catheter localization and dose calculations. For each type of brachytherapy implant, an imaging protocol should be developed and adhered to, with the parameters of these protocols chosen to minimize imaging artifacts, and to allow accurate applicator, target, and critical organ localization. Breathing artifacts in computed tomography (CT) scans or orthogonal radiographs will significantly
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Maximum Treatable Distance (from indexer faceplate) Minimum Treatable Distance (from indexer faceplate) Maximum number of dwell positions
Step sizes Gynecological transfer tube Flexible catheter transfer tube Stainless steel needle transfer tube
Nominal Values 1500 mm
Comment 4 mm additional catheter length for check cable test
725 mm 48
# of dwell positions multiplied by step size must be between min. and max. treatable distance
2.5 mm, 5 mm, 10 mm 1200 mm 1000 mm 1200 mm
increase the uncertainties in applicator and catheter localization accuracy, as well as in the delineation of treatment target and critical organs. Figure 2 shows an orthogonal pair of radiographs for the HDR treatment of bile duct, where the patient breathing motion artifact caused up to 1 cm differences in the y-coordinates of the x-ray markers. The physicist needs to evaluate the dosimetric consequences of such imaging artifacts, and communicate with the treating physician to arrive at the optimal actions to be taken in the patient’s treatment. When volumetric imaging is used for brachytherapy treatment planning, the field of view (FOV), slice thickness, table pitch, and gantry angle of the scans need to be reviewed to ensure that they are per the scanning protocol established for the particular brachytherapy treatment. Deviations from the established scanning parameters may reduce the accuracy of applicator, target, and critical organ localization, and could result in significantly increased dose calculation uncertainties. When orthogonal or other types of planar radiographs are used, the isocenter of the radiographs should be placed near the center of the target volume. The magnification factors of the radiographs and gantry angles need to be confirmed. The coordinate system used on the simulator, such as a Varian Ximatron unit, may be different from the International Electrotechnical Commission (IEC) coordinate system used in some treatment planning systems such as the Nucletron Plato TPS. The conversion of gantry angles between the two coordinate systems must be done correctly to avoid significant source localization errors. The filming technique used should allow clear and unambiguous recognition of applicators, surgical clips, and radiographic markers on the film. Target and Critical Organ Segmentation Image-based brachytherapy treatment planning has become common for some treatment sites, such as LDR and HDR transrectal ultrasound (TRUS) guided prostate brachytherapy, accelerated partial breast brachytherapy using either interstitial or MammoSite technique and gynecological intracavitary brachytherapy using CT/magnetic resonance (MR)-compatible applicators. The segmentation of treatment target volumes and critical organs for each patient may be performed by a radiation oncologist or a dosimetrist. The physicist reviewing the treatment plan should be familiar with the anatomy of the treated site, as well as the institutional and/or national treatment guidelines related to this part of treatment planning. The contours of segmented target and critical organs should be reviewed for both anatomical accuracy when applicable, such as the prostate, seminal vesicles, bladder, rectum for prostate implant, or the seroma and MammoSite balloon for breast implant. The expanded planning target volume (PTV) should be
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(b)
Figure 2. AP and lateral films of a bile duct HDR treatment. (a) Patient instructed to breathe in and hold breath; (b) No patient instruction.
reviewed for appropriate application of expansion margins and for overlaps with critical organs such as skin of the breast.
Review of Input Parameters to Treatment Plans Once satisfied that the images used to construct treatment plans are of adequate quality for the purpose of treatment planning, and that the applicators and catheters are inserted to allow the desired dose distribution, the physicist can concentrate on the correctness and accuracy of applicator and catheter localization and input into the TPS. Applicator/Catheter Geometry Parameters The dimensions of brachytherapy applicators, catheter indexer lengths, and their numbering directly have an impact on the correctness and the accuracy of the treatment plan. Ovoid and vaginal cylinder diameters need to be verified by comparison with documented values or direct measurement from the patient images. Ovoids larger than 2 cm in diameters are often constructed by sliding a plastic cap on 2-cm diameter ovoids, and are therefore often indistinguishable from the 2 cm ovoids. The diameters of such ovoids, as well as CT/MR-compatible vaginal cylinders, therefore must be documented in the operating room immediately following their insertion. The physicist reviewing the treatment plan should review such documentation and compare the values used in the treatment plan to ensure that they are correctly recorded and used in the treatment plan. Diameters of vaginal cylinders with metal identification rings can be measured off the radiographs, and are thus readily verifiable. Indexer lengths of HDR catheters must be reviewed carefully, with full understanding of the type of catheters used, for example, stainless steel needles or plastic flexible catheters. Depending on the transfer tubes used, treatment distance errors of up to 50 cm have occurred. Every attempt is warranted in avoiding such systematic errors that resulting in significant dose delivery errors and medical events.
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Source/Applicator/Catheter Localization Accuracy The input of sources and applicators/catheters are typically performed in treatment planning using coded x-ray markers or CT-compatible markers. The physicist reviewing a treatment plan needs to be familiar with the coding scheme of the x-ray markers, or the physical characteristics of CT-compatible markers, so that the distance information encoded in these markers can be correctly entered into the TPS. When interstitial or intraluminal brachytherapy treatments using multiple catheters are planned, the individual catheters should be numbered, such that the source loading and/or dwell position and time calculations individualized for each catheter can be correctly reproduced in every treatment fraction. The physical numbering of the catheters needs to be compared with the data entered into the treatment plan. If coded x-ray markers are used instead of numbering of treatment catheters, the x-ray markers must remain in the treatment catheters until source loading time. The physicist should be familiar with their appearances on the radiographs or CT images. Treatment planning systems often report source localization errors by comparing the reconstructed source length values with their expected values. These reports should be carefully reviewed to ensure that no significant errors have occurred during source localization. The use of spacer and the value of step sizes in the treatment plan require special attention. Their locations and values need to be compared with those indicated on the patient images, and double-checked prior to treatment delivery. Errors in their use in treatment planning can result in geographically displaced dose distributions, or potentially treatment of wrong site. Point of Interest/Critical Organ Localization Accuracy Points of interest are often used in brachytherapy treatment plans to represent prescription point or critical organ doses. They are commonly measured on orthogonal films or from volumetric patient images, and entered manually into the TPS for calculation of doses to these points. The localization of points of interest is of particular importance in gynecological tandem and ovoids treatment planning, as they may directly affect the overall delivery of the treatment. The classic Manchester system definition of point A measures level of point A superiorly from vaginal fornices (typically assumed to be the superior ovoid surfaces). Errors in localization of point A often result in greater than 10% error in delivered dose. While point B, rectum and bladder points are typically not used for brachytherapy treatment prescription, their dose values are sometimes used to optimize treatment plans such as changing the source loading pattern and dwell times to reduce rectum and bladder doses, or to assist in subsequent patient management, such as the need to include parametrium boost external beam radiation treatment. In the treatment planning for MammoSite breast treatment, the balloon diameter and its center may be measured off radiograph films. These measurements need to be reviewed for accuracy, as they are directly used for dose prescription and delivery. For intraluminal implants or planar or volumetric interstitial implants, the doses are often prescribed to points at a given distance to the catheter(s). The placement of these points again directly has an impact on the total dose delivered to the patient. Such points may be automatically generated, such as in the case of the Nucletron Plato TPS, or manually entered for LDR treatments. The physicist reviewing a treatment plan should fully understand the mechanism through which these points are determined and localized relative to the treatment catheters and patient anatomy, and determine its appropriateness as applied to the plan under review as well as its adherence to institutional protocols. Source Characteristics TPS acceptance testing and commissioning usually includes a complete testing of the source characteristics, including benchmark data, physical and active source dimensions, as well as single and multiple
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source dose distribution in a predetermined configuration. Treatment plan QA review therefore can concentrate on the particular sources used for a given patient treatment. Aspects of source characteristics review include verifying that the correct source model is used for the treatment plan, that the source decay was calculated accurately, and that the number of sources used in the treatment plan agrees with the prescription. Plan Optimization On a modern brachytherapy TPS, a number of automatic or semi-automatic plan optimization algorithms are typically available. While it is not the intention of this chapter to discuss these algorithms, a physicist reviewing a treatment plan, in which one of these algorithms is used, needs to have a detailed understanding of these algorithms. The selection of a particular algorithm should be specific to the goals of the treatment, in addition to keeping with institutional consistency. Automatic optimization algorithms most often are used for the preplanning of permanent prostate implants and various HDR treatment implants. For permanent prostate implants, the optimization algorithm may be a simple geometry model that places seeds at locations in or near the prostate, based on a set of geometry rules, such as distance between needles and seeds, columns to avoid, etc. Alternatively, the optimization algorithm may attempt to distribute the seeds based on dosimetric specifications in terms of target coverage and critical organ doses. The physicist should work with the treating physician in selecting an optimization algorithm for his or her institution’s treatment planning for prostate implants. Attempts should be made to use the same optimization algorithm and optimization parameters for all prostate implants. This effort in maintaining treatment planning consistency will help achieving more predictable treatment plans in the form of a relationship between prostate volume and the number and total activity of sources. Such information is critical in determining whether a plan “makes sense,” and should help maintain consistency of patient treatment. For HDR treatments, the selection of optimization algorithms and optimization parameters can directly impact the quality of the treatment plan as well. The Geometry Optimization algorithm (Edmundson 1990) and its variants are often used for interstitial planar and volume implants, while optimization based on dose points has been successfully used for gynecological implants (Stitt et al. 1992; Thomadsen et al. 1992). The placement of dose points relative to the applicators directly affects the resulting dose distribution. Physicists reviewing such treatment plans should pay special attention to how the dose points were placed. When dose volume histogram (DVH)-based optimization algorithms are used, the physicist should attempt to establish DVH optimization parameters for each treatment site, and subsequently review each treatment plan for its adherence to such protocols.
Plan Quality Evaluation The evaluation of a brachytherapy treatment plan quality includes the aspects of target coverage and normal organ doses, as well as dose homogeneity of the plan. Similar to external beam radiation therapy, the evaluation of the quality of a treatment plan is specific to a treatment site, as well as the delivery technique used in the treatment.
Point-based Plan Quality Evaluation Parameters Traditionally, brachytherapy plans based on two-dimensional radiographs do not have explicit quantitative representation of treatment target or critical organs beyond what can be estimated based on patient bony anatomy, surgical clips, or markers inserted into the patient, or reconstructed from Foley balloons and vaginal packing for gynecological treatments. When necessary, doses to these points can be calculated
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as surrogates to target coverage and critical organ doses. Significant over- or underdosage to those landmarks needs to be discussed with the treating physician. For interstitial implants, the International Commission on Radiation Units and Measurement (ICRU) (ICRU 1997) defines several parameters that are useful in the evaluation of the quality of a treatment plan. The report defines the central plane of an interstitial implant to be a plane through the major (long) direction of the implant. In practice, this often is taken to be the direction of the implant needles or catheters. For a complex implant, in which a single central plane cannot be defined with respect to all implanted catheters, subvolumes may be defined, such that a central plane may be defined for each subvolume. The mean central dose (MCD) is an extension of the basal dose of the Paris implant system, defined as the average of local minimum doses between the sources in the central plane(s) of an implant. For implants with parallel needles, the MCD can be determined geometrically as the midpoint between a group of neighboring needles. Figure 3 shows the definition of central planes and the corresponding MCD calculation for a complex interstitial implant. The Dose Homogeneity Index (DHI), defined as the ratio of the prescription dose to the MCD, has found frequent use in determining the whether a treatment plan is acceptable in terms of dose homogeneity.
Volume-based Plan Quality Evaluation Parameters When the three-dimensional volumetric dose distribution is calculated for a brachytherapy treatment plan, additional parameters may be defined for the plan quality evaluation. Dose distribution in the treated volume, which may be calculated with only the knowledge of source locations, has proven to be useful for dose homogeneity evaluations. When a set of volumetric images is available for a given patient, dose distribution to anatomically defined treatment target, as well as organs at risk (OARs), may be critically evaluated in the quality evaluation of a brachytherapy treatment plan. Given a dose distribution from implanted sources, the ICRU (1997) defines the high dose (HD) and low dose (LD) volumes, and the prescribed dose (PD) that help to evaluate the dose homogeneity and
Figure 3. A complex implant with two central planes defined. The Mean Central Dose (MCD) is the average of the doses at points A–G. (Reprinted from ICRU Report 58: Dose and Volume Specifications for Reporting Interstitial Therapy. Bethesda, MD: ICRU, © 1997 with permission from ICRU, Bethesda, MD.)
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coverage of an implant. Details on the definitions and applications of those quantities can be found in the chapter on ICRU interstitial reporting recommendations, and are not repeated here. A useful parameter in brachytherapy plan evaluation is the maximum contiguous dose, defined to be the minimum dose that encloses all sources in an implant (Neblett et al. 1985). This can be estimated on a TPS by plotting consecutively decreasing isodose clouds until an isodose cloud is seen to enclose all sources. This all-source-enclosing enclosing isodose cloud should cover the intended treatment target. Many other tools exist that help to describe quantitatively the quality of a brachytherapy plan, especially in regard to dose homogeneity. Anderson (1995) and Thomadsen (1999) described them in detail. As many of the brachytherapy treatment planning systems today are able to calculate DVHs of various types (cumulative, differential, natural), DVH values of the treated volume, i.e., the volume of tissue receiving a given dose level, are often used for evaluation of dose homogeneity of a treatment plan, while DVH values for a given segmented organ, either target or OARs, are useful for evaluation of the target coverage of the treatment as well as risk of treatment complications. A simple plotting of the volume of an organ (the sum of voxels) receiving doses in a given range produces a differential DVH is shown in Figure 4. The cumulative DVH is calculated from the differential DVH by summing all voxels receiving doses up to a given dose level, then plotting these “cumulative” volumes against the dose levels, shown in Figure 5. A disadvantage of differential or cumulative DVH plots for brachytherapy is that, due to the high dependence of brachytherapy dose distribution on the distance between the sources and the points of interest, all DVH plots will appear similar, as demonstrated in Figure 6, which shows the cumulative DVH plot for a single HDR dwell position treatment (MammoSite), compared to the cumulative DVH of an optimized HDR interstitial breast brachytherapy treatment of Figure 5. Anderson (1986) developed a form of DVH, the natural DVH, which mathematically removes the inverse-square-dependence of dose
Figure 4. Differential DVH plot of the treated volume of a breast brachytherapy treatment.
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Figure 5. Cumulative DVH plot of the treated volume of an optimized breast brachytherapy treatment.
Figure 6. Cumulative DVH of the treated volume of a MammoSite® treatment.
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distribution in a brachytherapy implant, resulting in a DVH plot as shown in Figure 7. Note that a peak appears near the prescription dose of 340 cGy. In comparison, for a MammoSite breast brachytherapy treatment using a single dwell position (effectively a point source), the natural DVH appears like a flat line, as seen in Figure 8. Various parameters can be calculated from a natural DVH plot, as seen in Figure 7, that may help comparative evaluation of a treatment plan. While it is difficult to generalize the acceptable values of these parameters, in general, sharper peaks and rapid fall-off in the high-dose side of a natural DVH indicate more homogeneous dose distribution. Saw and Suntharalingam (1991) described the dose non-uniformity ratio (DNR), defined as the ratio of a high-dose volume and the total volume of tissue receiving the prescription dose. As the magnitude of high-dose volume may be related to complications such as soft tissue necrosis, the value of DNR needs to be minimized. When volumetric imaging is used, such as in prostate or breast brachytherapy, the target coverage and critical organ doses of the treatment plan should be evaluated based on existing institutional protocols or national guidelines. The Coverage Index (CI), defined as the ratio of PTV/CTV (clinical target volume) receiving the prescribed dose to the volume of the PTV/CTV, is typically used to describe target coverage. Significant over- or underdosage to the target or the critical organs may require modification to the plan. The physicist performing plan review should have intimate knowledge of those protocols and guidelines, and be ready to recommend revision of a treatment plan or suggest remedial action. The use of DVH plots for brachytherapy plan evaluation can be largely empirical. Similar to external beam DVH plots, acceptable DVH values depend on the specific disease being treated, as well as the treatment technique adopted. For permanent prostate implants, the AAPM Task Group 64 (TG 64) report (Yu et al. 1999) recommended a set of DVH values. Stock et al. (2002) reported that keeping the dose that covers 90% of the PTV, D90, between 140 Gy to 180 Gy appears to be optimal, with D90 < 140 Gy
Figure 7. The Natural DVH plot of the optimized interstitial breast brachytherapy implant shown in Figure 6.
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Figure 8. Natural DVH of a MammoSite® breast brachytherapy treatment.
associated with increased biochemical failure, and D90 > 180 Gy with a small increase in long-term urinary symptoms. For interstitial breast HDR treatments, Arthur et al. (2003b) recommended a set of parameters extracted from the DVH of the treated volume for evaluation of HDR accelerated breast brachytherapy treatments, as shown in Table 3.
Independent Plan Calculation Check An immediate question that arises after all the detailed and tedious review of the plan for a given brachytherapy treatment is why perform an independent plan check calculation? The answer to this question determines the complexity of the check calculation process and procedure. The author submits the following goals for an independent brachytherapy plan calculation check: 1. Prevention of catastrophic data input errors. Such errors may be due to misidentification of dose prescription points, incorrect data transfer, or misinterpretation of treatment prescription.
Table 3. Accelerated Partial Breast Brachytherapy Plan Evaluation Criteria Volume of Treated Tissue Receiving 150% of Prescription Dose (V150)
Volume of Treated Tissue Receiving 200% of Prescription Dose (V200)
Dose Volume Ratio
50 cc
20 cc
0.75
V150 1 − V 100
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2. Prevention of unexpected software errors. Even though a TPS has been subject to rigorous testing during its commissioning and acceptance testing process, it is conceivable that certain parts of the software were never tested, or the use of a specific function was never anticipated. The complexity of a modern brachytherapy TPS is such that it is impossible to test all possible functions and the sequences in which these functions are utilized to arrive at a treatment plan. As treatment planning systems are often designed with a certain operation sequence in mind, any deviation from the intended sequence may result in unexpected results. 3. Regulatory compliance. The NRC requires that all brachytherapy calculations, either manually performed or computer-generated, must be checked. While the NRC does not specify how such checks should be performed, a quantitative calculation check would appear to satisfy this requirement. An independent calculation check, therefore, should be designed to quickly detect significant/catastrophic errors, while not significantly slowing down the treatment planning and delivery process. This is especially important in HDR brachytherapy treatments, where everything must happen in a compressed time frame.
Tools for Independent Calculation Checks Many tools have been available, proposed, or can be modified for use for independent brachytherapy plan checks. Given the strengths or dwell times or the sources and their spatial distributions, an independent plan calculation check scheme calculates dose delivered to a given point. Considering that such algorithms have been available before the arrival of computerized treatment planning, it is no surprise that many traditional dose calculation formalisms or implant systems have found youth in the independent calculation check of brachytherapy plans. In addition to performing a gross dosimetry accuracy check, they may help provide an insight into the relationship of source strengths and resulting doses and their points of specification, compared to computerized number manipulations. An independent plan calculation check can be done, in the more complex form, using a second brachytherapy TPS or software systems specifically developed for this purpose. Saw et al. (1998) described the use of an LDR brachytherapy TPS for the independent calculation check of HDR brachytherapy treatment plans. The source positions are transferred into the LDR TPS via digitization, with appropriate scaling of LDR source strengths to reflect the varying dwell times of HDR source at the active dwell positions. The authors stated that all calculations between the HDR system and the LDR system agreed to within 10%, with 80% of them agreeing to within 5%. Such a scheme, by its completeness, helps to identify human errors in data input as well as potential algorithmic errors in the HDR system in the aspect of dose prescription point placement, coordinate system translation and rotation, and dwell time weight optimization. The authors stated that each independent calculation check took less than 20 minutes. Cohen et al. (Cohen, Amols, and Zaider 2000) and Lachaine et al. (Lachaine, Gorman, and Palisca 2003) developed their own in-house programs to perform independent dose calculations for HDR treatments, either based on the source position input into the HDR TPS (Cohen et al.) or on the HDR treatment unit control file exported from the HDR TPS (Lachaine et al.). Cohen et al. observed a typical discrepancy of up to 3%, while Lachaine et al. saw maximum differences of 2%, compared with the HDR TPS. It should however be noted that simply transferring the already-localized source and point of interest coordinates into a second TPS and performing a second dose calculation will only verify the dose calculation algorithm accuracy, but may not detect the arguably more serious and frequent errors in source localization. At the other end of the spectrum, classical dose calculation algorithms, such as away-and-along tables for cesium tube sources, Sievert integral (Williamson 2003) for unfiltered line sources, Paterson-Parker system (Williamson 2003) for planar and volume implants (with appropriate accounting for modern source
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strength units and correction factors, such as shown in Table 4), and prostate implant nomograms (Anderson 1976; Anderson, Moni, and Harrison 1993; Wang and Potters 2001), serve to provide quick and often as accurate estimates of the relation between total source strength and prescription dose rates. Other methods (Mayo and Ulin 2001; Rogus, Smith, and Kubo 1998; Ezzell 1994, 2000a,b; Miller, Davis, and Horton 1996; Das et al. 2004; Cohen et al. 2002) have been developed that facilitate the quick manual calculation for specific brachytherapy treatment sites. They are typically easy to use and achieve adequate accuracy that, depending on the application, may substitute for computerized treatment planning.
Independent Calculation Check of Gynecological Intracavitary Implants Gynecological intracavitary implants are often performed using tandem and ovoid applicators or cylinder applicators, either using cesium tubes for LDR applicators or HDR remote afterloaders. At Mallinckrodt Institute of Radiology, the loading patterns of tandem ovoids follow strict rules depending on the length of the tandem and diameter of the ovoids. Computer-generated dose values for points A, B, and P (at 6 cm away from patient midline and level of point A) can be compared to manual calculations based on the calculated, standard values in Table 5. The average of the dose values at left and right points A, B, and P, generated by the TPS, typically are within 10% of the predicted values. While the table was calculated for LDR implants, HDR treatments at Mallinckrodt Institute of Radiology have maintained the same loading pattern by weighting the dwell times accordingly. The table therefore is useful for both LDR and HDR tandem and ovoid treatments. Mayo and Ulin (2001) described a method for checking the treatment time calculation for HDR vaginal cylinder treatments. For dose prescription points located at 5 mm away from cylinder surface in the vaginal apex region, tapering off to cylinder surface at points along the tranverse vaginal apex, the authors proposed the determination of a scaling factor K, that relates the prescribed dose D, the source strength Sk , and the total treatment time TT, in the form of TT = K × D/Sk .
(1)
Given that the dose prescription points are at varying distances from the vaginal cylinder surface, the authors arrived at an equation that described K as a function of vaginal cylinder diameter, the length of the cylinder with prescription point at 5 mm away from cylinder surface, as well as the length of the cylinder with prescription points on cylinder surface. The authors observed a maximum discrepancy of 5% of their independent calculation from the HDR TPS calculations.
Independent Calculation Check of Intraluminal Treatments Intraluminal brachytherapy treatments, such as endobronchial, esophageal, and bile duct treatments, typically use no more than two treatment catheters with long active lengths. When a single catheter of relative little curvature is used for the treatment, such as is often the case for esophageal treatments, the Sievert unfiltered line source integral may be used. Let L be the active length, Sk the total source strength, Λ the single-source dose rate constant, x the away distance of the dose calculation point away from the bisector of the line source, and T the total treatment time, then the dose D at this point can be approximated by D = 2×T ×
Sk × Λ L×x
L . 2 × x
tan −1
(2)
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Table 4. Paterson-Parker System Table for Plane and Volume Implants Volume implantsPlanar implants Volume (cm3) (cm2)
mgRaeq·h* 1000 ‘P-P’R
Minimum dose/ IRAK† cGy/(µGy·m2)
Area (cm2)
mgRaeq·h* 1000 ‘P-P’R
Minimum dose/ IRAK† cGy/(µGy·m2)
1 2 3 4 5 10 15 20 25 30 40 50 60 70 80 90 100 110 120 140 160 180 200 220 240 260 280 300 320 340 360 380 400
34 54 70 85 99 158 207 251 291 329 398 462 522 579 633 684 734 782 829 919 1005 1087 1166 1242 1316 1389 1459 1528 1595 1661 1725 1788 1851
3.49 2.20 1.68 1.38 1.194 0.752 0.574 0.474 0.408 0.361 0.298 0.257 0.228 0.206 0.188 0.174 0.162 0.152 0.143 0.129 0.118 0.110 0.102 0.0958 0.0904 0.0857 0.0815 0.0779 0.0746 0.0716 0.0690 0.0665 0.0643
0 2 4 6 8 10 12 14 16 18 20 24 28 32 36 40 44 48 52 56 60 64 68 72 76 80 84 88 92 96 100 120 140 160 180 200 220 240 260 280 300
30 97 141 177 206 235 261 288 315 342 368 417 466 513 558 603 644 685 725 762 800 837 873 908 945 981 1016 1052 1087 1122 1155 1307 1463 1608 1746 1880 2008 2132 2256 2372 2495
3.97 1.23 0.844 0.672 0.578 0.506 0.456 0.413 0.378 0.348 0.323 0.285 0.255 0.232 0.213 0.197 0.185 0.174 0.164 0.156 0.149 0.142 0.136 0.131 0.126 0.121 0.117 0.113 0.109 0.106 0.103 0.0910 0.0813 0.0740 0.0682 0.0633 0.0593 0.0558 0.0527 0.0502 0.0477
1000 ‘P-P’R, 1000 Manchester system roentgens; IRAK, integrated reference air-kerma. * Original Manchester values from Paterson, R, Parker HM: A dosage system for interstitial radium therapy. Br J Radiol 11:313–339, 1938. † Modified from original values for 192Ir assuming 8.6 Gy minimum peripheral dose per 1000 ‘P-P’R and 7.227 µGy·m2/mgRaeq·h. (Reprinted from Williamson, J. F. “Clinical Brachytherapy Physics” in Principles and Practice of Radiation Oncology, 3rd ed. C. A. Perez, L. W. Brady, E. C. Halperin, and R. K. Schmidt-Ullrich (eds.). © 1998 with permission from Lippincott Williams & Wilkins.)
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Zuofeng Li Table 5. Table for Tandem and Ovoid Plan Manual Calculation Check Used at Mallinckrodt Institute of Radiology. HDR Treatment Dwell Times Are Scaled to Achieve Identical Loading Patterns. Applicator Component
Small tandem Medium tandem Standard tandem Endometrial tandem 2.0 cm colpostats* 2.5 cm colpostats* 3.0 cm colpostats* Mini-ovoids (1.6 cm colpostats)
Dose Rate (cGy·h-1) per mgRaeq+
Loading (mgRaeq)
Point A
Point B
Point P
20 10-20 10-10-20 10-20-10 20-20 25-25 30-30 10-10
1.545 1.543 1.070 1.308 0.553 0.474 0.418 0.660
0.295 0.297 0.260 0.278 0.250 0.244 0.228 0.255
0.205 0.207 0.185 0.195 0.183 0.182 0.173 0.190
Multiply above numbers by total mgRaEq in each component. + 3M 137Cs tubes, 1.4 cm active length. No correction for decay. *Includes 6% applicator attenuation correction.
Modification of the Sievert integral for HDR treatments is necessary to account for significant variations of source dwell times. In a typical HDR optimized single catheter treatment plan, the dwell times at the ends of the catheters are often significantly higher than those in the middle section of the catheter, while those in the middle section remain mostly constant. The Sievert integral, in the form above, can then be combined with point source dose calculations for the dwell positions at the ends of the catheters, to arrive at a dose estimate often within 5% of the TPS calculations. Rogus et al. (Rogus, Smith, and Kubo 1998) investigated the relationship of catheter length and treatment distance for HDR single catheter treatments. Assuming straight-line or moderately curved catheters, the authors proposed a fitting equation for the total treatment time t as shown below:
(
)
t ( d , L )ref = −1.35 + 7.74 d + 0.322 d 2 +
L − 50 50
( −0.591 + 6.92 d + 0.0230 d ) , 2
(3)
for the reference condition of a prescription of 500 cGy and a source strength of 10 Ci, where d is the distance of prescription point away from the catheter, and L is the active length of the catheter. This equation is simply scaled to apply to other prescription doses and source strengths. Ezzell (2000a) studied the influence of catheter curvature on dose calculation accuracy based on straight-line assumption such as outlined above, and concluded that the ratio of the distance between end dwell positions (the cord) and the active length of a single, curved HDR treatment catheter can be used as a measure of the catheter curvature. Figure 9 shows the cord-to-active-length ratio required to maintain dose homogeneity within 10% (Ezzell 2000a). For treatments using two catheters, Miller et al. (Miller, Davis, and Horton 1996) and Ezzell (2000b) described methods to verify their treatment times. Miller et al. eliminated 50% of the active dwell positions located in areas where the two catheters overlap, such that the two catheters can be considered independently as if no overlap occurs. Ezzell’s formalism actually calculates the expected dwell time for each dwell positions, based on the distance of a dwell position on a given catheter to the second catheter.
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Figure 9. Cord length relative to active length required to maintain 10% homogeneity for a single, curved catheter. (Reprinted from Ezzell, G. A. (2000). “Limitations of the straight-line assumption for endobronchial HDR brachytherapy treatments.” Med Phys 27(1):151–153. © 2000 with permission from AAPM.)
A sample Microsoft Excel spreadsheet is available that can be used for performing these calculations, available at ftp://ftp.aip.org/epaps/medical_phys/E-MPHYA6-27-008005/example_worksheets.pdf.
Independent Calculation Check of Planar Interstitial Implants In general, the Paterson-Parker system applies well to the independent calculation check of planar implants. The width and the length of the implant can be estimated from orthogonal radiographs of the implant, or measured from a TPS display, with the orientation of the implant rotated to facilitate the measurement. The area of the implant thus measured can then be used for table lookup on a standard Paterson-Parker planar implant table, such as presented by Williamson (2003). The resulting mgRaeq·hrs for the implant should then be corrected for the elongation factor of the implant, and compared to the mgRaeq·hrs calculated by the brachytherapy TPS for the implant, following appropriate unit conversion. It is important to keep in mind that the original Paterson-Parker system’s prescription dose for planar and volume implants is the so-called “modal-dose,” or the dose at 10% higher the minimum dose in the surface that contains the dose prescription point. If the intention is to cover the implant volume by at least the prescription dose, the mgRaeq·hrs looked up from such a table would need to be increased by 10% to be comparable with TPS calculations. Ezzell (1994) reviewed 66 rectangular HDR planar implants, and described a method to estimate the total treatment time required to deliver a prescription dose from the planar implant. Let I be the Dose area index, defined as I=
Dose × area Source strength × Total time
,
(4)
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then I = A(T) + B(T) × E + C(T) × E2 ,
(5)
where T is the thickness of treatment, E is the length of the equivalent square of implant. For source strength specified in cGy cm2 s-1, Ezzell gave the calculations of the fitting coefficients A, B, and C as follows: A = 3.245 – 1.269 × T + 0.1014 × T2 B = 1.030 – 0.0728 × T
(6)
C =-0.02083 + 0.001925 × T . Testing of this method on real patients showed that it mostly predicts the total treatment time to within 10%.
Independent Calculation Check of Volume Interstitial Implants Volume interstitial implants can typically be easily checked by use of the Paterson-Parker tables (Williamson 2003). Das et al. (2004) reported this use for accelerated partial breast brachytherapy treatments, where better than 7% agreement was observed when corrections were made for elongation factor, and the original Paterson-Parker table was corrected for modern units. The volume of tissue treated to the prescription dose is looked up on the DVH plot calculated by the HDR TPS. When a DVH is not available, the implanted volume can be estimated by measuring the width, length, and height of the volume bounded by the implanted catheters. The Paterson-Parker system was designed for radium sources, which are minimally attenuated in tissue. For low-energy sources such as 125I and 103Pd, used for permanent prostate seed implant, Paterson-Parker system is no longer applicable. Cohen et al. (2002) reported the use of nomograms for the independent calculation check of prostate seed implants. Let davg be the average distances between pre-planned needles/seeds in the lateral, anterior-posterior and superior-inferior directions; the authors reported the following nomograms: davg + 0.8 = 1.524 1.09 U cm Sk
2.2
(7)
for 125I seed implants with a prescription dose of 144 Gy, treated to the prostate volume with a 5 mm margin in all directions except the posterior, and davg + 0.8 = 5.395 1.09 U cm Sk
2.56
(8)
for 103Pd seed implants with a prescription of 140 Gy. Compared with computerized treatment plans generated using a genetic algorithm, the authors reported agreement of better than 10% in the total activity required.
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Conclusions This chapter describes briefly the goals and methods in the QA review of brachytherapy treatment plans. As modern brachytherapy treatment plans become more complex, and as the use of brachytherapy becomes more widespread, physicists are challenged constantly to come up with efficient and accurate methods to quickly evaluate the quality of a brachytherapy treatment plan. The physicist reviewing a treatment plan needs to remain vigilant in performing this important task.
References Anderson, L. L. (1976). “Spacing nomograph for interstitial implant of 125I seeds.” Med Phys 3(1):48–51. Anderson, L. L. (1986). “A ‘natural’ volume-dose histogram for brachytherapy.” Med Phys 13:898–903. Anderson, L. L. (1995). “Dose Specification and Quantification of Implant Quality” in Brachytherapy Physics. J. F. Williamson, B. R. Thomadsen, and R. Nath (ed.). Madison, WI: Medical Physics Publishing, 1995. Anderson, L. L., J. V. Moni, and L. B. Harrison. (1993). “A nomograph for permanent implants of palladium-103 seeds.” Int J Radiat Oncol Biol Phys 27(1):129–135. Arthur, D. W., F. A. Vicini, R. R. Kuske, D. E. Wazer, and S. Nag. (2003a). “Accelerated partial breast irradiation: An updated report from the American Brachytherapy Society.” Brachytherapy 2(2):124–130. Arthur, D. W., D. E. Wazer, D. Koo, L. Berle, L. Cuttino, M. Yunes, M. Rivard, D. Todor, S. Tong, T. Tenenholz, and T. DiPetrillo. (2003b). “The importance of dose volume histogram evaluation in partial breast brachytherapy: A study of dosimetric parameters.” Int J Radiat Oncol Biol Phys 57(2):S361–S362. Beyer, D., R. Nath, W. Butler, G. Merrick, J. Blasko, S. Nag, and C. Orton. (2000). “American brachytherapy society recommendations for clinical implementation of NIST-1999 standards for (103)palladium brachytherapy.” Int J Radiat Oncol Biol Phys 47(2):273–275. Cohen, G. N., H. I. Amols, and M. Zaider. (2000). “An independent dose-to-point calculation program for the verification of high-dose-rate brachytherapy treatment planning.” Int J Radiat Oncol Biol Phys 48(4):1251–1258. Cohen, G. N., H. I. Amols, M. J. Zelefsky, and M. Zaider. (2002). “The Anderson nomograms for permanent interstitial prostate implants: A briefing for practitioners.” Int J Radiat Oncol Biol Phys 53(2):504-511. Das, R. K., R. Patel, H. Shah, H. Odau, and R. R. Kuske. (2004). “3D CT-based high-dose-rate breast brachytherapy implants: treatment planning and quality assurance.” Int J Radiat Oncol Biol Phys 59(4):1224–1228. Edmundson, G. “Geometry Based Optimization for Stepping Source Implants” in Brachytherapy HDR and LDR. A. A. Martinez, C. G. Orton, and R. F. Mould (eds.). Columbia, MD: Nucletron Corporation, 1990. Ezzell, G. A. (1994). “Quality assurance of treatment plans for optimized high dose rate brachytherapy—planar implants.” Med Phys 21(5):659–661. Ezzell, G. A. (2000a). “Limitations of the straight-line assumption for endobronchial HDR brachytherapy treatments.” Med Phys 27(1):151–153. Ezzell, G. A. (2000b). “A manual algorithm for computing dwell times for two-catheter endobronchial treatments using HDR brachytherapy.” Med Phys 27(5):1030–1033. Gaspar, L. E., S. Nag, A. Herskovic, R. Mantravadi, and B. Speiser. (1997). “American Brachytherapy Society (ABS) consensus guidelines for brachytherapy of esophageal cancer.” Int J Radiat Oncol Biol Phys 38(1):127–132. Grigsby, P. W., J. F. Williamson, and C. A. Perez. (1992). “Source configuration and dose rates for the Selectron afterloading equipment for gynecologic applicators.” Int J Radiat Oncol Biol Phys 24(2):321–327. International Commission on Radiation Units and Measurements (ICRU). Report 58: Dose and Volume Specifications for Reporting Interstitial Therapy. Bethesda, MD: ICRU, 1997. Jhingran, A., and P. J. Eifel. (2000). “Perioperative and postoperative complications of intracavitary radiation for FIGO stage I-III carcinoma of the cervix.” Int J Radiat Oncol Biol Phys. 46(5):1177–1183. Katz, A., and P. J. Eifel. (2000). “Quantification of intracavitary brachytherapy parameters and correlation with outcome in patients with carcinoma of the cervix.” Int J Radiat Oncol Biol Phys 48(5):1417–1425. Lachaine, M. E., J. C. Gorman, and M. G. Palisca. (2003). “A fast, independent dose check of HDR plans.” J Appl Clin Med Phys 4(2):149–155.
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Mayo, C. S., and K. Ulin. (2001). “A method for checking high dose rate treatment times for vaginal applicators.” J Appl Clin Med Phys 2(4):184–190. Miller, A. V., M. G. Davis, and J. L. Horton. (1996). “A method for verifying treatment times for simple high-doserate endobronchial brachytherapy procedures.” Med Phys 23(11):1903–1908. Neblett, D. L., A. M. N. Syed, A. A. Puthawala, R. Harrop, H. S. Frey, and S. E. Hogan. (1985). “An interstitial implant technique evaluated by contiguous volume reference.” Endocuriether/Hypertherm 1:213–221. Nag, S. (2000). “Brachytherapy for prostate cancer: summary of American Brachytherapy Society recommendations.” Semin Urol Oncol 18(2):133–136. Nag, S., A. A. Abitbol, L. L. Anderson, J. C. Blasko, A. Flores, L. B. Harrison, B. S. Hilaris, A. A. Martinez, M. P. Mehta, and D. Nori. (1993). “Consensus guidelines for high dose rate remote brachytherapy in cervical endometrial, and endobronchial tumors.” Int J Radiat Oncol Biol Phys. 27(5):1241–1244. Nag, S., P. E. Cole, I. Crocker, S. K. Jani, K. V. Kishnasastry, V. Massullo, R. Nath, D. Nori, S. Parikh, P. Rubin, B. Speiser, P. S. Teirstein, P. Tripuraneni, R. Waksman, and J. F. Williamson. (1999a). “The American Brachytherapy Society perspective on intravascular brachytherapy.” Cardiovasc Radiat Med 1(1):8–19. Nag, S., D. Beyer, J. Friedland, P. Grimm, and R. Nath. (1999b). “American Brachytherapy Society (ABS) recommendations for transperineal permanent brachytherapy of prostate cancer.” Int J Radiat Oncol Biol Phys 44(4):789–799. Nag, S., W. Bice, K. DeWyngaert, B. Prestidge, R. Stock, and Y. Yu. (2000a). “The American Brachytherapy Society recommendations for permanent prostate brachytherapy postimplant dosimetric analysis.” Int J Radiat Oncol Biol Phys 46(1):221–230. Nag, S., B. Erickson, B. Thomadsen, C. Orton, J. D. Demanes, and D. Petereit. (2000b). “The American Brachytherapy Society recommendations for high-dose-rate brachytherapy for carcinoma of the cervix.” Int J Radiat Oncol Biol Phys 48(1):201–211. Nag, S., B. Erickson, S. Parikh, N. Gupta, M. Varia, and G. Glasgow. (2000c). “The American Brachytherapy Society recommendations for high-dose-rate brachytherapy for carcinoma of the endometrium.” Int J Radiat Oncol Biol Phys 48(3):779–790. Nag, S., R. R. Kuske, F. A. Vicini, D. W. Arthur, and R. D. Zwicker. (2001a). “Brachytherapy in the treatment of breast cancer.” Oncology (Huntingt) 15(2):195–205. Nag, S., D. Shasha, N. Janjan, I. Petersen, and M. Zaider, (2001b). “The American Brachytherapy Society recommendations for brachytherapy of soft tissue sarcomas.” Int J Radiat Oncol Biol Phys 49(4):1033–1043. Nag, S., E. R. Cano, D. J. Demanes, A. A. Puthawala, and B. Vikram. (2001c). “The American Brachytherapy Society recommendations for high-dose-rate brachytherapy for head-and-neck carcinoma.” Int J Radiat Oncol Biol Phys 50(5):1190–1198. Nag, S., C. Chao, B. Erickson, J. Fowler, N. Gupta, A. Martinez, and B. Thomadsen. (2002). “The American Brachytherapy Society recommendations for low-dose-rate brachytherapy for carcinoma of the cervix.” Int J Radiat Oncol Biol Phys 52(1):33–48. Nag, S., J. M. Quivey, J. D. Earle, D. Followill, J. Fontanesi, and P. T. Finger. (2003a). “The American Brachytherapy Society recommendations for brachytherapy of uveal melanomas.” Int J Radiat Oncol Biol Phys 56(2):544–555. Nag, S., R. Dobelbower, G. Glasgow, G. Gustafson, N. Syed, B. Thomadsen, and J. F. Williamson. (2003b). “Intersociety standards for the performance of brachytherapy: A joint report from ABS, ACMP and ACRO.” Crit Rev Oncol Hematol 48(1):1–17. Paterson, R., and H. M. Parker. (1938). “A dosage system for interstitial radium therapy.” Br J Radiol 11:313–339. Potter, R., E. Van Limbergen, W. Dries, Y. Popowski, V. Coen, C. Fellner, D. Georg, C. Kirisits, P. Levendag, H. Marijnissen, H. Marsiglia, J. J. Mazeron, B. Pokrajac, P. Scalliet, and V. Tamburini. (2001). “Prescribing, recording, and reporting in endovascular brachytherapy. Quality assurance, equipment, personnel and education.” Radiother Oncol 59(3):339–360. Rogus, R. D., M. J. Smith, and H. D. Kubo. (1998). “An equation to QA check the total treatment time for singlecatheter HDR brachytherapy.” Int J Radiat Oncol Biol Phys 40(1):245–248. Saw, C. B., and N. Suntharalingam. (1991). “Quantitative assessment of interstitial implants.” Int J Radiat Oncol Biol Phys 20(1):135–139.
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Saw, C .B., L. J. Korb, B. Darnell, K. V. Krishna, and D. Ulewicz. (1998). “Independent technique of verifying highdose rate (HDR) brachytherapy treatment plans.” Int J Radiat Oncol Biol Phys 40(3):747–750. Stitt, J. A., J. F. Fowler, B. R. Thomadsen, D. A. Buchler, B. P. Paliwal, and T. J. Kinsella. (1992). “High dose rate intracavitary brachytherapy for carcinoma of the cervix: The Madison system: I. Clinical and radiobiological considerations.” Int J Radiat Oncol Biol Phys. 24(2):335–348. Stock, R. G., N. N. Stone, M. Dahlal, and Y. C. Lo. (2002). “What is the optimal dose for 125I prostate implants? A dose-response analysis of biochemical control, posttreatment prostate biopsies, and long-term urinary symptoms.” Brachytherapy 1(2):83–89. Thomadsen, B. R., S. Shahabi, J. A. Stitt, D. A. Buchler, J. F. Fowler, B. R. Paliwal, and T. J. Kinsella. (1992). “High dose rate intracavitary brachytherapy for carcinoma of the cervix: The Madison system: II. Procedural and physical considerations.” Int J Radiat Oncol Biol Phys 24(2):349–357. Thomadsen, B. R. (1999). Achieving Quality in Brachytherapy. Bristol: Institute of Physics, 1999. Wang, X. H., and L. Potters. (2001). “A theoretical derivation of the nomograms for permanent prostate brachytherapy.” Med Phys 28(4):683–687. Williamson, J. F. “Clinical Brachytherapy Physics” in Principles and Practice of Radiation Oncology, 4th ed. C. A. Perez, L. W. Brady, E. C. Halperin, and R. K. Schmidt-Ullrich (eds.). Philadelphia: Lippincott Williams & Wilkins, 2003. Yu, Y., L. L. Anderson, Z. Li, D. E. Mellenberg, R. Nath, M. C. Schell, F. M. Waterman, A. Wu, and J. C. Blasko. (1999). “Permanent prostate seed implant brachytherapy: Report of the American Association of Physicists in Medicine Task Group No. 64.” Med Phys 26(10):2054–2076. Also available as AAPM Report No. 68.
Chapter 24
Quality Management for Interstitial Implants Anil Kumar Sharma, Ph.D. Radiation Oncology Department Long Beach Memorial Medical Center Long Beach, California Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 459 Permanent Interstitial Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 461 Temporary Interstitial Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 461 Quality Management for the Implantation Procedure . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 461 Template-based Implants Using Rigid Guides or Needles . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 461 Prostate Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 462 Gynecological Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 463 Template-based Breast Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 463 Multiple Site Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 464 Flexible Catheter Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 464 Breast/Chest Wall Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 464 Head and Neck Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 465 Other Sites . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 466 Practical Considerations of Implant Reconstruction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 466 Templates and Rigid Needles . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 466 HDR Prostate Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 467 Gynecological Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 467 Flexible Catheters . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 468 Breast Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 468 Head and Neck Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 468 Quality Management of Treatment Planning and Evaluation . . . . . . . . . . . . . . . . . . . . . . . . . . 468 Prostate Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 469 Gynecological Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 470 Head and Neck Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 471 Quality Assurance of Source Application and Removal . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 471 Interstitial Brachytherapy Treatments with LDR . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 471 Interstitial Brachytherapy Treatments with HDR . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 472 Needle or Catheter Movements and Possible Corrections . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 473 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 473
Introduction The aim of quality management in brachytherapy is to ensure that each individual radioactive source application to the area of interest is administered accurately according to the radiation oncologist’s intent, consistent with the written directives and within the scope of the institution’s radioactive material license. Also, the procedure followed for the treatment should be safe for the patient, staff, and others who may be exposed to radiation during the course of treatment. In order to provide satisfactory brachytherapy services to the patient with full compliance of the federal and state regulations, an institution is required to develop a quality management program (QMP), which becomes an essential part of their brachytherapy program. The federal and state regulations are meant for the safety of the patient, personnel, and general public; and without making any clinical judgments relative to the patient’s treatment, they help define the standard of practice for all those involved in delivering brachytherapy services. They form the basis of a necessary, but not sufficient, quality assurance (QA)
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program in brachytherapy. The regulatory agencies, like Nuclear Regulatory Commission (NRC) or the agreement states’ Radiological Health Branch, require that the institution’s practices assure that the radiation from byproduct material will be administered as directed by the authorized user. So, like all other medical treatments, the prescription for interstitial brachytherapy (written directive) is an order in writing for a specific patient, dated and signed by the authorized user prior to the administration of radiation by brachytherapy. For low dose rate (LDR) applications, this prescription must contain the treatment site, name of the radioisotope, number of sources and source strength, and prior to completion of treatment, time of implant or total dose. For high dose rate (HDR) brachytherapy, prior to the treatment, treatment site, isotope, and total dose must be mentioned in the written directive (USNRC 1991a). The objectives of federally enforced QMP (the term “QMP” is dropped in the current part 35, but the requirements are retained) are all related to the written directive as follows: • Written directive before administration of radiation • Identification of the patient before administration of radiation (by more than one way) as the individual named in the written directive • Final plan and related calculations in accordance with the written directive • Administration of radiation in accordance with the written directive • Identification and evaluation of any unintended deviation from the written directive and implementation of appropriate action. NRC requires annual review of the actual treatment with respect to the written directive for a number of patients based upon statistical sampling principles. Also, NRC regulatory guide 8.33, Task DG-8001, contains requirements and recommendations of a QA nature to implement the QMP (USNRC 1991b). The recommendations in this guide can be adapted by individual licensees to arrive at a QMP suitable for their practice and needs, the net effect of which is to emphasize QA in brachytherapy. Besides NRC’s and the agreement states’ limited requirements and recommendations, detailed QA recommendations have come from several organizations. The brachytherapy section of the AAPM’s Task Group 40 (TG 40) (Kutcher et al. 1994) report, discusses both LDR and HDR brachytherapy QA in detail. For HDR brachytherapy practice, AAPM’s TG 59 (Kubo et al. 1998) is a comprehensive guide with several recommendations. Interstitial Collaborative Working Group (ICWG) of the American Brachytherapy Society (ABS) in their report (ICWG 1990) recommended source calibration procedures and other QA requirements in brachytherapy. From time to time ABS has come up with recommendations for several brachytherapy procedures (Nag et al. 2000, 2001, 2002). A major focus of the QA program is to assure accurate operation of all software and mechanical and radioactive devices used for planning, delivery, or QA of brachytherapy treatments. Each institution practicing interstitial brachytherapy should develop a written, periodic QA protocol defining the tests to be performed and their frequency (Nath et al. 1997). Most references cited in the previous paragraph deal with the QA program related to radiation safety in brachytherapy, source calibration, safe handling and application of radioisotopes either manually or through remote-controlled devices, QA of these units and source calibration devices, and QA of treatment planning systems. The goal of periodic device QA is to ensure that the operating characteristics of the device remain unchanged with time. The other equally important part of a brachytherapy QMP is the procedure-specific quality assurance, which is defined as a set of actions selected to ensure that each important step leading to delivery of a brachytherapy procedure is correctly carried out. In this chapter, quality management of interstitial brachytherapy will be discussed with special emphasis on this procedure-specific QA for manual and remote afterloaded interstitial implants. Interstitial brachytherapy, which is the direct application of radioactive isotopes into the tumor, can broadly be divided into two categories: permanent and interstitial brachytherapy.
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Permanent Interstitial Brachytherapy In permanent interstitial brachytherapy the radioactive sources are directly implanted into the tumor and left there permanently to decay. Normally, such applications are carried out in the operating room after careful preplanning. Permanent seed implants of the prostate, pancreas, nasopharynx, unresectable and unsuitable tumors of the head and neck, and stereotactic brain implants using 125I or 103Pd fall into this category. The radioactive sources are transported to the operating room (OR) and flashed and implanted by the radiation oncologist according to the extent of the tumor and safety of the surrounding structures, following the preplan carried out prior to the actual procedure. The number of sources used for such implants depends on the type of the radioactive isotope, size of the tumor, total dose to be delivered, and the initial activity of the source. Almost invariably all sources in permanent implants have the same air kerma strength. Quality management of such implants is dealt with elsewhere and will not be discussed in this chapter.
Temporary Interstitial Brachytherapy In temporary interstitial brachytherapy the radioactive sources are either directly implanted into the tumor or are afterloaded into a set of catheters, which may be rigid needles (metallic, ceramic or plastic) or flexible tubes (nylon or plastic) arranged in an appropriate geometrical pattern with an intention to cover the whole tumor while sparing surrounding normal tissue. In temporary implants, after an adequate amount of dose is delivered to the tumor, the sources are removed from the tissue or from the catheters into which they were sent. The total time of the implantation depends on the type of source, number and strength of each source, and the geometric pattern in which they are arranged. Most interstitial implants in the early days of brachytherapy were direct implantation of radium needles, which subsequently got replaced by 137Cs needles in late 1970s. Today, most interstitial implants are catheter based (rigid needles or flexible tubes) into which small radioactive sources (mostly 192Ir) are afterloaded either manually or through a remote afterloading machine. Quality management of such afterloaded implants is the main topic of this chapter.
Quality Management for the Implantation Procedure Interstitial brachytherapy techniques vary according to the site to be implanted, tumor volume, and whether the target is approachable from multiple sides or from just one side. Usually, the sites that can be approached from only one side are implanted with rigid guides using templates, whereas others can be implanted with flexible catheters with or without the aid of a template. Templates not only help in guiding the catheters into the target at the time of implantation, but also help in keeping the implant in position during the course of treatment. Given the wide variety of interstitial procedures performed in any institution, it is understandable that the philosophy of maintaining quality of the implantation is different for different types of applications. Selection of the type of applicators, catheters, or needles for placement in the patient’s tumor is under the control of the radiation oncologist in consultation with the medical physicist. Physics QA duties include documentation of the catheters or needles inserted and their correct correlation with target volume.
Template-based Implants Using Rigid Guides or Needles Template-based implants with rigid guides are very common for the treatment of prostate, rectum, breast, and gynecological tumors. Templates for such applications come in various shapes and sizes and can be re-useable or disposable after one application. It is very important to decide, before the team heads for the OR, which template is most suitable for the patient and what kind of guides are required for the proce-
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dure. For example, if the treatment is planned to be given by LDR 192Ir ribbons, the set of needles required is different from those used for HDR brachytherapy. Also, if interstitial radiofrequency hyperthermia is part of the treatment strategy, steel needles need to be implanted and used as antennae for hyperthermia. In such situations, the number of implanted needles will be more than required just for brachytherapy alone because for better uniformity of hyperthermia, the desired pattern of needles may be different. Therefore, a checklist of items required for the implantation procedure is the first step of a QMP. This list may vary with the site to be implanted. Some examples follow. Prostate Implants Syed-Neblett prostate template and the Martinez Universal Perineal Interstitial Template (MUPIT) are most commonly used for the treatment of carcinoma of the prostate. Whereas the MUPIT is essentially a grid of allowable guides in a square pattern with holes at every 0.5 cm interval, the Syed-Neblett prostate template has a provision for introducing a maximum of 18 needles arranged in two concentric circles. Occasionally, in order to encompass a very large prostate, templates designed for rectal applications can be used. The rectal template (also called the large prostate template) has an additional outer concentric circle of 3-cm radius around the inner two circles (of radii 1 and 2 cm) of a regular prostate template. These templates are designed to hold 17-gauge needles perpendicular to the template, but in some cases the needles can either be flared out or coned in to fit onto the target (Syed et al. 1992, 1997). The number of needles used is usually fewer than 18 in most prostate cases, most often from 14 to 18. Pre-procedure volume studies are carried out to determine the gland size, any extracapsular extension, calcification, median lobe involvement, seminal vesicle involvement, transurethral resection of the prostate (TURP) etc., so that the appropriate size template and strategy for treatment can be chosen. A three-way Foley balloon, with 50% Hypaque™ (contrast) solution in the balloon is required for contrast injection directly into the bladder before planning computed tomography (CT) is done. Before implantation of needles under transrectal ultrasound (TRUS) guidance, two gold markers are inserted into the base and two into the apex of the prostate. These markers can be used in applicator-based planning. Quality management for prostate implantation would include the following practical considerations: • Determination of the type of implant (LDR or HDR), so that appropriate needles can be chosen for implantation. LDR needles cannot be connected to the afterloader because only HDR needles have special features for connection to the afterloader. • Verification of the template orientation at the time of implantation. If the template is unintentionally rotated by 90°, the configuration of inner circle (urethral circle) becomes different from the usual convention, resulting in confusion at several stages, from numbering the channels to loading with a set of dummy trains to treatment planning and finally to the actual treatment, because the written procedure for each of these stages is meant for conventional position of the template. The template is in its correct orientation when two holes in the first circle coincide with the 12 o’clock and 6 o’clock positions. • Verification of the proper functioning of the ultrasound unit, and its calibration with respect to the template (if it is a transducer fixed system). • Post procedure notes in the patient’s chart by the physician, the diagram of the implant with written directives clearly spelling out the total dose, dose rate for LDR, or dose per fraction for HDR. • Sending the completed implant loading form to the brachytherapy dosimetry team in radiation oncology. Recently, two commercial systems, SWIFT™ from Nucletron Corporation (Columbia, MD) and Vitesse™ from Varian Medical Systems (Palo Alto, CA) have become available for ultrasound image-
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based implantation and dosimetry. For such systems extensive QA of the ultrasound imaging system and calibration of the template for one-to-one correspondence with the template in the ultrasound image is required on the same lines as for the permanent prostate seed implants. For details, QA of permanent seed implant devices may be referred to elsewhere in this monograph. Gynecological Implants The most used template systems for interstitial implantation, Syed-Neblett GYN and MUPIT, were introduced in 1974. Both systems use 17-gauge needles, but differ in their pattern of allowable guides. MUPIT allows needles in its type I holes at 1-cm intervals and perpendicular to the template, and in type II holes at 1.25-cm intervals at a 13° laterally outward direction (Martinez, Cox, and Edmundson 1984). The vaginal and rectal cylinders, available in two sizes, compose the intracavitary component of the system, and can accommodate an intrauterine tandem or drainage tube. Each of the cylinders can also accommodate up to eight needles. MUPIT can be used for both LDR and HDR applications. Commercialized templates are available through Nucletron Corporation for HDR applications. The Syed-Neblett template (Syed et al. 1986; Neblett 1995) for gynecological malignancies has a pattern of holes (1 cm apart) in five concentric circles or arcs at 1-cm radial increments and a central grooved vaginal obturator of 2 cm diameter and 15 cm length, which can accommodate six needles and a centrally located interauterine tandem of medium curvature. These templates are available as reuseable or disposable from Alpha Omega Services, Bellflower, CA, and Best Industries, Springfield, VA. All LDR needles used with these templates except the first needle (also called “guide”) have small, raised metal rings near the proximal end to prevent them from sliding through the template. HDR needles also are shaped at their proximal end for connection to an afterloading system. The implantation procedure is discussed in detail elsewhere in this monograph. The following practical considerations form part of the quality management program: • Choice of template (MUPIT or Syed-Neblett GYN 1, 2 or 3) needs to be made after careful evaluation of the pre-procedure imaging study. CT and particularly magnetic resonance (MR) images are extremely helpful in defining the volume to be implanted, and providing knowledge about the central and peripheral disease. If the disease is extensive in the anterior-posterior (AP) direction, choice of GYN 1 or 2 templates will lead to undercoverage of the target. • Type of implant (LDR or HDR) must be determined, so that appropriate needles can be chosen. Also, from the imaging studies, the depth of needles can be determined and whether the vaginal length requires the long obturator and longer needles (25 cm). • Implantation of gold markers into the cervical os requires an accuracy of better than 2 mm, since these markers are used for applicator-based planning. • Even though insertion of all the allowable needles is not necessary and would depend on the extension of disease, enough needles (usually more than 10 in either side of the obturator) need to be implanted to provide stability and rigidity to the template. • In LDR implants, obturator needle insertion is discouraged in the presence of an intrauterine tandem; but for HDR implants where dwell-time optimization is possible, implantation of four lateral obturator needles is found to be helpful in adequate coverage of centrally bulky disease. These needles are now regularly implanted in HDR applications, but their loading is decided at the time of planning. Template-based Breast Implants Conventionally, freehand breast interstitial implants have been performed as boost using either 192Ir ribbons or 125I seeds for LDR implants or for afterloading HDR fractionated treatments using 192Ir. Recently, partial
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breast irradiation by interstitial implants has shown its potential for selected patients with early stage disease (Vicini et al. 1999; Baglan et al. 2001). Implants are carried out using a standardized template, with interplane and intercatheter separation of about 1.5 cm. Intraoperative ultrasound-guided placement of afterloading catheters can also be used to ensure that the catheters encompass the biopsy cavity. Detailed description of breast brachytherapy is covered elsewhere in this monograph. Quality management for the template-based breast implants will include proper choice of the template and ensuring that it is secure at the time of implantation so that the needles are arranged with proper separation as per the pre-plan. If ultrasound guidance is used, proper QA of the ultrasound unit is required. Multiple Site Implants Multiple site implants in close proximity are usually mixed implants, i.e., a template-based implant and a freehand flexible catheter implant or an intracavitary application. Examples of the former are vulvar implants and breast and axillary lymph node implants, whereas interstitial-intracavitary cervical and some extensive maxillary implants fall in the latter category. Additional considerations about the possible overlap of two implants need to be taken into account. Quality management of flexible implants is discussed in the next section and its extent in the presence of a rigid implant requires clear demarcation of the boundaries of each implant, at least externally, before the start of the implantation process. If there is any intracavitary application within or close to the rigid needle implant, securing the intracavitary applicator in place within 2 mm is required.
Flexible Catheter Implants Several sites are easily accessible from at least two sides for freehand flexible catheter implants, which are normally performed in the OR under general anesthesia, except for breast implants, where local anesthesia may be used instead. General considerations for such implants are to ensure the patency of the implanted catheters, availability of sterile (angled) sharp needles (trocars), buttons (plastic buttons of different colors, metallic buttons to be used as stoppers, and gold buttons for shielding), appropriate catheters, sterile rulers, and marking pens. Breast/Chest Wall Implants Either a single-channel device, called MammoSite®, or multiple catheters implanted interstitially are used for breast implants. Quality assurance features of these implants are very different from each other. Single-Channel Breast Implants (MammoSite®) MammoSite applicator (Proxima Therapeutics Inc., Alpharetta, GA) is a single-channel applicator with a silicon inflatable balloon at its distal end and two ports; one for HDR source application and the other for injecting fluid into the balloon, on its proximal end. This device can either be implanted at the time of lumpectomy or later within a few weeks of lumpectomy. This subject is dealt with in depth elsewhere in this monograph. Before implantation, ensuring that the there is no leakage in the balloon is an important QA step. If the cavity size is not known, at least two sizes of the spherical/elliptical applicators should be taken to the OR. A small diameter catheter, such as a baby NG tube, may be implanted side by side with the MammoSite applicator, which may be used later to get rid of the air pockets. If implanted a few weeks after lumpectomy, ultrasound guidance is required in identifying the cavity. The ultrasound unit should be calibrated before its use. Ensuring that the applicator is properly secured in its place and the HDR port has the flexible stylet in it are part of the QA for this procedure.
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Freehand Planar Breast Implants After marking the intended entry and exit points of these needles on the breast surface, 17-gauge stainless steel trocars are inserted. Most often two or three planar implants are carried out, interdigitating needles in the neighboring planes. • Interplanar distances and intercatheter distances should follow Paris or Manchester system rules for LDR implants for 192Ir wires and ribbons, respectively. For HDR implants, this spacing, though desirable, is not very crucial. • Catheters should be secured by the buttons at both ends, so that the unintended movement of catheters is no more than 2 mm. • Catheters of each plane should be of different colors or at least a different color button should be used for securing them. • Flexible metallic or plastic stylets should be inserted after the implant is done to ensure patency of the catheters. • Single-plane chest wall implants require careful spacing to avoid hot and cold spots in the implant. Head and Neck Implants Using flexible afterloading plastic catheters, previously unimplantable lesions involving retromolar trigone, soft palate, tonsellar pillars, phrangeal walls, and base of tongue can be effectively implanted (Puthawala et al. 1986; Syed and Puthawala 1996). Planar Implants Mostly used for neck nodes, these planar implants have the same kind of requirements as mentioned above for the freehand breast implants. Gold Button/Button Technique This technique is used in most of the oral cavity and oropharyngeal cancers. In this technique, the plastic tubes are held between a gold and plastic button at its sealed end and a plastic button on the skin. The gold button protects the overlying tissues against irradiation and trauma. Loop Technique/Arch Technique This technique is used for cancers encroaching on the mandible and occasionally for base of tongue lesions. The technique involves insertion of one plastic tube through the intraoral ends of a pair of hollow needles implanted on either side of the mandible. When the needles are removed, the plastic tube is wrapped over the alveolar ridge. This technique is also used for lesions involving the soft and hard palate and tonsillar pillars. In this technique, one plastic tube is inserted through the intraoral ends of a pair of hollow needles implanted from either side of the neck or cheek through the palate. In most oral cavity and oropharyngeal cancers, a combination of various techniques is used to encompass the tumor adequately. Multiple Site Implants It is a very common practice in head and neck brachytherapy that multiple sites are implanted at the same time. Various combinations of techniques and sites could be very confusing for the implanter as well as for the planner. Documentation of each site with the number of catheters and the technique used for
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implantation should be made. Whereever the catheters from one site overlap into the other, marking on the skin should be made as a cautionary measure. QA considerations for head and neck implants are as follows: • Before implantation, marks on skin for proper catheter separation • Different color buttons for different planes of an implant. Different color catheters for different implants, in the case of multiple-site implants • Each catheter secured in its place with a metallic button so that the catheter movement is secured within 2 mm. Distal end blocked to ensure source ribbon does not go through • The intraoral plastic button(s) which serve as a spacer to allow the 192Ir sources to project beyond the mucosa • Not using the loop applications as for LDR implants for HDR applications since the HDR source may not be able to take sharp turns in the loop • For LDR implants, measurement of the catheter length and determination of number of seeds for each catheter • Clear diagram of the color-coded implant for brachytherapy dosimetry team. Other Sites Many other sites can be implanted with flexible catheters, abdomino-pelvic implants using a HAM applicator, or freehand implantation of tumor bed with multiple catheters in a single plane for recurrent tumors of rectum, bladder, sarcomas, etc.
Practical Considerations for Implant Reconstruction Rigid needles or flexible catheters implanted in the tumor and the critical structures located in the vicinity of the implant need to be reconstructed for the purpose of treatment planning in both LDR and HDR applications. AAPM TG 40 recommends that the position of all applicators or catheters be verified either radiographically or by CT. The ABS recommends that target volume delineations include (in the order of preference) determination on CT or MR images, through surgical clips or by projection drawn on orthogonal radiographs. Template-based implants and freehand implanted catheters are considered separately.
Templates and Rigid Needles To calculate a dose distribution for an interstitial implant, coordinates of all the radioactive sources must be determined relative to an arbitrary reference point that is used as the origin. This can be accomplished by orthogonal radiographs, variable angle radiographs, stereo-shift films, 3-film technique, or more elegantly by CT or MR scanning. For identification of individual catheters, coded dummy trains having a unique pattern are loaded into the catheters. The following practical considerations can be used as guide for the quality management of implant reconstruction of template-based interstitial implants: • Each template should be named and its allowable set of holes numbered in a unique fashion. For example, numbering of Syed-Neblett templates, at our institution, starts from the top (1 o’clock position) needle of the innermost circle and increases clockwise, from innermost to the outer concentric circles or arcs.
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• After the patient is implanted, all the implanted channels are numbered on a printed diagram of the template following the above convention. • For film-based reconstruction, each channel is loaded with a uniquely defined dummy train and the loaded dummy’s identification is marked against the channel in which it is loaded. • For CT-based reconstruction, there is no need to load dummy trains in steel needles, but plastic catheters may be loaded with dummy markers to be able to see them in scout views. • All dummy markers should truly represent the source wire, with the cold tip or leader of the same size as that of the source train. • The patient should be positioned in a reproducible position (we use three-point set up) so that when radiographs or scout films are taken before each fraction, the patient is in the same position as for the planning radiographs or CT. It is extremely important to ensure patient immobilization during the entire procedure of localization films, because identification of various channels and calculation points may require data entry from different sets of films. HDR Prostate Implants • A set of radiographs may be taken in the OR to ensure that the implant is in the correct place and the needles have not moved after the patient is brought to supine position from the earlier modified lithotomy position. This is important for implants with the needle tips exactly below the bladder base. • For film-based planning, a set of orthogonal or stereo-shift films with dummy trains is required. It has been reported (Sharma, Williamson, and Cytacki 1982) that accuracy of localization is better for orthogonal views; a set of AP/Lat films is desirable. • Three main critical structures in the prostate cases are bladder, urethra, and rectum. The Foley balloon, introduced previously into the bladder, should be filled with diluted (50%) contrast solution, otherwise the balloon may obscure dummy markers in the needles. To view the implant position with respect to a full bladder position, diluted (20%) Hypaque solution is injected directly into the bladder (that’s why three-way Foley) allowing easier urethral marking. • For CT-based planning, 2-mm transverse slices are desirable to identify the distal ends of the needle tips accurately. Gynecological Implants • Gynecological implants usually are characterized with numerous channels that can extend laterally about 7 cm on either side of the central obturator. Lateral radiographs result in severe overlapping of the channels. Several sets of orthogonal radiographs with selective loading of channels with dummy markers will be required. However, if two oblique (45° from the anterior) are taken with all the channels loaded with dummy ribbons, it is sometimes possible to work with just one set of radiographs. • Identification of the bladder points can be done according to ICRU 38 (ICRU 1985), but the rectal points cannot be defined in the same way as there is no vaginal packing. However, using a rectal marker or contrast, rectal points can be assigned. CT-based planning is preferable, since it allows delineation of structures.
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Flexible Catheters Even though most freehand catheters are implanted without the aid of any well-defined geometrical pattern, their numbering should follow a very consistent order to avoid any kind of confusion for treatment planning and source loading. It is imperative that a preplan is carried out in such cases and guidelines from classical systems like Manchester, Paris, or Quimby be observed. Once the implant is complete, a diagram closely resembling the implant may be drawn showing different catheters in at least two planes, e.g., transverse and coronal or sagittal planes. Breast Implants Localization of MammoSite is carried out by inserting a dummy wire through the HDR port right up to the tip of the applicator. Either orthogonal films can be taken or CT scan performed after ensuring that the balloon is inflated with the planned amount of fluid (20% contrast). Without CT, however, the dose to the skin cannot be assessed accurately, nor can the trapped air be visualized, two problems that would lead to aborting the procedure. On the other hand, radiographic verification of the size of the balloon before treatment (to assure that it has not leaked) should be a routine QA procedure. Numbering of catheters is crucial in the case of multichannel implants. It is important to follow a convention consistently to avoid loading the catheters incorrectly. We follow the convention of shallow to deep planes with numbering starting at the most superior catheter of the shallow plane. All catheters are marked and the diagram of these catheters is made for planning and dose delivery. Head and Neck Implants The following points may be useful in proper reconstruction of complex head and neck implants: • Mark the catheters from superior to inferior starting at the shallow plane. • If the implantation is carried out in sagittal planes, start from the left side plane, numbering the catheters from anterior to posterior. • If there are multiple implants in close proximity, catheters should be marked continuously without separating them. • Numbering should be done for the corresponding flexible stylets as well, because catheters and their corresponding stylets will generally be of different sizes. • Patients usually are in pain (or sometimes even disoriented); carefully positioning them on the table with a headrest and possibly a masking tape to stabilize the head may be useful. Immobilization is essential because several films may be required and the patient’s position reproducibility is crucial. • Dummy ribbons should be loaded and their position in the catheters noted on the diagram. • Orthogonal films or variable angle oblique films are preferable. • For CT-based planning, the dummy ribbons are generally not required, but in some cases artifacts due to dental fillings or the gold buttons (used for providing shielding) may obscure the catheter’s distal end, necessitating the need to use dummy ribbons.
Quality Management of Treatment Planning and Evaluation Treatment planning for interstitial implants should be carried out on planning systems, which use ICWG formulism adopted in AAPM’s TG 43 report (Nath et al. 1995). Source strength should be specified as air kerma strength of the source to be used, with proper corrections for the source anisotropy along its
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axis. The planning system should be thoroughly tested before its clinical implementation. Also, periodic performance re-evaluation should be carried out by using reproducible catheter placement in a simple geometrical pattern. Post implant dose calculations should be performed after localization of the catheters or needles with respect to the target volume and the organs at risk. Dose distribution should be obtained in multiple planes after volume dose calculations. All patient dose calculations should be verified by the brachytherapy physicist by reviewing that no gross errors have occurred and by performing an independent dose calculation to one critical point. Dosimetry should be performed by a dosimetrist or a medical physicist and should be reviewed with the attending radiation oncologist before the start of an HDR fraction or during the first half of an LDR treatment.
Prostate Implants • After proper reconstruction of the implant, loading of channels is carried out as per the delineated target volume or according to the gold seeds previously implanted into the prostate. • Activity per seed in the central six needles is half of that used for the outer needles for LDR implants, dose rate about 65 cGy per hour. • For HDR implants, catheter length parameter, i.e., the distance from the indexer to the tip of the needles should be measured for all channels and entered into the planning system. This should be with 1mm for all channels. • For HDR implants, geometrical optimization works very well, but for further urethral dose reduction, the central six needles’ dwell time should be decreased to about one-third of the original dwell time. • Dose to the bladder points and rectal points should be calculated and be less than the prescribed dose; urethral dose also should be controlled. At our institution, no point in the urethra gets more than 110% of prescribed dose. • For TURP and re-irradiation cases, the dose to urethra should further be reduced by decreasing the dwell time in the central six needles (producing a hollow cylinder or doughnut shaped distribution). • Dwell positions for HDR implants should remain within the target volume. • The dose prescription and dose limitation to organs at risk will however take precedence over the above suggestions. • Dose evaluation for LDR or HDR prostate implants should be carried out and reported as per the ABS and ICRU 58 (ICRU 1997) guidelines. • Geometrical optimization generally results in doses in the center of the implant that are higher than the peripheral (or reference) dose. However, homogenous dose distributions can be achieved by manually reducing the dwell time for the inner circle catheters or differentially loading the catheters for LDR implants (Thomadsen et al. 1990). • Reading point dose values, particularly in the center of the implant, and drawing dose profiles at various levels help in assessing the dose homogeneity in the implanted volume. • Dose distributions are visually evaluated by inspecting the isodose lines in multiple planes throughout the planned target volume. The foremost optimization goal for prostate implants is to have the target volume covered as closely as possible with the reference isodose surface. The smallest-volume isodose surface covering the entire target volume is taken to be the planned target dose (PTD).
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Anil Kumar Sharma • Though high-dose gradients exist around the source dwell positions, the total volume of the hyper dose sleeve (200% PTD) should be restricted to less than 10% of the treatment volume • Further analysis of the dose distribution should be carried out using differential dose volume histograms (DVHs) of the optimized plans. If the prostate volume implant is optimized to the same dose midway between the catheters, then the differential DVH will show a peak for that dose. • This peak is somewhat distorted by an inverse square law effect. These obscurative effects are eliminated in the “natural” DVH (Anderson 1986), a differential histogram for which the derivative of the baseline variable with respect to dose is proportional to the –5/2 power of the dose. Using the natural DVH, evaluation of the dose distribution between the catheters and in the treated volume is evaluated and the effect of dose optimization is assessed. • From the point dose values and the shape of dose profiles and DVHs, the extent to which the dwell time values for the central catheters are to be lowered is inferred. Changes can be made in a graphical optimization mode by simply dragging the graphical bar view of dwell times with the mouse or by typing in the required values. • Since attempts are made to make the dose distribution in the target volume homogeneous, it should approach Manchester system type loading. The mg-hrs required from Paterson-Parker tables (Paterson and Parker 1938) for the reference isodose surface volume can be compared with the actual mg-hrs obtained after appropriate conversions from the planned total air kerma strength for second calculation. In our center for prostate implants, these values matched within 5% for more than 97% of the cases, in fact for 68% of the treatment plans the mg-hrs were within 2% (Sharma et al. 1998).
Gynecological Implants Planning of gynecological tumors also follows most of the considerations mentioned above for the prostate implants, because the channels tend to be straight and implanted almost equidistantly and, loaded with almost the same spacing, form a geometrical matrix which is most suitable for geometrical optimization. However, several features are different from the prostate implants: • Central dose should be higher, because most tumors are at or around the center. For LDR implants 125% higher dose in the center is acceptable for non-tandem cases. With tandem a in place, the central dose is much higher. • For LDR planning, the central obturator needles are not supposed to be loaded if tandem is present, which alone is loaded with 10-5-5 mgRaeq type of loading. However, for HDR planning four lateral obturator needles can also be lightly loaded for better uniformity of the dose. • Obturator needles should not be loaded inferior to the external os, whether or not a tandem is in place, unless clinically indicated for extended vaginal disease. • Usually, several catheters are implanted (up to 44, and almost all are loaded for LDR implants with 192 Ir seeds—inner channels either loaded with lesser activity per seed or differentially unloaded in time) and tandem (if present) with 137Cs tubes. • For HDR planning, most planning systems allow more channels for planning than the maximum number of channels treatable on the afterloader. If the treatment planning or treatment unit only allows a maximum number of catheters equal to the number of channels, a composite plan should be carried out and later divided for the sake of treatment on the HDR machine. Extreme vigilance is required in such cases, where two subplans are used for treatment as far as the channel numbers on the patient and their corresponding planning catheter numbers are concerned.
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• Treatment planning and reporting for gynecological tumors should be volume based, as per ICRU 38 (ICRU 1985) and ICRU 58 reports (ICRU 1997). • Modified Paterson-Parker tables can again be used for a second check of dose calculations. For plans without tandem, the results are usually within 7%, and for plans with tandem within 15%.
Head and Neck Implants • Treatment planning of flexible catheters is in general more difficult as there can be more than one simultaneous implant and each one requires a different type of planning. For example, for tongue implants, the most distal source needs to be loaded above the mucosa for LDR implants, and with much higher dwell time for HDR implants. Whereas for neck implants dwell times near the skin should be reduced and dwell positions kept at least 5 mm from the skin. Tonsillar fossa catheters are loaded only to or slightly beyond the mid line, and the dwell times in the distal positions will depend on the position of other catheters that are implanted in planes orthogonal to the tonsillar fossa catheters. • A composite plan for multiple implants always should be made for dosimetry and evaluation. For HDR treatments, if required, different plans can be exported separately to the treatment machine. • For HDR treatments, all catheter length parameters should be measured individually, at both the entrance to the skin and the exit (or most distal) ends. The length of the catheter within the tissue as obtained by radiographic or CT means should be compared with these measurements. Both should agree within 2 mm. • Dose prescription in these implants is volume based; the target volume is either delineated on the CT slices or taken from the surgical markers. Effective interaction between physicist and radiation oncologist forms the basis of an effective treatment plan. • Plan evaluation should be carried out in all the transverse and sagittal planes to check for any hot or cold spots and any overlapping of loaded catheters. Natural DVHs are of great help in picking up the right isodose surface and for evaluating the plan using its dose uniformity and quality indices.
Quality Assurance of Source Application and Removal The most common radionuclide for LDR temporary interstitial implants is 192Ir, in the form of seeds arranged at specific intervals (usually 1 cm) in ribbons or wires of desired lengths. Source calibration, leak testing, receiving, logging, safe handling, and returning to the suppliers, etc., are regulated by NRC and agreement states. The institutions are required to manage their radiation sources in storage and in the patients as per their Radioactive Material Licenses (RAMLs). An excellent review on this topic is can be found in Williamson, Thomadsen, and Nath (1995), Thomadsen (1995), Thomadsen et al. (1994), and in chapter 9 on regulations in this monograph. In this section, only the practical considerations pertaining to the LDR and HDR interstitial brachytherapy treatment are presented.
Interstitial Brachytherapy Treatments with LDR • Based on the treatment plan, 192Ir source ribbons of proper strength and size are ordered and received in the department. • The number of sources and their strengths are compared with the user’s order and logged.
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Anil Kumar Sharma • All sources should be stored in a secure room, with a radioactive material sign posted on its door. When opening the package, it should be determined that there is no contamination due to damage during shipping. Exposure at 1 m and at surface, along with a wipe test of the outer container, should be noted and logged. • Ribbons should be transported to the patient’s room in a cart with proper labeling of radioactive material after calibration as per AAPM TG 56 report (Nath et al. 1997). • Loading should be carried out by the authorized user and physics staff using long-handled forceps. A radiation survey with a calibrated meter should be done after the sources are loaded and secured in the respective channels as per the plan. The time of source loading and survey results are noted down in the patient’s chart. • A “Caution: Radiation Area” sign is posted on the patient’s door with the type of sources and radiation dose at 1 m, at bed side, and at the door noted on the radiation label. • Emergency phone numbers of the radiation safety officer (RSO) and the physician should also be on the label. • After the completion of treatment, all sources should be removed using long forceps and placed in a shielded container. • A radiation survey should be carried out and its results noted down. When satisfied that all sources have been removed and secured in the container, all radiation labels are removed from the door and room. • Sources are removed from the container and stored in the hot lab for later shipment to the supplier.
Interstitial Brachytherapy Treatments with HDR The AAPM Report No. 41 (TG 41) (Glasgow et al. 1993), Report No. 61 (TG 59) (Kubo et al. 1998). and Report No. 46 (TG 40) (Kutcher et al. 1994) guidelines on remote afterloading technology should be followed for daily, monthly, and other periodic guidelines. • All HDR applications should be done after a proper treatment day QA, which includes position verification of proper operation of the dummy cable and positioning of the source wire to within 1 mm, and functioning of all interlocks including the door, indicator lights, radiation detectors, catheter integrity, timer, and source decay correction. • New source calibration and periodic source calibration as per the RAML must be carried out. Differences up to 5% from the manufacturer’s calibration certificate are generally acceptable, but it is good practice to check with the manufacturer if the difference is more 3%. • Authorized user and medical physicist or RSO must be present for all HDR treatments. • The treatment plan should be compared with that from the treatment planning system by checking the dwell times in each channel and total treatment time. For large implants with numerous dwell positions, random dwell position and dwell times may be sufficient. • The implanted needles or catheters should be connected by one person to their respective numbers on the afterloader turret and verified by a second. • There should be a proper audiovisual contact with the patient during the treatment.
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• Emergency kit should be available in the treatment room, and two radiation survey meters (at least one of them being ionization chamber type with high range (500 R) should be available at the console. • Once the treatment is finished, the physicist surveys the patient, again taking readings at the afterloader surface and at a distance (some states require at 20 cm, NRC requirement is at 1 m) from the afterloader surface. • Before any HDR fraction, the patient with template and needles is either scanned or radiographed to ensure that the applicator is in the right position. MammoSite patients should be radiographed (or scouted) to ensure that the balloon is intact. • For the MammoSite applicator, it is mandatory to ensure that the luer-lock is firmly connected. Also, for all MammoSite patients, endobronchials and endoesophageals should be treated in the scanning position, i.e., supine on a treatment table or a gurney. • All source transfer tubes should be as far off the patient as possible to reduce any contact dose to other areas of the patient. • HDR fractionated treatments should be at least 6 hours apart; any shorter gap between the fractions should be justified and accounted for through recalculation in the dose prescription. For head and neck tumors, the ABS recommends at least a 6-hour gap (Nag et al. 2001). • Any treatment plans with more than the number of available channels on an older treatment planning system that requires the division into two separate plans should be checked very carefully by the authorized user, physicist, and therapist to ensure the planned channel is connected to the treatment channel.
Needle or Catheter Movements and Possible Corrections HDR brachytherapy is delivered over a number of fractions, allowing for potential movement of afterloading needles between fractions. Our own experience has shown that about 25% of the patients require adjustment of the catheter length parameter (Damore et al. 2000). Most movements take place for the second fraction after an overnight stay in the hospital. By keeping adequate margins for treatment, i.e., implanting needles deeper than required, but treating the target by leaving required space at the distal end, one can have enough room for applying corrections in the cleft lip and palate. Inferior displacement of interstitial needles between fractions is a potential source of error in the delivery of HDR brachytherapy for prostate, gynecological, and rectal cancers. Localization films or CT scouts are necessary before any treatment and corrections based on these films can be made so that the target volume is given the same dose pattern approved after planning every time.
References Anderson, L. L. (1986). “A ‘natural’ volume-dose histogram for brachytherapy.” Med Phys 13:898–903. Baglan, K. L., A. A. Martinez, R. C. Frazier, V. R. Kini, L. L. Kestin, P. Y. Chen, G. Edmundson, E. Mele, D. Jaffray, and F. A. Vicini. (2001). “The use of high-dose-rate brachytherapy alone after lumpectomy in patients with earlystage breast cancer treated with breast-conserving therapy.” Int J Radiat Oncol Biol Phys 50:1003–1011. Damore, S. J., A. M. N. Syed, A. Puthawala, and A. Sharma. (2000). “Needle displacement during HDR brachytherapy in the treatment of prostate cancer.” Int J Radiat Oncol Biol Phys 46(5):1205–1211. Glasgow, G. P., J. D. Bourland, P. W. Grisby, J. A. Meli, and K. A. Weaver. “Remote Afterloading Technology.” AAPM Report No. 41. New York: American Institute of Physics, 1993.
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International Commission on Radiation Units and Measurement (ICRU) Report No. 38. “Dose Volume Specification for Reporting Intracavitary Therapy in Gynecology. Bethesda, MD: ICRU, 1985. International Commission on Radiation Units and Measurement (ICRU) Report No. 58. “Dose and Volume Specifications for Reporting Interstitial Therapy.” Bethesda, MD: ICRU, 1997. Interstitial Collaborative Working Group (ICWG). Interstitial Brachytherapy, Physical, Biological and Clinical Considerations. New York: Raven Press, 1990. Kubo, H. D., G. P. Glasgow, T. D. Pethel, B. R. Thomadsen, and J. F. Williamson. (1998). “High dose-rate brachytherapy treatment delivery: Report of the AAPM Radiation Therapy Committee Task Group 59.” Med Phys 25(4):375–403. Also available as AAPM Report No. 61. Kutcher, G. J., L. Coia, M. Gillin, W. F. Hanson, S. Leibel, R. J. Morton, J. R. Palta, J. A. Purdy, L. E. Reinstein, G. K. Svensson, M. Weller, and L. Wingfield. (1994). “Comprehensive QA for radiation oncology: Report of AAPM Radiation Therapy Committee Task Group 40.” Med Phys 21:581–618. Also available as AAPM Report No. 46. Martinez, A, R. Cox, and G. Edmundson. (1984). “A multiple-site perineal applicator (MUPIT) for treatment of prostate, anorectal, and gynecological malignancies.” Int J Radiat Oncol Biol Phys 10:297–305. Nag, S., B. Erickson, S. Parikh, N. Gupta, M. Varia, and G. Glasgow. (2000). “The American Brachytherapy Society recommendations for high-dose-rate brachytherapy for carcinoma of the endometrium.” Int J Radiat Oncol Biol Phys 48(3):779–790. Nag, S., E. Cano, D. J. Demanes, A. A. Puthawala, B. Vikram; American Brachytherapy Society. (2001). “The American Brachytherapy Society recommendations high-dose-rate brachytherapy for head and neck carcinoma.” Int J Radiat Oncol Biol Phys 50(5):1190–1198. Nag, S., C. Chao, B. Erickson, J. Fowler, N. Gupta, A. Martinez, B. Thomadsen; American Brachytherapy Society. (2002). “The American Brachytherapy Society recommendations for low-dose-rate brachytherapy for carcinoma of the cervix, Int J Radiat Oncol Biol Phys 52(1):33–48. Nath, R., L. L. Anderson, J. A. Meli, A. J. Olch, J. A. Stitt, and J. F. Williamson. ( 1997). “Code of practice for brachytherapy physics: Report of AAPM Radiation Therapy Committee Task Group No. 56.” Med Phys 24(10):1557–1598. Also available as AAPM Report No. 59. Nath, R., L. L. Anderson, G. Luxton, K. A. Weaver, J. F. Williamson and A. S. Meigooni. (1995). “Dosimetry of interstitial brachytherapy sources: Recommendations of the AAPM Radiation Therapy Committee Task Group No. 43.”Med Phys 22:209–234. Also available as AAPM Report No. 51. Neblett, D. C. (1995). “Clinical Techniques on Applicators Available for interstitial Implantation” in Brachytherapy Physics. J. F. Williamson, B. R. Thomadsen, and R. Nath (eds.) Madison, WI: Medical Physics Publishing, pp. 282–300, 1995. Paterson, R., and H. M. Parker. (1938). “A dosage system for interstitial radium therapy.” Br J Radiol 11:252–266. Puthawala, A. A., A. M. N. Syed, D. L. Eads, D. Neblett, L. Gillin, and T. C. Gates. (1985). “Limited external irradiation and interstitial 192-iridium implant in the treatment of squamous cell carcinoma of the tonsillar region.” Int J Radiat Oncol Biol Phys 11(9):1595–1602. Sharma, S. H., J. F. Williamson, and E. Cytacki. (1982). “Dosimetric analysis of stereo and orthogonal reconstruction of interstitial implants.” Int J Radiat Oncol Biol Phys 8:1803. Sharma, A., A.M.N. Syed, A. Puthawala, and A. A. Farid. (1998). “HDR planning for prostate cancer implants using Varisource remote afterloader.” J Brachytherapy International 14:1–14. Syed, A. M. N., and A. A. Puthawala. Proceedings of the Workshop on HDR and LDR Brachytherapy Techniques for Prostate, Head and Neck and GYN Malignancies. Long Beach, CA: Long Beach Endocuritherapy Foundation, 1996. Syed, A. M. N., A. Puthawala, P. Austin, J. Cherlow, J. Perley, L. Tansey, A. Shanberg, D. Sawyer, R. Baghdassarian, B. Wachs, et al. (1992). “Temporary iridium-192 implant in the management of carcinoma of the prostate.” Cancer 69:2515–2524. Syed, A. M. N., A. Puthawala, N. Barth, A. Sharma, and A. Londrc. (1997). “High dose rate brachytherapy in the treatment of carcinoma of prostate: Preliminary results.“ J Brachytherapy No vol given:1–14. Syed, A. M. N., A. Puthawala, D. Neblett et al. (1986). “Transperineal interstitial/intracavitary ‘Syed-Neblett’ applicator in the treatment of carcinoma of the uterine cervix.” Endocurither/Hypertherm 2:1–13.
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Thomadsen, B. R., P. V. Houdek, R. van der Laarse, G. Edmundson, I. K. Kolkman-Deurloo, and A. G. Visser. “Treatment Planning and Optimization” in High Dose Rate Brachytherapy: A Textbook. S. Nag (ed.). Armonk, NY: Futura Publishing, pp. 79–145, 1994. Thomadsen, B. R., S. Shahabi, D. Buchler, W. Giese, and M. Mehta. (1990). “Differential loadings of brachytherapy templates.” Endocuriether/Hypertherm Oncol 6:197–202. Thomadsen, B. R. “Clinical Implementation of Remote Afterloading Interstitial Brachytherapy” in Brachytherapy Physics. J. F. Williamson, B. R. Thomadsen, and R. Nath (eds.). Madison, WI: Medical Physics Publishing, pp. 679–698, 1995. U.S. Nuclear Regulatory Commission. (1991a). Code of Federal Regulations, title 10, Rules and Regulations. Washington, DC: Nuclear Regulatory Commission. U.S. Nuclear Regulatory Commission. (1991b) Code of Federal Regulations, title 10, part 35, para 35.32, Energy, Rules and Regulations. Washington, DC: Nuclear Regulatory Commission. U.S. Nuclear Regulatory Commission. (1991c). Regulatory Guide 8.33 (Task DG-8001), Quality Management Program, para 3.1.6 and 3.2.10. Washington, DC: Nuclear Regulatory Commission. Vicini, F. A., V. R. Kini, P. Y. Chen, E. Horwitz, G. Gustafson, P. Benitez, G. Edmundson, N. Goldstein, K. McCarthy, and A. Martinez. (1999). “Irradiation of tumor bed alone after lumpectomy in selected patients with early-stage breast cancer treated with breast-conserving therapy.” J Surg Oncol 70:33–40. Williamson, J. F., G. Ezzell, A. Olch, and B. R. Thomadsen. “Quality Assurance of High Dose Rate Brachytherapy” in High Dose Rate Brachytherapy: A Textbook. S. Nag (ed.), Armonk, NY: Futura Publishing, pp. 147–212, 1994.
Chapter 25
Post-Implant Evaluation William S. Bice, Jr., Ph.D. University of Texas Health Science Center at San Antonio International Medical Physics Services San Antonio, Texas Background . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 477 The Post Plan . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 477 Essential Characteristics of the Treatment-Planning System . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 477 Desirable Characteristics of the Treatment-Planning System . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 478 The Dose Evaluation Hierarchy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 478 Individual Implant Assessment . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 479 Procedural Assessment . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 480 Reporting . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 482 Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 482 Author’s Note . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 482 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 483
Background Post-implant evaluation is performed after the implant procedure. While this may sound inane, or at least obvious, restricting this evaluation to completion of the implant has a purpose. Completion of the implant means after the removal of sources in temporary implant or after the implantation of sources for permanent brachytherapy. The intent is to distinguish post-implant evaluation from implant design or adjustment, performed during the implant procedure. Post-implant evaluation then is performed for two purposes: (1) determining whether the individual implant meets the desired goals and (2) evaluating the performance of the implant team.
The Post Plan The heart of the post-implant evaluation is the post plan. Even though the evaluation includes much more than the radiation dose distribution—examples might be an assessment of methods and techniques or a review of appropriate patient selection, the focus remains on achieving optimal dosimetry. The post plan provides immediate quantitative feedback about the implant, indicative of clinical success and a measure of procedural success.
Essential Characteristics of the Treatment-Planning System Modern treatment planning is based upon three-dimensional (3-D) imagery, typically a set of contiguous axial images. An image set or sets that show the patient’s anatomy as well as the source locations are essential to realistic depiction of implant quality. Knowing where the dose was delivered is just as important as knowing how much dose was delivered. Dose calculations for brachytherapy, because the dose changes so rapidly over short distances, should be performed on a grid of voxels no larger than 2 mm on a side (Yu et al., 1999; Nag et al., 2000). Calculations must be performed in three dimensions using the formalism and values described by AAPM Task Group 43 (TG-43) (Nath et al., 1995) or the more recent update (TG-43U1) (Rivard et al., 2004). Linear and point source calculations should be available where appropriate. For permanent seed implants, where the number of sources precludes individual source
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identification, an automated seed sorting algorithm is essential (Brinkman and Kline 1998; Bice et al., 1999; Li et al., 2001; Liu et al., 2003; Davis et al., 2004; Lam et al., 2004; Holupka et al., 2004). Automated localization and sorting algorithms have improved dramatically since their introduction in the mid-1990s. As a time-saving device their utility is obvious. More importantly, from the perspective of providing high-quality information, the results of these routines are consistent and repeatable. A planning system that can produce a digitally reconstructed radiograph (DRR), or even just an antero-posterior (AP) localization plot of the source locations, provides an extremely useful tool in evaluation of the automated procedures; properly scaled, the sorted seed locations should match an AP radiograph of the patient postimplant. Dose and structures must be available in the same display, preferably with both 2-D and 3-D available. The planning system must be capable of calculating and displaying structure-based dose-volume histograms (DVHs). Individual point placement, calculation, and display are essential.
Desirable Characteristics of the Treatment-Planning System Other tools that may be considered useful for post planning include dose-surface histograms (DSHs), line plots, or dose traces (Anagnostopoulos et al., 2003). Automated generation of quantifiers from the DVH/DSH is a time saver as is automated location of possible source positions. In contrast to seed sorting, where the number of possible source locations is reduced to the number of sources in the implant, automated location of possible source positions is dependent upon contrast differences in the image between the sources and the surrounding tissues. The computer, able to distinguish sources in the image, picks out these possible locations automatically. There has been some work on techniques that allow retention of some of the spatial information that has been lost in the DVH by dividing the target into sectors and generating a DVH for each sector (Bice, Prestidge, and Sarosdy 2001). Projection imagery, digitally reconstructed radiography, is useful for comparing to projection radiographs. Localization from multiple projection radiography can be useful, particularly in situations where this localization is used in conjunction with—fusion with—an image set where anatomy is displayed (Gong et al., 2002).
The Dose Evaluation Hierarchy Figure 1 shows the dose evaluation displays available from the treatment-planning systems described above. Isodose displays show the location of the dose distribution with the size and location of regions of higher and lower dose shown graphically. DVHs are derived from the isodose distributions, removing the spatial information pertaining to the location of the dose within a structure. In return, the DVH reduces the complex information available in the isodose display to a single graph. Further simplification of the information contained in the DVH leads to the generation of quantifiers, single numbers that describe a corresponding characteristic of the dose distribution. Thus there are three tiers to the dose evaluation hierarchy; in order of decreasing information but increasing simplicity, isodose curves, DVHs, and quantifiers. There is a fourth, limited tier that consists of linear dose displays or dose traces. Linear displays, from which quantifiers may also be drawn, sacrifice none of the spatial information. Such displays are appropriate only for linear structures. Examples include the urethra and the neurovascular bundles (DiBiase et al., 2000). Post-implant evaluation at the top of the hierarchy (isodose curves) is usually restricted to analysis of individual implants. For the individual implant it is important to know as much dosimetric information as possible. While it may be desirable to compare quantifiers for single implants, assessing the location of dose is vastly more important. An example of this is the implant that adequately covers the site of positive biopsy results, but does not quite cover the entire planning target volume (PTV). The coverage quantifiers may not be quite up to par, but the brachytherapist may conclude that the coverage is adequate to preclude further therapeutic intervention.
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Information
Analytical tool Isodose displays Dose-Volume Histograms Dose Trace Displays Quantifiers
Spatial information X
Dose information X X
X
X
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Dose analysis X X X
Simplicity Group analysis
X
Figure 1. The dose-evaluation hierarchy. Dose information from an implant can be presented in several ways. Dose calculations are usually performed on a grid of voxels whose resolution depends on the desired accuracy of the calculation. Isodose curves and surfaces can be derived from this volumetric calculation, detailing the location of dose—if anatomical information is available from an image set, the relationship of the dose distribution to the patient’s anatomy can be determined. DVHs contain no spatial information over and above the fact that the stated dose-volume relationship pertains to the structure for which the DVH was generated, i.e., all the calculated dose points are located within the structure. A quantifier is a single number that describes some aspect of the implant, coverage, high dose regions, etc. Because of the large amount of information available from an isodose display, detailed analysis using isodose displays is usually restricted to single implants. Group analysis is better suited to the use of quantifiers due to their simplicity.
Group analysis, that used to evaluate the procedural methods and the skill of the implant team, is usually restricted to the use of quantifiers. A set of well-chosen quantifiers can be very revealing about the institution’s implant procedure, even with just a few implants in the database.
Individual Implant Assessment While the radiation following an implant procedure cannot be taken back, it is still important to assess the dose delivered to the target volume and the dose delivered to any critical structures. Figure 2 shows a paradigm that can be used in evaluating post-implant dosimetry for an individual implant. The decision scheme shown can be used after the post-implant dosimetry has been created. If the target volume received an inadequate dose, or inadequate coverage, then adjuvant therapies should be considered. An additional implant procedure may be considered, an additional application in the case of temporary implants or additional sources implanted in the case of permanent implants. External beam radiotherapy may be used, as well as other adjuvant therapies, chemotherapy or surgery, for instance. This is in contrast to salvage procedures—those performed following a procedure that has failed clinically rather than just dosimetrically. In evaluating the dose to critical structures the task of post-implant analysis becomes, arguably, more difficult. Ideally, no dose would be delivered to normal tissues. In addition to determining the dose to these structures then, the evaluator must determine how much dose is too much. The fuzzy world of risk versus benefits must be inhabited, where bowing to reality is a necessity. Worse, the radiation has been delivered, usually with attendant physical and biological uncertainties. Too some extent the brachytherapist can only watch while the results of his handiwork play out before him. Increased watchfulness can take the form of more frequent follow-up visits. In some cases, the use of therapeutic modalities—even as prophylactic—can be prescribed. An example of this is hyperbaric oxygen used to avoid radiation-induced rectal injury.
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Figure 2. A paradigm that may be used to evaluate an individual implant. This decision tree outlines a technique for conducting an evaluation on a single implant chart. Two evaluations are performed, one to determine the adequacy of coverage of the target volume and the other to adjudge the dose and the effect of dose on the critical structures. A “no” with regard to the first results in consideration of adjuvant therapies or at least a level of increased watchfulness on the part of the caregiver. A “no” concerning acceptable critical structure doses can result in the same increased watchfulness if complications are possible from the doses to these structures. All decisions eventually lead to filing the chart for eventual review as part of a group review (labeled programmatic view in the figure).
Procedural Assessment Analyzing a series of implants to study the quality of the implant program can be of enormous value. Any effort in this regard is worthwhile if the evaluation is performed honestly; publishable results are not the goal. A large database is not required. Series with as few as 10 brachytherapy procedures can still be useful (Bice et al., 1998). A suggested paradigm for evaluating the post-implant dosimetry on a series of implants is shown in Figure 3. Note the dissimilarity between this figure and the technique presented in Figure 2 showing individual implant assessment. In the group assessment shown in Figure 3 implant characteristics and dosimetric results—as described by quantifiers—are categorized for each implant. The choices of the quantifiers and the implant characteristics to evaluate are made by the implant team. The quantifiers are used to discover group traits, the characteristics correlated with the quantifier values in order to determine what kinds of choices can be made to improve the dosimetric results. Two points about this evaluation should be mentioned. First, the list of implant characteristics and implant quantifiers needs to be short. A long, comprehensive list will prove burdensome and the process will not be used, particularly in busy clinic. Additionally. a full set of characteristics and quantifiers has the tendency to be confusing particularly in terms of answering a specific question. Planning and a little thought can promote an intelligent choice of characteristics and quantifiers. Consider the following scenario. Suppose that the brachytherapist notices, anecdotally, that there appears to be an unusual amount of rectal bleeding in his permanent prostate brachytherapy patients
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Figure 3. A paradigm that may be used to evaluate a series of implants. This is a possible technique for conducting a group or programmatic review. Checking for regulatory compliance is separated from the quantifier analysis. This was done to highlight the difference between the two and the intended effect of the quantifier analysis on improving implant quality. While the number of patients has been widened from one to many in comparing the review to the previous one, the scope of the analysis has been severely restricted. The implant team should determine what characteristics and quantifiers to extract. The text describes an example illustrating the utility of a limited scope in performing the post-implant evaluation. The waiting periods shown in the figure can be set as time periods or periods determined by the number of patients treated before another evaluation needs to be performed.
implanted with iodine. A review of the literature shows several papers that correlate dose and rectal symptoms. Choosing one that appears reputable, the brachytherapy team decides to keep track of the volume of rectum that receives more than 160 Gy (R160 Gy). The team can then compare their dosimetric results to those seen at the publishing institution, which recommended keeping R160 Gy to less than 1.3 cm3 to achieve less than a 5% chance of inducing Grade 2 proctitis (Snyder et al., 2001). Now a reasonable choice for implant characteristics to correlate might be needles placed in row one of the middle three columns of the implant template. Note that this kind of analysis can be limited to a dosimetric comparison. It can be done quickly on a relatively small number of patients, looking only at post-implant dosimetry. An analysis of actual bleeding rates would take longer and can be done using this technique, but remember the idea is not to be able to publish the results, but to perform better implants. If a large number of the institution’s patients had R160 Gy greater than 1.3 cm3, the team might choose to modify their implant technique. If the number of needles placed in the offending rows correlated to increased rectal doses, they might want to modify their preplans to move sources further away from the rectum. In this case, it would be appropriate to follow the path shown at the bottom of Figure 3, wait until another series of implants with this new design technique has been collected, and analyze the results on this new series. This finally brings us to the second point. Correlation of quantifier results to implant characteristics need not introduce sophisticated statistical analysis. A Fisher test can be used to perform this correlation, or even a table showing the data can highlight correlations. Just taking averages are often quite telling. A Student’s t-test is simple and can be used. In the example above, the correlation could probably have even been omitted: it is pretty clear that moving needles and sources away from the rectum would lower the dose. Remember, the idea is to improve implant quality, not to display mathematical prowess.
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Reporting From the physicist’s point of view, the implant record must at least contain a written directive; order, shipment, and receipt records; assay and other quality control results; survey results; nursing and patient release instructions and records; pre- and post-planning records and calculation checks; and an implant report (Anderson et al., 1991, Nath et al., 1997). The role of reporting is central to post-implant evaluation efforts. A complete implant report includes the results of post-implant dosimetry. At our institution, as at most, the implant report is part of the special physics consult. From the perspective of the post-implant evaluators, this report is an organizational tool, ensuring completeness of the information conveyed. If properly developed, this report can serve as the sole, or at least the primary, source of data for the post-implant evaluation used in group analysis. Suggested formats for implant reporting change with the type of implant. Not all dose specification methods are applicable to all implant types. Regardless of the form of the dose specification, volumetric, surface or points, the criteria need to be clearly stated, as well as the localization method used. Doses used in the written directive need to state the relationship between the directive and the specified dose. The specification needs to be consistent and reproducible. The format for reporting should conform to that recommended by consensus groups (Nag et al., 2003). Table 1 lists references for reporting formats for different kinds of implants. While every institution should feel free to revise these formats to fit their post-implant-evaluation needs, using the referenced formats will help convey information among institutions in a clear and consistent manner.
Conclusions Post-implant evaluation is essential to determine the quality of the implant, first to assess the outcome of the implant for the patient, second to appraise the implant technique employed by the brachytherapy team. The post plan is key in this evaluation. Proper post planning requires the proper tools, specifically a modern, image-based treatment-planning system. Proper, consistent reporting is essential to post implant evaluation, completely conveying the intent and result of the implant. Group, or programmatic, analysis is greatly facilitated by well-chosen report formats.
Author’s Note This chapter on post-implant evaluation is intentionally general in nature, in an attempt to address all types of implants. A subsequent chapter in this book deals with a specific type of implant, low dose rate (LDR) permanent prostate brachytherapy as an example of post-implant evaluation methods. LDR prostate implants were chosen presumably because of the impact that post-implant evaluation has had
Table 1. Suggested reporting formats. Brachytherapy Procedure
Reporting Format
Gynecological brachytherapy
ICRU 38 (1985)
Gynecological brachytherapy (HDR)
ICRU 38 (1985), Nag et al. (2000)
Interstitial implants
ICRU 58 (1997)
Prostate brachytherapy
Gillin et al. (1990), Yu et al. (1999), Nag et al., (2002)
General Brachytherapy
Anderson et al. (1991), Nath et al. (1997)
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on improving the modality and the available tools. The detail presented in the subsequent chapter can be applied to other types of implants as well.
References Anagnostopoulos, G., D. Baltas, A. Geretschlaeger, T. Martin, P. Papagiannis, N. Tselis, and N. Zamboglou. (2003). “In vivo thermoluminescence dosimetry dose verification of transperineal 192Ir high-dose-rate brachytherapy using CT-based planning for the treatment of prostate cancer.” Int J Radiat Oncol Biol Phys 57(4):1183–1191. Anderson, L., R. Nath, A. J. Olch, and J. Roy. (1991). “American Endocurietherapy Society recommendations for dose specification in brachytherapy.” Endocuriether/Hypertherm Oncol 7:1–12. Bice, W. S., Jr., B. R. Prestidge, and M. F. Sarosdy. (2001). “Sector analysis of prostate implants.” Med Phys 28(12):2561–2567. Bice, W. S., Jr., B. R. Prestidge, P. D. Grimm, J. L. Friedland, V. Feygelman, M. Roach 3rd, J. J. Prete, D. F. Dubois, and J. C. Blasko. (1998). “Centralized multiinstitutional postimplant analysis for interstitial prostate brachytherapy.” Int J Radiat Oncol Biol Phys 41(4): 921–927. Bice, W. S., Jr., D. F. Dubois, J. J. Prete, and B. R. Prestidge. (1999). “Source localization from axial image sets by iterative relaxation of the nearest neighbor criterion.” Med Phys 26(9):1919–1924. Brinkmann, D. H., and R. W. Kline. (1998). “Automated seed localization from CT datasets of the prostate.” Med Phys 25(9):1667–1672. Davis, B. J., D. H. Brinkmann, J. J. Kruse, M. G. Herman, W. N. LaJoie, D. J. Schwartz, T. M. Pisansky, and R. W. Kline. (2004). “Selective identification of different brachytherapy sources, ferromagnetic seeds, and fiducials in the prostate using an automated seed sorting algorithm.” Brachytherapy 3(2):106–112. DiBiase, S. J., K. Wallner, K. Tralins, and S. Sutlief. (2000). “Brachytherapy radiation doses to the neurovascular bundles.” Int J Radiat Oncol Biol Phys 46(5):1301–1307. Gillin, M.T., D.L. Zellmer, D.F. Grimm and K. Sherwood. (1992). “Practical Considerations for Interstitial Brachytherapy” in Advances in Radiation Oncology Physics, Dosimetry, Treatment Planning and Brachytherapy. J. A. Purdy (ed.). AAPM Medical Physics Monograph Number 19. New York: American Institute of Physics, pp. 703-727, 1992. Gong, L., P. S. Cho, B. H. Han, K. E. Wallner, S. G. Sutlief, S. D. Pathak, D. R. Haynor, and Y. Kim. (2002). “Ultrasonography and fluoroscopic fusion for prostate brachytherapy dosimetry.” Int J Radiat Oncol Biol Phys 54(5):1322–1330. Holupka, E. J., P. M. Meskell, E. C. Burdette, and I. D. Kaplan. (2004). “An automatic seed finder for brachytherapy CT postplans based on the Hough transform.” Med Phys 31(9):2672–2679. International Commission on Radiation Units and Measurements (ICRU). Report No. 38. “Dose and Volume Specification for Reporting Intracavitary Therapy in Gynecology.” Bethesda, MD: ICRU, 1985. International Commission on Radiation Units and Measurements (ICRU). Report No. 50. “Prescribing, Recording and Reporting Photon Beam Therapy.” Bethesda, MD: ICRU, 1993. International Commission on Radiation Units and Measurements (ICRU). Report No. 58. “Dose and Volume Specification for Reporting Interstitial Therapy.” Bethesda, MD: ICRU, 1997. International Commission on Radiation Units and Measurements (ICRU). Report No. 62. “Prescribing, Recording and Reporting Photon Beam Therapy (Supplement to ICRU Report 50).” Bethesda, MD: ICRU, 1999. Lam, S. T., P. S. Cho, R. J. Marks II, and S. Narayanan. (2004). “Three-dimensional seed reconstruction for prostate brachytherapy using Hough trajectories.” Phys Med Biol 49(4):557–569. Li, Z., I. A. Nalcacioglu, S. Ranka, S. K. Sahni, J. R. Palta, W. Tome, and S. Kim. (2001). “An algorithm for automatic, computed-tomography-based source localization after prostate implant.” Med Phys 28(7):1410–1415. Liu, H., G. Cheng, Y. Yu, R. Brasacchio, D. Rubens, J. Strang, L. Liao, and E. Messing. (2003). “Automatic localization of implanted seeds from post-implant CT images.” Phys Med Biol 48(9):1191–1203. Nag, S., W. Bice, K. DeWyngaert, B. Prestidge, R. Stock, R., and Y. Yu. (2000). “The American brachytherapy society recommendations for permanent prostate brachytherapy postimplant dosimetric analysis.” Int J Radiat Oncol Biol Phys 46:221–230.
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Nag, S., R. J. Ellis, S. Merrick, R. Bahnson, K. Wallner, and R. Stock. (2002). “American Brachytherapy Society recommendations for reporting morbidity after prostate brachytherapy.” Int J Radiat Oncol Biol Phys 54(2):462–470. Nag, S., R. Dobelbower, G. Glasgow, G. Gustafson, N. Syed, B. Thomadsen, and J. F. Williamson. (2003b). “Inter-society standards for the performance of brachytherapy: A joint report from ABS, ACMP and ACRO.” Crit Rev Oncol Hematol 48(1):1–17. Narayanan, S., P. S. Cho, and R. J. Marks II. (2004). “Three-dimensional seed reconstruction from an incomplete data set for prostate brachytherapy.” Phys Med Biol 49(15):3483–3494. Nath, R., L. L. Anderson, G. Luxton, K. A. Weaver, J. F. Williamson, and A. S. Meigooni. (1995). “Dosimetry of interstitial brachytherapy sources: Recommendations of the AAPM Radiation Therapy Committee, Task Group No. 43.” Med Phys 22:209–234. Also available as AAPM Report No. 51. Nath, R., L. L. Anderson, J. A. Meli, A. J. Olch, J. A. Stitt, and J. F. Williamson. (1997). “Code of practice for brachytherapy physics: Report of the AAPM Radiation Therapy Committee Task Group No. 56.” Med Phys 24(10):1557–1598. Also available as AAPM Report No. 59. Rivard, M. J., B. M. Coursey, L. A. DeWerd, W. F. Hanson, M. S. Huq, G. S. Ibbott, M. G. Mitch, R. Nath, and J. F. Williamson. (2004). “Update of AAPM Task Group No. 43 Report: A revised AAPM protocol for brachytherapy dose calculations.” Med Phys 31:633–674. Also available as AAPM Report No. 84. Snyder, K. M., R. G. Stock, S. M. Hong, Y. C. Lo, and N. N. Stone. (2001). “Defining the risk of developing grade 2 proctitis following 125I prostate brachytherapy using a rectal dose-volume histogram analysis.” Int J Radiat Oncol Biol Phys 50(2):335–341. Yu, Y., L. L. Anderson, Z. Li, D. E. Mellenberg, R. Nath, M. C. Schell, F. M. Waterman, A. Wu, and J. C. Blasko. (1999). “Permanent prostate seed implant brachytherapy: Report of the American Association of Physicists in Medicine Task Group No. 64.” Med Phys 26(10):2054–2076. Also available as AAPM Report No. 68.
Chapter 26
ICRU Recommended Dose and Volume Specifications for Reporting Interstitial Brachytherapy Ali S. Meigooni, Ph.D. Department of Radiation Medicine University of Kentucky, Lexington, Kentucky Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 486 Definitions of Terms and Concepts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 486 Temporary and Permanent Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 486 Source Specification . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 487 Description of Source Patterns . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 487 Total Reference Air Kerma . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 487 Volume and Planes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 488 Gross Tumor Volume (GTV) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 488 Clinical Target Volume (CTV) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 488 Planning Target Volume (PTV) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 488 Treatment Volume . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 489 Central Plane . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 489 Description of Dose Distribution . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 489 General Concepts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 489 Dose Distribution in One or More Planes Through the Implant . . . . . . . . . . . . . . . . . . . . . . . . . . 489 Prescription Dose . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 490 Minimum Target Dose . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 490 Mean Central Dose . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 490 High-Dose Volume . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 493 Low-Dose Volume . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 494 Dose Uniformity Parameters . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 494 Additional Representations of the Dose Distribution . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 494 Time-Dose Factors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 495 General Considerations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 495 Time and Dose Rates for Temporary Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 496 Time-Dose Pattern for Temporary Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 496 Recommendations for Recording and Reporting . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 498 Parameters Required for Recording and Reporting . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 498 Description of Volumes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 498 Description of Sources . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 498 Description of Technique and Source Pattern . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 498 Description of Time Pattern . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 498 Total Reference Air Kerma . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 499 Description of Dose Distribution . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 499 Priority . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 500 Practical Applications of the Recommendations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 500 Temporary Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 500 Permanent Implant . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 500 Single Stationary Source Line . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 500 Moving Sources . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 500 Surface Applicators . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 500 Appendix: Quantities and Units . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 506 Basic Quantities and Units . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 506 Definition of Quantities . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 506
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Introduction One of the important technological developments in brachytherapy over the past two decades was the introduction of miniaturized and highly flexible sources. These sources can be used in afterloading devices with radionuclides of different activities that can produce a wide range of dose distribution within the treatment volume. At the same time, the sophisticated three-dimensional (3-D) source localization methods have been developed in this field and can be linked to computerized methods for dose calculation and representation of dose distribution. These developments have led many clinicians to depart from the longestablished implant systems and to utilize new methodologies for their patient treatments. Therefore, a common language is valuable to provide a method of dose specification and reporting which can be used for implants of all types and can be common to all those involved in interstitial brachytherapy. The International Commission on Radiation Units and Measurements (ICRU) has previously published reports dealing with dose and volume specification for reporting intracavitary therapy in gynecology (ICRU Report 38) (ICRU 1985). However, in 1997 a new report was published to address the problem of absorbed dose specification for reporting interstitial therapy (ICRU Report 58) (ICRU 1997). The aim of report 58 was to develop a common language based on the presently existing concepts. This report introduced concepts to describe what has been done in a way that can be more closely related to the outcome of treatment and one that was generally understood. This chapter reviews the ICRU 58 recommended concepts and procedures for dose specifications and reporting of the interstitial brachytherapy implants.
Definitions of Terms and Concepts Temporary and Permanent Implants Interstitial implants generally fall into two categories, temporary or permanent. In permanent implants, the brachytherapy sources remain in the patient, and they will not be removed. In contrast, in temporary implants, the radioactive sources are removed from the tissue after the treatment is completed. Traditionally, the temporary implants were performed with the linear (wires) radioactive sources or seeds arranged in a linear fashion (i.e., ribbons, etc.), whereas in permanent implants, multiple loose sources were distributed in random orientations. With the design of the low-energy radioactive sources in the form of stranded seeds or radioactive wires, now the permanent implant can also be performed with linear or pre-arranged ribbons. The loose seed implants are considered separately in this report. In planning temporary implants, the total time of implantation depends on the number of sources, their strengths, and the pattern of source distribution within the implant volume. In contrast, in permanent implants, the number of sources depends on their initial strength and type of the radioisotope. Both temporary and permanent implants are occasionally used as boost radiation therapy for external beam irradiation. In that case, the number of seeds and the initial activity of the seeds are adjusted upon the external beam radiation dose. In temporary implants, in the event of a non-ideal source pattern, there may be the possibility of improving the associated dose distribution through manipulation of the dwell time of some sources in the implant. However, at this time, such adjustment is not possible for permanent implants.
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Source Specification The apparent source activity (mCi) and/or equivalent milligram of radium were two commonly used units in specifying the source strength. However, the recent American Association of Physicists in Medicine (AAPM) Task Group 43 (TG 43) (Nath et al. 1995) has recommended using air kerma (kinetic energy released in matter) strength for specification of the brachytherapy source strengths. The ICRU recommendation is also consistent with the TG 43 recommendation for specification of source strength. The reference air kerma rate of a brachytherapy source is the kerma rate to air, in air, at a reference distance of 1 m, corrected for air attenuation and scattering. For this purpose, this quantity is expressed in µGy·h–1 at 1 m, or mGy·h–1 at 1 m. For a linear brachytherapy source, the source strength is defined as the reference air kerma rate of the source divided by the equivalent active length of the source. Therefore, the quantity is expressed as µGy·h–1·cm–1. Please note that 1 µGy·h–1 at 1 m is equivalent to 1 cGy·h–1 at 1 cm, which is more commonly used in clinical applications.
Description of Source Patterns Since all the implants are essentially irradiating a volume of tissues, the term “volume implant” should not be used to describe a specific implant. Therefore, a more accurate description of the source pattern is briefly described as follows: 1. A single-plane implant is defined as an implant containing two or more sources that lie in the same plane. In some instances, the sources lie in a single curved surface. 2. A two-plane implant contains two planes, which are generally parallel to each other. 3. A multiplane implant contains three or more planes, and can often be described according to the number of planes of sources used. 4. A non-planar implant, which is not formed in recognizable planes, may be described by the location of the sources relative to a plane passing through the center of the implant or by a specific geometrical configuration (e.g., sphere or cylinder).
Total Reference Air Kerma The total reference air kerma (TRAK) is the sum of the products of the reference air kerma rate (Si) and the irradiation time (ti) for each source, as: TRAK = ∑ ti ⋅ Si .
(1)
This quantity is analogous to mg·h, which is proportional to the integral dose to the patient. Also, TRAK can serve as a useful index for radiation protection of personnel. The simple determination of the total reference air kerma does not, however, allow one to derive, even approximately, the absorbed dose in the immediate vicinity of the sources (i.e., in the tumor or target volume).
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Volume and Planes The gross tumor volume (GTV) and clinical target volume (CTV) in brachytherapy are defined entirely based on general oncological principles. Therefore, these definitions are identical to those given for external beam radiotherapy (see ICRU Report 50) (ICRU 1993). Gross Tumor Volume (GTV) The GTV is the gross palpable or visible/demonstrable extent and location of the malignant growth. According to this definition, there is no GTV after complete “gross” surgical resection. There is no GTV when there are only a few individual cells or “subclinical” involvement. Clinical Target Volume (CTV) The CTV is defined as the volume of the tissue that contains a GTV and/or subclinical microscopic malignant disease which has to be eliminated. Therefore, this volume of tissue has to be treated adequately in order to achieve the aim of therapy: cure or palliation. The CTV must always be described, independent of the dose distribution, in terms of the patient’s anatomy and the tumor volume. The CTV is a volume of tissue that needs to be irradiated, according to a specified dose-time pattern. As a minimum, the physical dimensions of the CTV are described in terms of its maximum dimension (cm) in three orthogonal directions (see Figure 1). Unlike the external therapy, the CTV in interstitial brachytherapy is sometimes selected at the time of implantation, on the assumption that it is contained within the minimum target isodose (see the section herein, Description of Dose Distribution, Minimum Target Dose). This procedure is not recommended by ICRU. The CTV should be clearly described in the patient chart before the implant is planned. Planning Target Volume (PTV) In external therapy, in order to ensure that all the tissues included in the CTV receive the appropriate dose, it is necessary to plan for a larger volume to be irradiated. This volume of tissue receiving the irradiation is defined as the planning target volume (PTV). In interstitial brachytherapy, the PTV is, in general, identical to the CTV with very few exceptions. For example, with some techniques in which there are uncertainties of consistency of source positions [high dose rate (HDR), moving sources, fractionated tech-
Figure 1. The physical dimensions of the CTV to be reported are the three maximum diameters measured in orthogonal directions. They are, in general, noted as width W, length L, and height H. (Reproduced from Figure 2-1, ICRU Report No. 58. © 1997, with permission from ICRU.)
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niques] or alteration of source position (permanent implants) during the application, the PTV may be larger than the CTV to take these factors into account. However, in ICRU report 58, the term “CTV” is used rather than “PTV.” Treatment Volume The treatment volume is defined as the volume of tissue that is encompassed by an isodose surface that has been selected or specified by the radiation oncologist for a particular treatment. This isodose surface should, ideally, encompass the CTV. The dose value at this isodose surface is the minimum target dose (see the section herein, Description of Dose Distribution, Minimum Target Dose). Central Plane For source patterns in which the source lines are straight, parallel, of equal length, and with centers, which lie in a plane perpendicular to the direction of the source lines, this plane is the central plane (see the section herein, Description of Dose Distribution, Mean Central Dose). In an actual implant, all source lines may not necessarily be straight, parallel, and of equal length. In such cases, the central plane should be chosen perpendicular to the main direction of the source lines and passing through the estimated center of the implant. For more complex implants, it may be necessary to subdivide the target volume into two or more subvolumes for dose evaluation. In this event, a central plane may be defined for each of these subvolumes (see the section herein, Description of Dose Distribution, Mean Central Dose). The calculation of dose distributions in multiple planes throughout the target volume shows that a variation of a few millimeters, in the position of the central plane, is not critical.
Description of Dose Distribution General Concepts In brachytherapy, the dose distribution is nonhomogenous and includes steep dose gradients and regions of high dose surrounding each source. However, within the volume of the implant, there are regions where the dose gradient approximates a plateau (see Figure 2). In an interstitial implant, the regions of plateau dose are equidistant between adjacent neighboring sources, for sources of identical linear activity. They are regions of local minimum doses. Variations in the dose between the different plateau doses can be used to describe the dose uniformity of an implant. A region of plateau dose is the place where the dose can be calculated most reproducibly and compared easily by different departments. Dose Distribution in One or More Planes Through the Implant Although, in modern computer systems, the 3-D dose distribution can be computed and presented as isodose surfaces, these facilities are not yet available in many departments. In order to provide the minimum of information needed about the dose or dose-rate distribution, the calculation of isodose curves in at least one chosen plane is necessary. Methods to present dose information, either in tabular form or by graphical presentation have been discussed in ICRU report 42 (ICRU 1987). If only one plane is chosen for isodose calculation, the central plane of the implant (as defined in the previous section, Volume and Planes) should be chosen for this purpose. To assess the dose distribution in other areas of the implant, multiple planes for isodose calculation can be chosen, either parallel or perpendicular to the central plane.
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Figure 2. Plateau dose region between radioactive sources. The dose distribution in a plane perpendicular to linear and parallel sources shows a plateau dose region of low-dose gradient. (Reproduced from Figure 2-2, ICRU Report No. 58. © 1997, with permission from ICRU.)
Prescription Dose For purposes of this ICRU report, the prescribed dose is defined as the dose, which the physician intends to give, and enters in the patient’s treatment chart. Depending on the system used, the approach for dose prescription may be different. It is not the intention of this report to encourage users to depart from their normal practice of dose prescription. Minimum Target Dose The minimum target dose (MTD) is the minimum dose at the periphery of the CTV. It should be equal to the minimum dose decided upon by the clinician as adequate to treat the CTV. The minimum target isodose is the isodose surface corresponding to the MTD. It defines the treatment volume and should entirely encompass the clinical target volume (see the previous section, Volume and Planes). The MTD is known in some American centers as the “minimum peripheral dose.” The word “peripheral” is not recommended as being too vague and leading to confusion with the concept of peripheral dose in external therapy referring to the dose to healthy structures outside of the target volume. The MTD is known as the “reference dose” in the Paris system. The MTD is equal to about 90% of the prescribed dose in the Manchester system for interstitial therapy. Mean Central Dose In the field of brachytherapy, the mean central dose (MCD) is taken to be the arithmetic mean of the local minimum doses between sources, in the central plane, or in the central planes if there is more than one. In the case of a single-plane implant, the MCD is, in the central plane, the arithmetic mean of the doses at mid distance between each pair of adjacent source lines, taking into account the dose contribution at that point from all sources in the pattern (see Figures 3a and 3c). In the case of implants with line sources in more than one plane, the MCD is the arithmetic mean of the local minimum doses between each set of three adjacent source lines within the source pattern (Figures 3.b and 4). The minimum dose lies at the intersection of perpendicular bisectors of the sides of the triangles (geometric center) formed by these source lines. This point is equidistant from all three-source lines (see Figure 2).
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(b)
(c) Figure. 3. In an implant where the source lines are rectilinear, parallel, and of equal length, the central plane is perpendicular to the direction of the source lines and passes through their centers. The mean central dose Dm is the arithmetic mean of the local minimum doses D in the plateau dose regions; (a) a planar implant; (b) a two-plane implant; (c) an actual single-plane implant where sources are not rectilinear. (Reproduced from Figure 2-3, ICRU Report No. 58. © 1997, with permission from ICRU.)
In some complex implants, a single central plane may not bisect or even include all the sources. In these cases, a MCD based on one plane can be misleading and it is advisable to subdivide the volume and to choose a separate central plane for each subvolume (see Figure 4). Three practical methods are acceptable for determining MCD. They include the following: 1. In the case of implants with parallel lines, identify triangles consisting of three adjacent source lines for all the sources, so that the triangles formed constitute as many acute triangles as possible. Determine the intersection points of the perpendicular bisectors of each triangle and calculate the local minimum dose at each of these points. The mean of these local minimum doses is the MCD (see Figures 3 and 4). This method is the most precise one for parallel lines. 2. Evaluation of dose profiles: Calculate dose profiles for one or more axes through the center of the implant expected to pass through as many local minima as possible. Determine, by inspection, the local minimum doses. The mean of these local minimum values is the MCD (see Figure 5). In a single surface implant performed following a curved surface, a profile may lead to an underestimation of the MCD. In a complex implant, it may be difficult to find axes passing through the minima and profiles may lead to an overestimation of the MCD. However, examples show that
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Figure 4. Two central planes (a) for the longest source lines and (b) for the shortest ones, in an implant. Two mean central doses are determined in the two subvolumes (a) Dma and (b) Dmb. Open circles are the intersections of sources with central plane, and closed circles are the points where the local minimum doses are calculated. (Reproduced from Figure 2-4, ICRU Report No. 58. © 1997, with permission from ICRU.)
(a)
(b)
Figure 5. Three profiles (b) are drawn along two orthogonal directions through a two-plane implant (a) with 8 parallel line sources, 10 cm long, 1.8 cm spacing. The profiles are calculated in percentage of the minimum target dose (thick line) along axes XX, YY, and Y’Y’ in the central plane. The profile along the axis YY is the most representative to estimate the MCD. The mean of the local minimum doses is the MCD. (Reproduced from Figure 2-5, ICRU Report No. 58. © 1997, with permission from ICRU.)
the error lies within acceptable limits. This method is some times preferred for seed implants. In a seed implant, such as the one presented in Figure 6, the dose should be calculated along several random profiles passing through the implant. 3. Inspection of dose distribution: Plot the dose distribution in the central plane. With isodose lines varying by 5% (at most 10%) of the local dose in the central region, the local minima can be deter-
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mined by inspection. The mean of these local minima is the MCD (see Figure 7). This method is often preferred for complex implants with line sources. High-Dose Volume In order to correlate radiation dose with late damage, the high-dose volumes around sources should be assessed. There will inevitably be a high-dose zone around each source. Although it is often small and well tolerated, the exact tolerance dose and volume for interstitial therapy are not known. However, it is necessary, for intercomparison purposes, to agree on a way to describe the high dose volumes. It is suggested that a dose of approximately 100 Gy is likely to be the amount used in determining late effects. In those patients who receive 50 to 60 Gy as minimum target dose or 60 to 70 Gy as MCD, 100 Gy corresponds approximately to 150% of the MCD. It is therefore recommended that the size of the region receiving more than 150% of the MCD be reported. This will indicate that the high-dose volumes should be defined as the volumes encompassed by the isodose corresponding to 150% of the mean central dose around the sources in any plane parallel to the central plane where a high-dose region is suspected. The maximum dimension of the largest region in all planes calculated should be reported (see Figure 8).
Figure 6. The central plane is perpendicular to the main direction of the lines of implantation and passes through the center of the implant. (Reproduced from Figure 2-6, ICRU Report No. 58. © 1997, with permission from ICRU.)
Figure 7. Dose distribution in the central plane of an implant with six parallel iridium line sources, 6 cm long, 1.5 cm spacing, reference air kerma-rate 14.5 µGy·h–1 at 1 m. The dose varies by 5% between plotted isodose lines in the region of interest (A, B, C, D). The local minima, A, B, C, and D, can be easily estimated by inspection. DA and DD approximate 31 cGy·h–1 and DB and DC approximate 34 cGy·h–1. The estimated mean central dose is Dm = 32.5 cGy·h–1. (Reproduced from Figure 2-7, ICRU Report No. 58. © 1997, with permission from ICRU.)
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(a)
(b)
Figure 8. Dose distributions in two planes with three nonparallel loops loaded with seed ribbons. In (a), in a plane YY’ parallel to the main direction of the implant, the dose distribution does not display any unexpected high-dose region. In (b), in the central plane XX’ of the implant, the maximum dimension of the 150% isodose line (dotted line), varies from 6 mm to 27 mm. In lower perpendicular planes, the dimension of the high-dose region increases slowly to 33 mm. (Reproduced from Figure 2-8, ICRU Report No. 58. © 1997, with permission from ICRU.)
Low-Dose Volume A low-dose volume should be defined as a volume within the CTV, encompassed by an isodose corresponding to 90% of the prescribed dose. The maximum dimension of the low-dose volume in any plane calculated should be reported. In implants where the CTV is included within the minimum target isodose, the occurrence of a lowdose region is exceptional. If the CTV is not covered by the minimum target isodose, there will be low-dose regions due to geographical miss. To correlate the local recurrence rate with the dose distribution, it is recommended that low-dose volumes be reported. Dose Uniformity Parameters Several indices quantifying the homogeneity of the dose distribution have been proposed (see, for example, Paul, Koch, and Philip 1988; Wu, Ulin, and Sternick 1988; Saw and Suntharalingam, 1991). In this ICRU report, two parameters describing dose uniformity for interstitial implants are recommended. They can be derived directly from the concepts of MTD and MCD (see Figure 9): 1. The spread in the individual minimum doses used to calculate the MCD in the central plane expressed as a percentage of the MCD. 2. The dose homogeneity index (DHI), defined as the ratio of MTD to the MCD. Additional Representations of the Dose Distribution To obtain a full perception of the dose distribution of an implant, the use of volume-dose calculations has been advocated (see, for example, Neblett et al., 1985; Anderson 1986; Bridier et al., 1988; McCrae,
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Figure. 9. Dose distribution in the central plane of a two-plane breast implant, with seven line sources, 10 cm long, 2 cm spacing, 90 cGy·h–1 at 1 m. The MCD is 70.9 cGy·h (the local minima in cGy·h are DA = 65.4, D = DD = 74.4, and D = 75.3). The MTD (100%) is 58.1 cGy·h–1. The spread in the individual minimum doses is from –8% to +6%. The dose homogeneity index, expressed as the ratio of the minimum target dose and the mean central dose, is 0.82 = 58.1/70.9. (Reproduced from Figure 2-9, ICRU Report No. 58. © 1997, with permission from ICRU.)
Rogers, and Dritschilo 1987). For this purpose, the CTV (or a larger volume including an additional margin) is subdivided in subvolumes (e.g., voxels) and the dose rate is calculated at the center of each subvolume. The volume receiving at least a specified dose is then defined as the sum of all subvolumes where, at the center at least, that dose is received. Examples of results are shown in Figure 10. Because of high-dose gradients, significant differences in calculated volumes can be observed, depending upon the size of the elementary subvolumes. The size of the grid and of the elementary subvolumes (voxels) used in dose and volume calculations should be clearly stated. Volume-dose data can also be represented by means of histograms, showing the distribution of fractions of the CTV receiving doses within chosen intervals. The value of these alternative representations of the dose distribution as possible prognostic factors for treatment outcome has still to be established in clinical research.
Time-Dose Factors General Considerations Considerable experience has been gained over many years with conventional dose rates. For removable implants, 60 to 70 Gy has usually been delivered in 4 to 8 days at a dose rate of 30 to 90 cGy·h. For permanent implants with 125I, doses of 120 to 160 Gy have been delivered with 50% of the dose received in the first 2 months and the majority of the remainder over the succeeding 6 months. Even with these conventional treatments, it has been recognized that the dose rates within and adjacent to the target volume vary considerably as a function of the distance from the sources and that these variations may be significant in determining effects on both tumor and normal tissues. The development of new afterloading techniques and, in particular, the use of HDR introduces new dose-time patterns that require evaluation. These include: • Continuous low dose rate (LDR); with short scheduled or unscheduled interruptions • Single moving source HDR used to treat several channels of an implant over several days to simulate continuous LDR (See the next section under Time-Dose Pattern for Temporary Implants, Pulsed irradiation.) • HDR; in a single fraction • Fractionated HDR
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Figure 10. Volume (sum of subvolumes receiving at least a certain dose) vs. dose, for two different patterns of parallel source lines: a two-plane implant with six sources 5 cm long (upper curve), and a cylindrical implant with seven sources 4 cm long (lower curve). The dose is expressed in percentage of the minimum target dose. The size of the voxel used for calculations is 1 mm (Bridier et al., 1988). (Reproduced from Figure 2-10, ICRU Report No. 58. © 1997, with permission from ICRU.)
• Combinations of external beam radiation with any type of brachytherapy with varying intervals between the two The overall treatment time for brachytherapy, and the duration of the interval(s) between treatments, can have an important effect on outcome. Therefore, the dose-time pattern should be recorded, and it is sometimes necessary to group different patterns for the purpose of analysis of outcome and intercomparisons. Time and Dose Rates for Temporary Implants Irradiation time is the time during which a radioactive source is present in the patient. The overall treatment time is the total time elapsed from the beginning of the first irradiation to the end of the last one. Instantaneous dose rate is the quotient of the dose and the irradiation time, for a given fraction or pulse. Average overall treatment dose rate is the quotient of the total dose and the overall treatment time. Average overall treatment dose rate is a concept useful for continuous LDR irradiations with or without short interruptions and for some pulsed irradiations (see the next section). Time-Dose Pattern for Temporary Implants Continuous irradiation. The overall treatment time does not differ from the irradiation time: only the instantaneous dose rate is considered. When the irradiation times of individual sources are different, the instantaneous dose rate varies with time and the average overall treatment dose rate is, in general, meaningless (Figure 11a). Noncontinuous irradiation. With the advent of remote afterloading, in most instances the overall treatment time is greater than the total irradiation time (which is the sum of the partial irradiation times) due to incidental or planned short interruptions during the treatment. The instantaneous dose rate is greater than the average overall treatment dose rate. In LDR irradiations, when the duration of one given
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Figure 11. Overall treatment time and irradiation time for different types of treatment: A, continuous irradiation; B, noncontinuous irradiation; C, fractionated irradiation; D, hyperfractionated irradiation; E, pulsed irradiation in two fractions. (Reproduced from Figure 2-11, ICRU Report No. 58. © 1997, with permission from ICRU.)
interruption is longer than 10% of the total irradiation time, the irradiation should be considered “fractionated” irradiation. In this type of treatment, irradiation time is subdivided into multiple fractions. In fractionated irradiation, the overall treatment time is much greater than the total irradiation time. For fractionated irradiation, the instantaneous dose rate is the ratio of the dose per fraction and the irradiation time per fraction, and the aver age overall treatment dose rate is, in general, meaningless. Although the time interval between fractions is usually of the order of magnitude of a day or days, an LDR irradiation is considered fractionated when one given interruption is longer than 10% of the total irradiation time. The special case of multiple short irradiations with HDR source(s) is considered in the following paragraph. Hyperfractionated irradiation. When two or more fractions are given per day, the irradiation is considered a “hyperfractionated” irradiation. When the time interval between short HDR irradiations reaches or exceeds 4 hours, the irradiation should be considered a hyperfractionated irradiation. It should be considered as fractionated when the time interval is equal to one or several days. Pulsed irradiation. When a single HDR source is used to give a sequence of short irradiations (pulses) to simulate continuous LDR irradiation, the irradiation should be considered a “pulsed” irradiation as long as the time interval is shorter than 4 hours.
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Recommendations for Recording and Reporting It is recommended that adequate information be recorded to give a consistent description of any implant. The guidelines for reporting dose will make it possible to compare results of future brachytherapy practice and to better relate outcome to treatment. To report an implant the following should be recorded.
Parameters Required for Recording and Reporting Description of Volumes The description of the volumes should include, as a minimum, the GTV, the CTV, and the treatment volume. Description of Sources The description of the sources employed should include details of: 1. Radionuclide used, including filtration, if relevant. 2. Type of source used, i.e., wire, seeds, seed ribbon, hairpin, needle, etc. 3. Length of each source line used. 4. Reference air kerma rate of each source or source line. 5. The distribution of the strength within the source should be described (uniform or differential loading, etc.). Description of Technique and Source Pattern If the source distribution rules of a standard system have been followed, this shall be specified; if not, the source pattern should be described as explained previously in the section, Description of Source Patterns In addition, the following data should also be recorded: 1. Number of sources or source lines. 2. Separation between source lines and between planes. 3. Geometrical pattern formed by the sources with the central plane of the implant (e.g., triangles, squares), where relevant. 4. The surfaces in which the implant lies, i.e., planes or curved surfaces. 5. Whether crossing sources are placed at one or more ends of a group of linear sources. 6. The material of the inactive vector used to carry the radioactive sources, if any (e.g., flexible or rigid), whether rigid templates are used at one or both ends. 7. Type of remote afterloading, if used. Description of Time Pattern The description of the time pattern should include the type of irradiation with the necessary data on treatment and irradiation times as described below. The information on dose and time should provide the necessary data to calculate instantaneous and average dose rates. 1. Continuous irradiation: the overall treatment time should be recorded.
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2. Noncontinuous irradiation: both the overall treatment time and the total irradiation time should be recorded, together with information about lengths of gaps. 3. Fractionated and hyperfractionated irradiation: the irradiation time of each fraction or pulse, the interval between fractions or pulses, and the overall treatment time should be recorded. 4. When the irradiation times of the different sources are not identical, they should be recorded. 5. Moving sources: a. Stepping sources: Step size and dwell time should be recorded if constant. Variation of the dwell times of a stepping source can be used for manipulating the dose distribution. If such a dose optimization is applied, this should be specified (e.g., optimization on dose points defined in the implant or geometrical optimization (Kolkman-Deurloo et al., 1994). For pulsed irradiation, at least two statements of dose rate may be necessary. One is the “pulse-average dose-rate,” which is the quotient of the pulse dose by the time from beginning to end of the pulse. The other is the maximum local dose rate at 1 cm from the stepping source. b. Oscillating sources: Speed in different sections of the vectors should be recorded. Total Reference Air Kerma The Total Reference Air Kerma (TRAK) for the total irradiation time should be recorded (see the previous discussion in the Definitions of Terms and Concepts section herein). Description of Dose Distribution The following doses should be recorded: 1. Prescribed Dose. If the dose is not prescribed at the level of either the MTD or the MCD, the method of dose prescription should be recorded. If, for clinical or technical reasons, the dose received differs from the prescribed dose, it should be noted (see the section Description of Dose Distribution, Prescription Dose). 2. The MTD should be recorded (see the section Description of Dose Distribution, Minimum Target Dose). 3. The MCD should be recorded (see the section Description of Dose Distribution, Mean Central Dose). 4. The following additional information, when available, should be recorded: a. Dimension of high-dose volume(s) (see the section Description of Dose Distribution, High-Dose Volume). b. Dimension of any low-dose volume (see the section Description of Dose Distribution, LowDose Volume). c. Any dose uniformity data (see the section Description of Dose Distribution, Dose Uniformity Parameters). d. Additional representation of dose distribution, if any (see the section Description of Dose Distribution, Additional Representations of the Dose Distribution).
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Priority Three levels of priority are recognized for reporting an interstitial therapy application. They are linked to the different levels of dose computation sophistication needed to fulfill the reporting requirements (Visser 1989). The wide variation in the availability of treatment-planning systems is recognized and taken into account in Tables 1 to 5.
Practical Applications of the Recommendations It is not the intention of this report and is not the task of the ICRU to encourage radiation oncologists to depart from their current practice of dose prescription or technique of application. Application of the reporting recommendations to existing systems and techniques is developed below.
Temporary Implants The recommended hierarchy of dose reporting, for a temporary implant is presented in Table 1.
Permanent Implant The majority of permanent implants are not done with recognizable source lines and it is therefore difficult to identify a central plane or to calculate a mean central dose. The recommended hierarchy of dose reporting for a permanent implant is presented in Table 2.
Single Stationary Source Line The source can be intraluminal or sometimes interstitial: LDR or HDR techniques can be applied. The recommended hierarchy of dose reporting, for a single source line, is presented in Table 3.
Moving Sources In addition to simulating a uniform line source, a moving source can be used to modify the dose distribution by changing the dwell time between moves or the speed of movement. The hierarchy of dose reporting of a moving source takes the form presented in Table 4.
Surface Applicators Although surface applicators are not interstitial implants, the physical factors which govern dose distribution from surface applicators are similar. At present, surface applicators are most commonly used for treating lesions involving skin or mucosal surfaces and the choroidal layer of the eye. When describing treatment by a surface applicator, the recommended hierarchy of reporting for surface applicators is presented in Table 5.
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Table 1. Levels of Priority for Reporting Temporary Interstitial Implants Parameters for reporting temporary interstitial implants Description of Volume Gross Tumor Volume Clinical Target Volume Treated Volume Description of Source and Technique Radionuclide, type of source Source size and shape, source pattern Reference air kerma rate Inactivity vector (applicator), if any Description of Time Pattern Total Reference Air Kerma Description of Dose Prescribed dose including point or surface of prescriptionc. Reference Dose in Central Plane a. Mean Central Dose b. Minimum Target Dose Description of High and Low Dose Volume Uniformity Parameters Alternative Representation of Dose Distribution Dose Rates at Point or Surface of Prescription a
Prioritya
Levelb of computation
1 1 1 1
1 1 3 1
1 1
1 1
1
1
1 2 3 3 3
2 2 3 3 3 3
Priority 1. Concerned with doses in the central plan. 2. Required calculations outside the central plan. If this is not available, then a more detailed description of source pattern under priority 1 is required. 3. Additional information mostly of clinical research interest. b Level of computation 1. No computer needed. 2. Hand calculation and/or computer calculation in central plane. 3. 3-D computation needed. c Essential to establish consistent reporting and related to past experience, necessary for comparison of brachytherapy data and for relating outcome to treatment. If a classical system is used, the system should be identified. (Reproduced from Table 4.1 of ICRU Report 58, © 1997, with permission of ICRU.)
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Ali S. Meigooni Table 2. Levels of Priority for Reporting Permanent Interstitial Implants Parameters for reporting permanent interstitial implants
Description of Volume Gross Tumor Volume Clinical Target Volume Treated Volume Description of Source and Technique Radionuclide, type of source Source size and shape, source pattern Reference air kerma rate Inactivity vector (applicator), if any Total Reference Air Kermac Description of Dose Prescribed dose including method of prescription Reference Dose in Central Plane c. Mean Central Dose d. Minimum Target Dose Description of High and Low Dose Volume Uniformity Parameters Alternative Representation of Dose Distribution a
Prioritya
Levelb of computation
1 1 1 1
1 1 3 1
1
1
1
1
1 1 2 3 3
2 2 3 3 3
Priority 1. Concerned with doses in the central plan. 2. Required calculations outside the central plan. If this is not available, then a more detailed description of source pattern under priority 1 is required. 3. Additional information mostly of clinical research interest. b Level of computation 1. No computer needed. 2. Hand calculation and/or computer calculation in central plane. 3. 3-D computation needed. c For permanent implants, the TRAK is calculated as the product of the total air kerma rate at the time of the implantation and the decay time constant λ, also known as the mean life time (λ = T/ln 2, where T is the half life). (Reproduced from Table 4.2 of ICRU Report 58, © 1997, with permission of ICRU.)
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Table 3. Levels of Priority for Reporting Implants with a Single Stationary Source Line Parameters for reporting implants with a single stationary source line Description of Volume Gross Tumor Volume Clinical Target Volume Treated Volume Description of Source and Technique Radionuclide Length Shape (Straight/Curved) Reference air kerma rate Strength distribution (uniform linear strength is assumed, if not, distribution must be specified.) Diameter of inactive vector (applicator) Description of Time Pattern Total Reference Air Kerma Description of Dose and Prescription Point (Distance from the source line position along the source line) Minimum target dose if different from prescribed dose Dose at 1 cm from axis of the source line at its center Dose at the surface of applicator in contact with tissue Additional representation of dose distribution Dose Rate Average overall treated dose rate at the point or surface of dose prescription a
Prioritya
Levelb of computation
1 1 1 1
1 1 3 1
1 1 1
1 1 1
1 1 3 3
1 1 3 3
3
1
Priority 1. Concerned with doses in the central plan. 2. Required calculations outside the central plan. If this is not available, then a more detailed description of source pattern under priority 1 is required. 3. Additional information mostly of clinical research interest. b Level of computation 1. No computer needed. 2. Hand calculation and/or computer calculation in central plane. 3. 3-D computation needed. (Reproduced from Table 4.3 of ICRU Report 58, © 1997, with permission of ICRU.)
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Ali S. Meigooni Table 4. Levels of Priority for Reporting Implants with Moving Source Parameters for reporting implants with moving source
Description of Volume Gross Tumor Volume Clinical Target Volume Treated Volume Description of Source and Technique Radionuclide, source type and source Type of movement Range of motion (effective length of source) Applicator-including diameter of inactive vector. Number of inactive vectors Reference air kerma rate Description of Time Pattern Intervals between fractions Irradiation time per fraction Total Reference Air Kerma Description of Dose Prescribed dose Minimum target dose For single source line or bifurcationc, dose at 1 cm For complex implant, mean central dose Method of dose optimization, if applicable Description of high and low dose volume Uniformity Parameters Additional representation of dose distribution Alternative Representation of Dose Distribution Dose Rates at point or surface of prescription a
Prioritya
Levelb of computation
1 1 1 1
1 1 3 1
1
1
1
1
1 1 1 1 2 2
1 2 2 2 3 2
3 3 3
3 3 3
Priority 1. Concerned with doses in the central plan. 2. Required calculations outside the central plan. If this is not available, then a more detailed description of source pattern under priority 1 is required. 3. Additional information mostly of clinical research interest. b Level of computation 1. No computer needed. 2. Hand calculation and/or computer calculation in central plane. 3. 3-D computation needed. c If there is a bifurcation, more than one target volume should be considered. d Continuous/step wise, size of step, unidirectional/oscillating. Uniform motion is assumed; if not, motion must be specified. (Reproduced from Table 4.4 of ICRU Report 58, © 1997, with permission of ICRU.)
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Table 5. Levels of Priority for Reporting Use of Surface Applicators Parameters for reporting use of surface applicators Description of Clinical Target Volume Treated Volume Description of Applicator Shape (flat/curved, round/square, etc.) Size Description of Source Radionuclide and chemical form Concentration (seeds/tubes/plated)c Description of Time Pattern Intervals between fractions Irradiation time per fraction Total Reference Air Kerma Description of Dose Prescription Dose and point of dose prescription Dose at 5 mm in tissue at the center of the applicatord Minimum target dosee Description of high dose at tissue surface in contact with applicator, usually near the center of the applicator. Uniformity Parameters Additional representation of dose distribution Dose Rates Average dose rate at the point of dose prescription a
Prioritya computation
Levelb of
1 1 1
1 3 1
1
1
1
1
1
1
1
1
1 1 1
1 2 2
2
2
3
3
3
3
Priority 1. Concerned with doses in the central plan. 2. Required calculations outside the central plan. If this is not available, then a more detailed description of source pattern under priority 1 is required. 3. Additional information mostly of clinical research interest. b Level of computation 1. No computer needed. 2. Hand calculation and/or computer calculation in central plane. 3. 3-D computation needed. c Including distance from source(s) to the surface of the applicator. d For eye plaque, 5 mm from the internal sclera. e Dose at the distal extent of the clinical target volume. (Reproduced from Table 4.5 of ICRU Report 58, © 1997, with permission of ICRU.)
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Appendix: Quantities and Units Basic Quantities and Units The quantities and units used in brachytherapy are generally related to the dosimetric aspects of this treatment modality. ICRU report 33 (ICRU 1980) has a detail description of these parameters. The following information provides most relevant information for this chapter.
Definition of Quantities 1. Absorbed dose: The mean energy imparted by the radiation to a unit mass of medium. D=
dε dm
The SI unit of absorbed dose is J.kg–1. The special unit of the absorbed dose is gray (Gy) where 1 Gy = 1 J·kg–1. The subunit of absorbed dose is cGy, where 1 cGy = 1/100 Gy. 2. Kerma (kinetic energy released in matter): The sum of the initial kinetic energies of all the ionizing charged particles liberated by uncharged ionizing particles. K=
dEtr dm
The SI unit of kerma is J·kg–1. The special unit of the kerma is the gray (Gy) where 1 Gy = 1 J·kg–1. 3. The activity, A, is the Number of disintegration of a radioactive nucleus per second (dps). A=
dN dt
The SI unit of kerma is s–1. The special unit of the activity is the becquerel (Bq) where 1 Bq = 1 s–1. The more commonly used unit of activity is the curie (Ci) where 1 Ci = 2.7 × 1010 dps. 4. The air kerma-rate constant, Γδ, of a radioactive nuclide emitting photon is defined as: Γδ =
" 2 K# δ
,
A
where K# δ is the air kerma rate due to photons of energy greater than δ, at a distance, ", from a point source, with activity of A. The SI unit of air kerma constant is J.kg–1m2. The special unit of the air kerma constant is U, where 1 U = 1 µGy.m2.h–1 = 1 cGy.cm2.h–1.
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Relationship Between the Quantities Absorbed dose and kerma in a reference material m
µ Dm = K air tr (1 − g ) ρ air where g is the small part of the kerma which is lost to bremsstrahlung in medium, and (µtr/r ) is the energy transfer coefficient of the radiation for a specific medium.
Reference Air Kerma Rate At the time when radium was used as brachytherapy sources, their strength was determined by measuring the mass of the radium. Subsequently, 1 gram of radium was specified as 1 Ci. As other radioactive materials become available, the concept of apparent activity (the activity of an unencapsulated point source, which could produce the same exposure rate as the capsulated source, at a given point) has been introduced. However, for low-energy photon emitters, the accuracy of the geometric structure of the sources was affecting this parameter and introducing some discrepancies among various dosimetric information. To eliminate these discrepancies, air-kerma rate was introduced by the AAPM (AAPM, 1987; Nath et al. 1995) for brachytherapy source determinations. The unit of air kerma strength is U, where where 1U = 1µGy·m2·h–1 = 1cGy·cm2·h–1.
References American Association of Physicists in Medicine (AAPM). Report No. 21. “Specification of Brachytherapy Source Strength.” New York: American Institute of Physics, New York, 1987. Anderson, L. L. (1986). “A ‘natural’ volume-dose histogram for brachytherapy.” Med Phys 13:898–903. Bridier, A., H. Kafrouni, J. P. Houlard, and A. Dutreix. “Comparison des distributions de dose en curietherapie interstitielle autour des sources continues et discontinues.” Tnt. Symp. on Dosimetry in Radiotherapie. IAEA SM-298/23. Vienna: IAEA, 1988. International Commission on Radiation Units and Measurements (ICRU). Report No. 33. “Radiation Quantities and Units.” Bethesda, MD: ICRU, 1980. International Commission on Radiation Units and Measurements (ICRU). Report No. 38. “Dose and Volume Specification for Reporting Intracavitary Therapy in Gynecology.” Bethesda, MD: ICRU, 1985. International Commission on Radiation Units and Measurements (ICRU). Report No. 42. “Use of Computers in External Beam Radiotherapy Procedures with High-Energy Photons and Electrons.” Bethesda, MD: ICRU, 1987. International Commission on Radiation Units and Measurements (ICRU). Report No. 50. “Prescribing, Recording and Reporting Photon Beam Therapy.” Bethesda, MD: ICRU, 1993. International Commission on Radiation Units and Measurements (ICRU). Report No. 58. “Dose and Volume Specification for Reporting Interstitial Therapy.” Bethesda, MD: ICRU, 1997. Kolkman-Deurloo, I. K., A. G. Visser, C. G. Niel, N. Driver, and P. C. Levendag. (1994). “Optimization of interstitial volume implants.” Radiother Oncol 31:229–239. McCrae, D., J. Rogers, and A. Dritschilo. (1987). “Dose-volume and complication in interstitial implants for breast carcinoma.” Int J Radiat Oncol Biol Phys 13:525–529. Nath, R., L. L. Anderson, G. Luxton, K. A. Weaver, J. F. Williamson, and A. S. Meigooni. (1995). “Dosimetry of interstitial brachytherapy sources: Recommendations of the AAPM Radiation Therapy Committee, Task Group No. 43.” Med Phys 22:209–234. Also available as AAPM Report No. 51. Neblett, D., S. A. M.Nisar, A. A. Puthawala, R. Habrop, H. S. Frey, and S. E. Hogan. (1985). “An interstitial implant technique evaluated by contiguous volume analysis.” Hypertherm Oncol 1:213–221.
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Paul, J. M., R. F. Koch, and P. C. Philip. (1988). “Uniform analysis of dose distribution in intersti tial brachytherapy dosimetry systems.” Radiother Oncol 13:105–125. Saw, C. B., and N. Suntharalingam. (1991). “Quantitative assessment of interstitial implants.” Int J Radiat Oncol Biol Phys 20:135–139. Visser, A. G. (1989). “An intercomparison of the accuracy of computer planning systems for brachytherapy.” Radiother Oncol 15:245–258.
Bibliography American Association of Physicists in Medicine (AAPM). Report No. 41. “Remote Afterloading Technology.” New York: American Institute of Physics, New York, 1993. Bernard, M., B. Guille, and G. Duvalet. (1975). “Mesure du debit d’exposition linéique nominal des sources a une dimension, utilisées en curiethérapie.” J Radiol Electrol 56:785–790. British Institute of Radiology (BIR). “Recommendations for Brachytherapy Dosimetry.” Report of a joint working party of the BIR and the Institute of Physical Sciences in Medicine (IPSM). London: British Institute of Radiology, 1993. British Committee on Radiation Units and Measurements (BCRU). (1984). “Specification of brachytherapy sources, memorandum from the British Committee on Radiation Units and Measurements.” Br J Radiol 57:941–942. Boutillon, M. “Values of g for Photon Energies” in CCEMRI Report (1)185-18. Paris: Consultative Committee for Standards of Measuring Ionizing Radiations, 1985. Cance, M., and J. P. Simoen. (1983). “Etalonnage des sources de curiethérapie” Compte-Rendus du Congrès de la Société Française des Physiciens d’Hépitaux. Paris: French Society of Hospital Physicists (SFPH), pp. 87–104, 1983. Comité Francais de Mesure des Rayonnements Ionisants (CMFRI). “Recommendations pour la determination des doses absorbees en curiethérapie. Rapport du Comite Francais ‘Mesure des Rayonnements Ionisants’ No. 1.” Paris: Bureau National de Metrologie, 1983. Dutreix, A., and A. Wambersie. (1968). “Étude de la repartition des doses autour de sources ponctu elles alignées.”Acta Radiol 7:389–400. Dutreix, A., and A. Wambersie. (1975). “Specification of gamma-ray brachytherapy sources.” Br J Radiol 48:1034. Dutreix, A., G. Marinello, and A. Wambersie. Dosimetrie en Curiethérapie. Paris: Masson, 1982. Goetsch, S. J., F. H. Attix, D. W. Pearson, and B. R. Thomadsen. (1991). “Calibration of 192Ir high dose-rate afterloading systems.” Med Phys 18:462–467. International Commission on Radiation Units and Measurements (ICRU). Report No. 24. “Determination of Absorbed Dose in a Patient Irradiated by Beams of X or Gamma Rays in Radiotherapy Procedures.” Bethesda, MD: ICRU, 1976. International Commission on Radiation Units and Measurements (ICRU). Report No. 29. “Dose Specification for Reporting External Beam Therapy with Photons and Electrons.” Bethesda, MD: ICRU, 1978. International Commission on Radiation Units and Measurements (ICRU). Report No. 35. “Radiation Dosimetry: Electron Beams with Energies Between 1 and 50 MeV.” Bethesda, MD: ICRU, 1984. Kutcher, G. J., L. Coia, M. Gillin, W. F. Hanson, S. Leibel, R. J. Morton, J. R. Palta, J. A. Purdy, L. E. Reinstein, G. K. Svensson, M. Weller, and L. Wingfield. (1994). “Comprehensive QA for radiation oncology: Report of AAPM radiation therapy committee task group 40.” Med Phys 21:581–618. Also available as AAPM Report No. 46. Ling, C. C., and Z. C. Gromadzki. (1981). “Activity uniformity of 192Ir seeds.” Int J Radiat Oncol Biol Phys 7:665–669. Ling, C. C., E. D. Yorke, I. J. Spiro, D. Kubiatowicz, and D. Bennett. (1983). “Physical dosimetry of I-125 seeds of a new design for interstitial implant.” Int J Radiat Oncol Biol Phys 9:1747–1752. Marinello, G., M. Valéro, S. Leung, and B. Pierquin. (1985). “Comparative dosimetry between iridium wires and seed ribbons.” Int J Radiat Oncol Biol Phys 11:1733–1739. Meredith, W. J. (ed.). Radium Dosage: The Manchester System. Edinburgh: Livingston, 1967. National Council on Radiation Protection and Measurements (NCRP). Report No. 41. “Specification of GammaRay Brachytherapy Sources.” Bethesda, MD: NCRP, 1974.
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Netherlands Commission on Radiation Dosimetry (NCS). Report No. 4. “Recommendations for Dosimetry and Quality Control of Radioactive Sources Used in Brachytherapy.” Bilthoven: (Netherlands Commission on Radiation Dosimetry, 1991. Pierquin, B., A. Dutreix, C. H. Paine, D. Chassagne, G. Marinello, and D. Ash. (1978). “The Paris system in interstitial radiation therapy.” Acta Radiol Oncol 17:33–48. Pierquin, B., J. F. Wilson, and D. Chassagne. Modern Brachytherapy. New York: Masson, 1987. Quimby, E. H., and V. Castro. (1953). “The calculation of dosage in interstitial radium therapy.” Am J Roentgenol 70:739–749. Sralek, R. J., and M. Stovall. “Brachytherapy Dosimetry” in The Dosimetry of Ionizing Radiation. K. R. Kase, B. E. Bjarngard and F. H. Attix (eds.). New York: Academic Press, pp. 259–321, 1990. Steggerda, M. J., and B. J. Mijnheer. (1994). “Replacement corrections of a Farmer-type ionization chamber for the calibration of Cs-137 and Ir-192 sources in a solid phantom.” Radiother Oncol 31:76–84. Thomason, C., T. R. Mackie, and M. J. Lindstrom. (1991). “Effect of source encapsulation on the energy spectra of 192 Ir and 137Cs seed sources.” Phys Med Biol 36:496–506. Venselaar, J. L., W. F. Brouwer, B. H. van Straaten, and A. H. Aalbers. (1994). “Intercomparison of calibration procedures for Ir-192 HDR sources in The Netherlands and Belgium.” Radiother Oncol 30:155–161. Williamson, J. F. (1988). “Monte Carlo evaluation of specific dose constants in water for 125Iseeds.” Med Phys 15:686–694. Williamson, J. F., F. M. Khan, S. C. Sharma, and G. D. Fullerton. (1982). “Methods for the routine calibration of brachytherapy sources,” Radiology 142:511–516. Wu, A., K. Ulin, and E. S. Sternick. (1988). “A dose homogeneity index for evaluating Ir-192 interstitial implants.” Med Phys 15:104–107. Young, M. E. J., and H. F. Batho. (1964). “Dose tables for linear radium sources calculated by an electronic computer.” Br J Radiol 37:38–44.
Chapter 27
Intraoperative Radiation Therapy (IORT) Gil’ad N. Cohen, M.S., and Marco Zaider, Ph.D. Department of Medical Physics Memorial Sloan-Kettering Cancer Center New York, New York Clinical Overview . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 511 Hardware. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 512 Afterloading Devices . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 512 Applicators . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 512 Cost of IORT Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 513 Treatment Planning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 513 Dose Specification . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 513 Treatment-Planning Approaches . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 514 Computer-Assisted Plan Verification . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 514 Determination of Implant Geometry . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 515 Dose Calculation. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 516 Classical Dosimetric Systems. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 516 QA Considerations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 517 Facility . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 517 Design and Shielding Considerations. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 517 Personnel Requirements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 519 Emergency Procedures . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 519 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 519
Clinical Overview Ideally, radiation treatments would deliver a curative dose to the tumor and completely spare surrounding healthy tissue. In the clinic, however, delivering the dose required for local tumor control may result in unacceptable normal tissue complications, in effect placing an upper limit on the dose deliverable using external beam radiation therapy (EBRT). This is of particular relevance for locally recurrent tumors, some of which may have been treated previously, and pediatric tumors, where normal tissue toxicity requires special consideration.Various treatment techniques have been proposed to overcome this limitation, including external beam (EB) intensity-modulated radiation therapy (IMRT) and brachytherapy. Briefly, IORT refers to a single fraction treatment delivered to a surgically exposed target area. Two competing approaches to IORT are currently in practice: the first is linac based using electron beams; the second employs a high dose rate (HDR) 192Ir afterloader. The brachytherapy-based intraoperative approach is of particular interest because it is best suited to maximize the therapeutic ratio and deliver doses higher than any other treatment modality. For instance: • The treatment is performed at the time of surgery, when the target area (the tumor bed) is exposed and the applicator can be placed directly over the target. • Organs at risk may be retracted and shielded as necessary. • This technique is free of any accessibility limitations: the applicator can be used in virtually any anatomic location. This is most important in the treatment of colorectal malignancies where the tumor bed is often inaccessible to the cones of a linac-based system (Harrison et al. 1998).
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• The flexible applicator used for HDR-IORT easily conforms to the target area, ensuring uniform dose delivery throughout the target (avoiding cold spots and hot spots often encountered when using the electron beam approach as a result of angle of beam incidence and field matching). Finally, beyond the clinical and technical considerations, IORT offers the convenience and cost effectiveness of accelerated-course radiotherapy used more recently for the treatment of breast cancer. In this chapter it is assumed that the reader is familiar with the general requirements for HDR brachytherapy. What follows is a summary of the special needs and peculiarities of brachytherapy based IORT.
Hardware Afterloading Devices The three HDR afterloading machines most commonly used are manufactured by Nucletron, BV (The Netherlands) and by Varian Medical Systems, Inc. (GammaMed® and VariSource™), and are all suitable for use in IORT.
Applicators Several applicators are mentioned in the literature (Nag et al. 1994), however, some may not have Food and Drug Administration (FDA) 510K approval. The user is urged to contact the vendor of the afterloader for further information. Described here is the Memorial Sloan-Kettering Cancer Center (MSKCC) experience with the Harrison-Anderson-Mick (HAM) applicator (Mick Radio-Nuclear Instruments, Mount Vernon, NY), as used with a GammaMed 12i afterloader. Versions of this applicator are also available for the VariSource and Nucletron afterloaders. The standard HAM applicator (Figure 1) consists of standard-length (130 cm) catheters embedded at 1 cm spacing in 8 mm thick silastic rubber: with 5 mm from the center of catheters to the front of the appli-
Figure 1. Picture of a standard HAM applicator (courtesy of Mick Radio Nuclear Instruments, Inc., Mount Vernon, NY).
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cator and 3 mm to the back. The asymmetry, reducing the overall thickness of the applicator, was chosen for added flexibility. The number of catheters varies from 2 to 24, and the length of the silastic is 22 cm, resulting in a treatment area up to 23 cm in width and 20 cm in length. The prescription point is usually 0.5 cm from the surface of the applicator (or 1 cm from the plane of the catheters). A modified HAM applicator for use in breast-IORT is shown in Figure 2. Here, the silastic rubber is molded symmetrically with at least 1 cm from the surface to any of the catheters. This applicator is fitted with a tungsten shield, to protect the skin at the incision. Prescription using this applicator is usually 1 cm from the surface of the applicator (2 cm from the source plane).
Cost of IORT Brachytherapy On the average, the applicator cost is about $100.00/channel. Source exchange and periodic maintenance (assuming quarterly source exchanges) are approximately $50,000.00/year. With an average treatment width of five channels, and one treatment per week, the cost per treatment, excluding initial investments, is approximately $1,500.00.
Treatment Planning Dose Specification Being in an operating room (OR), this is arguably one of a few instances in which a medical physicist will be working upon oral instructions rather than on written directive (however, standard OR procedures should be followed: details of the prescription should be clearly repeated by the physicist and confirmed by the physician). An IORT prescription typically specifies the width (or number of channels), the length (or number of positions if the source stepping distance is known), the dose required, and the prescription point. Typically, a homogenous dose distribution is desired throughout the treatment area, with doses in the range of 15 to 20 Gy. Lower doses (10 to 15 Gy) are usually used for pediatric cases, and when IORT is used as a boost in conjunction with EBRT treatment. Occasionally, nonrectangular target areas and critical organs may be specified. The orientation of the applicator must be established to implement such dose prescriptions correctly.
Figure 2. The breast applicator modeled after the HAM applicator (Figure 1) is 2 cm thick and is used in tandem with a matching tungsten shield to protect the skin at the incision area.
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Treatment-Planning Approaches Most IORT treatment plans may assume fixed applicator geometry and be generated from pre-calculated plans or template-based plans. These are essentially fixed treatment geometries that are stored in, or automatically generated by, the treatment-planning system. These treatment-planning schemes are attractive mainly for their simplicity, reliability, and fast output (Anderson et al. 1997). Occasionally, custom plans are needed to accommodate regions of dose escalation or dose sparing. Figure 3 is an illustration of a plan with a dose escalation region in the middle of the treated area. Standard IORT plans are symmetric by nature. However, this often is not the case for custom treatments. In those cases, one should take care to note the orientation of the applicator with respect to the patient (e.g., channel 1-Med; applicator tip-Post); channel numbering tags are mounted on the catheters for this purpose (see Figure 1). While the most efficient of dose sparing methods employ a combination of retraction and shielding of the organs at risk, in some cases anatomic/geometric constraints require that the dose be tailored by means of treatment plan optimization. Breast IORT, for example, aims at delivering 20 Gy to a surface 1 cm away from the applicator while sparing the skin (ideally the skin dose should not exceed 10 Gy). In this type of geometry the skin is at the applicator surface, and very close to the treatment volume. It is usually possible to achieve a skin dose of 15 Gy by means of dose optimization, and further reduce the skin dose to 10 Gy using shielding (Figure 2). It is assumed that most facilities will be using the treatment planning system supplied with the afterloader, and that the physicist involved is familiar with that planning system. The details of the planning are left to be worked out by the physicist for the planning system available to him/her in a manner compatible with the clinical needs of the respective hospital.
Computer-Assisted Plan Verification Plan verification, unlike treatment planning, is not implemented in commercial software packages. Several schemes have been suggested (Ezzell 1994; Venselaer, Bierhuizen, and Klop 1995; Rogus, Smith, and Kubo 1998; Miller, Davis, and Horton 1996; Saw et al. 1998), most of which address only singlecatheter treatments. HDR brachytherapy plans are complex in nature; they involve many source-stopping positions with widely varying dwell times, and the time allotted for planning and verification is limited.
Figure 3. An example of a custom plan generated using Abacus (Varian Medical Systems, Charlottesville, VA) treating the center of the target to 17.5 Gy and the periphery of the target to 12.5 Gy.
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Consequently, it is impractical to manually review the plan for errors. To address this problem, various implementations of computerized quality assurance (QA) using spreadsheet calculations are also used. Typically, the physicist will type dwell times into a spreadsheet form and calculate the dose at one or more points of interest. However, manual data entry slows the QA process, and makes this method susceptible to errors, especially if used for large implants. As a means for independent plan verification, it is suggested that a computer program be used to assist in this task (Cohen, Amols, and Zaider 2000). Here, we describe in detail the program used for HDR plan verification at MSKCC. For completeness, we shall broaden the scope of this discussion to any procedure (whether or not invasive, or under anesthesia) for which a patient is on the treatment table, waiting for the plan to be completed and the treatment to commence. This definition includes (at our hospital) cylinders for vaginal treatment, ring and tandem for cervical cancer, and endobronchial treatments of the lung. The QA of HDR treatment planning in general has been discussed extensively in the report of AAPM TG-59 (Kubo et al. 1998). This program is an efficient tool assisting the physicist in implementing AAPM recommendations. Once a treatment plan is completed, the treatment-planning system generates a data file, which contains all the parameters necessary for treatment. This information, which is transferred to the afterloader console for treatment delivery, serves as the main data source for the independent verification program. Use of this file for secondary-dose calculations offers several obvious advantages. It provides an orderly way to: (a) read treatment parameters into the dose verification program, (b) confirm that the treatment data file corresponds to the actual treatment plan, (c) ensure that the data file is intact, and (d) use actual treatment parameters to both verify treatment geometry and perform a secondary dose calculation. The secondary verification consists of two steps: (a) determination/validation of treatment geometry, and (b) dose calculation based on that geometry. Once these two steps are completed the independent reviewer can verify that the treatment file conforms to the desired treatment parameters. User input is minimal. Determination of Implant Geometry The main task in the determination of implant geometry is the correct reconstruction of dwell positions (DPs). Each DP is defined by a nominal dwell time and a set of coordinates in the patient’s coordinate space (X, Y, Z). For fixed geometry implants, DP coordinates are implicitly defined by the configuration of active channels. In general, for other implant geometries such as template-based implants, DP coordinates may be approximated using explicit user input. However, this is beyond the scope of this discussion. Applicator geometry for all procedures mentioned above is assumed fixed and can be encoded in the software. The program detects automatically the type of applicator being used, virtually eliminating user input. Specifically, the number of active channels and channel index indicate which applicator is used. Because for most applicators the location of the prescription point is also fixed, the complete implant geometry can be automatically reconstructed. In the case of IORT, the active treatment area is defined by the length L (Y-axis) and width W (X-axis), with the origin at the center of the implant (Figure 1). Assuming a flat applicator (Z = 0 for all DPs) with N channels and M positions per channel, the coordinates of DP(n,m) are given by:
N −1
(X ,Y ) n. m =
2
− (n − 1) × ChSp,
M −1 − (m −1) × Stepn − Shiftn 2
(1.1)
where “Shift” is the distance the first dwell position is shifted from the tip of the applicator, “Step” is the distance between dwell point positions, and “ChSp” is the spacing between channels (which in the case of the HAM applicator is fixed and equals 1 cm; see Figure 1). In the case of a single channel application, where N=1, Xn,m reduces to X1,m = 0. It should be noted that in the specific case of the HAM applicator, each “atlas” treatment plan uses a fixed length (i.e., M is constant) for all channels. This, however, need
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not be the case (see description of custom plans above), because the program reads all 40 dwell times for each of the 24 channels, whether or not they are used. The detection of the applicator type and the calculation of the dose at the reference point(s) are performed as follows: 1.Single channel applications: (channel no. < = 19) indicates a vaginal applicator; (channel no. > = 20) indicates a bronchial/esophageal applicator. [The GammaMed afterloader automatically tests the length of channels 1 to 19 (of 24 channels) to ensure that the first dwell position corresponds with the end of the catheter/tube. Unlike vaginal applicators, endoesophageal and endobronchial applicators are subject to distortion and variations in length, and are typically attached to channel no. 24 of the afterloader.] At our institution, the dose reference point for a vaginal procedure is set at Pref = (0.0, 0.0, d/2+0.5) cm, where d is the cylinder diameter. Because d is variable, the distance of Pref from the source channel is entered manually. For bronchial or esophageal procedures, Pref = (0.0, 0.0, 1.0) cm. 2.Two active channels indicate GYN ring and tandem applicator. Here, the user is required to enter the length of the tandem. With the tandem orthogonal to the plane of the ring, and the channel in the ring section of the applicator also fixed, the applicator geometry is defined. In some cases (applicable only when the tandem is made of two sections), the angle of the sleeve may not be perpendicular to the plane of the ring; then user input is required. The reference points defined as points Aright and Aleft are fixed relative to the applicator hardware. The outer diameter of the ring (i.e., of the caps) is also entered, to calculate the dose to the vaginal mucosa. Uterine and cervix points1 are also fixed with respect to the applicator and are automatically calculated. 3.Three or more active channels indicate a HAM applicator. The assumption made here is that the applicator is flat. We found this to be a good approximation for most procedures. For HAM-IORT, we set the dose reference point at Pref = (0.0, 0.0, 1.0) cm (1.0 cm from the source plane, or 0.5 cm from the surface of the applicator) and for breast IORT at Pref = (0.0, 0.0, 2.0) cm. In the case of a convex curved application (e.g., rectal IORT) the dose calculation will predict an overdose of 5% to 10%. While this is still within regulatory requirements, our treatment planning “Atlas” and dose verification program can both accommodate the curved geometry. Dose Calculation Dose-to-point calculations consist of summation of dose contributions from all dwell positions, using a two-dimensional dose table, F(r,σ), for the 192Ir source, and the fifth-order Meisberger polynomial, M(r) (Meisberger, Keller, and Shalek 1986). Both the dose table and polynomial are hard coded in the program, making it self contained. Dose calculations are typically within 1% to 2% of the planned dose. By the end of plan verification, the total treatment time is also known. The total treatment time (including dummy tests) should be communicated to the anesthesiologist, to ensure proper medication of the patient. Also at this time, surgeons, the radiation oncologist, and the OR staff are ready to start treatment. The last step in the plan verification process is (oral) approval of the plan by the physician. It is important to “pause” and review the plan with the physician: this is the last chance to catch any miscommunications or other errors in an otherwise busy and time-pressured environment.
Classical Dosimetric Systems To complement the plan verification discussed above, the classical Manchester system may be used. The original Manchester tables were designed to give the total amount of radium (mg) times application time (h) required to deliver 1000 R for various implants, given the implant’s volume or area, and elongation (Merredith 1967). This quantity, M, differs only by a constant from total reference air kerma (TRAK) 1
The cervix and uterine points are defined in a similar manner as point A: the cervix is 1 cm superior to the flange and 1 cm lateral to the tandem, and the uterine points are 1 cm inferior to the tip of the tandem and 2 cm lateral to the tandem.
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expressed in gray (ICRU 1997). Because mg·h is still most commonly used, M is expressed as the total mg·h that produce 10 Gy in water. Using the ratio of exposure rate constants for radium and iridium and adjusting for units, the total dwell time for a nominal source activity of 370 GBq, expressed in s/Gy, is interpolated from the Manchester table, M (expressed in mg·h), as follows:
0.5 2 h 2 3/ 4 T = 0.0659 M A , 0.5 exp(0.05[E − 1] ) , h 0.5
(1.2)
where h is the distance (cm) from the plane of source to the treatment plane, A is the treatment area (cm2), and E is the elongation of A. The original Manchester tables were written for a treatment distance of 0.5 cm from the plane of the source. The equation above reflects the use of the table at other treatment distances as well. Predictions of the total nominal dwell time using this method are usually within 5% of the planned time.
QA Considerations In general, QA procedures should follow standard recommendations (Nath et al. 1997). However, it is imperative to remember that IORT is a single-fraction treatment. Proper QA of the hardware, software, and each treatment plan is essential for a successful IORT program. Special attention should be given to applicator QA: • The applicator is a single-use device that requires in-house sterilization. Each applicator should undergo QA prior to sterilization. Specifically, the QA procedure should include checks of catheter integrity and lengths, catheter labeling, and overall applicator integrity. If source transfer tubes are used, the QA should include them and they should be sterilized as well. • It is equally important to examine the applicator after the procedure is complete and to establish its integrity; if lead shields or source transfer guides were used, ensure that all are intact and accounted for. • Initial QA at the time of the IORT program is being set up should be performed to ensure the applicator and source transfer tubes are not damaged during the sterilization process. The operation room is a high-pressure environment. It is recommended that standardized forms and treatment protocols be used whenever possible. An example of a treatment setup form is provided in Figure 4.
Facility Design and Shielding Considerations The requirement of a shielded operating room is for the most part identical to those of a standard HDR treatment suite. A few design considerations are noteworthy: • Remote monitoring of the patient, usually placed at the treatment console, should include a monitoring station for anesthesia staff. • Essential OR staff and the radiation oncologist must remain “scrubbed” (i.e., they may not break sterility) during the procedure in case an emergency intervention is required. Room should be provided for them to wait throughout the procedure, preferably near the entrance to the OR.
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Figure 4. A template for an IORT worksheet. The redundancy of numerical entry (in the header of the form) and graphical representation helps ensure the oral directives are followed correctly.
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• In the event of a radiation emergency, the patient may not leave the OR. Instead, the facility should include a shielded enclosure into which the afterloader and source may be retracted and isolated, to enable completion of surgery. • In general, operating rooms tend to become crowded very easily, occasionally hampering access to the patient. To ensure easy setup of the remote afterloader, ensure all cables are long enough to reach all sides of the patient/room.
Personnel Requirements Radiation oncology staff required to be present for each procedure include: a physician, a therapist, and two physicists (one to plan and deliver the treatment, and one for independent verification). Surgery often starts early morning and in some cases may continue until late at night. When the patient will be ready for IORT is not always predictable, which may place yet an additional strain on busy physics staff. Good planning may often help, but do realize that long, waiting hours will sometimes be unavoidable.
Emergency Procedures Emergency procedures during IORT follow the same guidelines as other HDR emergency procedures but for one crucial exception: the patient may not leave the operating room. As discussed above, the source and afterloader should be isolated from the patient, preferably into a shielded housing in the corner of the operating room.
References Anderson, L. L., M. R. Hoffman, P. J. Harrington, and G. Starkschall. (1997). “Atlas generation for intraoperative high dose rate brachytherapy.” J Brachyther Int 13:333–340. Cohen, G. N., H. I. Amols, and M. Zaider. (2000). “An independednt dose-to-point calculation program for the verification of high-dose-rate brachytherapy treatment planning.” Int J Radiat Oncol Biol Phys 48(4):1251–1258. Ezzell, G. A. (1994). “Quality assurance of treatment plans for optimized high dose rate brachytherapy—planar implants.” Med Phys 21:659–661. Harrison, L.B., B. D. Minsky, W. E. Enker, B. Mychalczak, J. Guillem, P. B. Paty, L. L. Anderson, C. White, and A. M. Cohen. (1998). “High dose rate intraoperative radiation therapy (HDR-IORT) as part of the management strategy for locally advanced primary and recurrent rectal cancer.” Int J Radiat Oncol Biol Phys 42(2):325–330. International Commission on Radiation Units and Measurements (ICRU). Report No. 58. Dose and Volume Specification for Reporting Interstitial Therapy. Bethesda, MD: ICRU, 1997. Kubo, H. D., G. P. Glasgow, T. D. Pethel, B. R. Thomadsen, and J. F. Williamson. (1998). “High dose-rate brachytherapy treatment delivery: Report of the AAPM Radiation Therapy Committee Task Group No. 59.” Med Phys 25(4):375–403. Also available as AAPM Report No. 61. Meisberger, L., R. Keller, and R. Shalek. (1986). “The effective attenuation in water of the gamma rays of gold 198, iridium 192, cesium 137, radium 226 and cobalt 60.” Radiology 90:953–957. Meredith, W. J., (ed). Radium Dosage. The Manchester System. Edinburgh: Livingston, 1967. Miller, A. V., M. G. Davis, and J. L. Horton. (1996). “A method for verifying treatment times for simple high-doserate endobronchial brachytherapy procedures.” Med Phys 23:1903–1908. Nag, S., P. Lukas, D. S. Thomas, and L. B. Harrison. “Intraoperative High Dose Rate Remote Brachytherapy” in High Dose Rate Brachytherapy: A Textbook. S. Nag. Armonk NY: Futura Publishing Company, 1994. Nath, R., L. L. Anderson, J. A. Meli, A. J. Olch, J. A. Stitt, and J. F. Williamson. (1997). “Code of practice for brachytherapy physics: Report of the AAPM Radiation Therapy Committee Task Group No. 56.” Med Phys 24(10):1557–1598. Also available as AAPM Report No. 59.
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Rogus, R. D., M. J. Smith, and H. D. Kubo. (1998). “An equation to QA the total treatment time for single-catheter HDR brachytherapy.” Int J Radiat Oncol Biol Phys 40:245–248. Saw, C. B., L. J. Korb, B. Darnell, K. V. Krishna, and D. Ulewicz. (1998). “Independent technique of verifying highdose rate (HDR) brachytherapy treatment plans.” Int J Radiat Oncol Biol Phys 40:747–750. Venselaar, J. L., H. W. Bierhuizen, and R. Klop. (1995). “A method to check treatment time calculations in Ir-192 high-dose-rate volume implants.” Med Phys 22:1499–1500.
Chapter 28
Introduction to Prostate Brachytherapy Wayne M. Butler, Ph.D. and Gregory S. Merrick, M.D. Schiffler Cancer Center Wheeling Hospital Wheeling, West Virginia Prevalence and Mortality of Prostate Cancer . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 522 Frequency of Occurrence and Mortality . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 522 Trends Over Time . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 522 Effect of PSA Screening . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 523 Combined PSA and Digital Rectal Exam (DRE) Screening . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 524 Clinical Characteristics of Prostate Cancer . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 524 PSA . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 524 Gleason Score . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 525 Clinical Stage . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 525 Imaging Modalities . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 525 Other Prognostic Factors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 527 Patient Selection and Classification . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 527 Partin Tables . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 527 Risk Group Classification . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 527 Patient Quality of Life Factors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 528 Preimplant Urinary Function . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 528 Erectile Function . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 528 Rectal Function . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 528 Other Criteria Sometimes Used to Select or Reject Patients . . . . . . . . . . . . . . . . . . . . . . . . . . . . 528 Defining Clinical Success . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 532 Randomized Trials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 532 Survival . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 532 Biochemical Failure as a Survival Marker . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 533 When To Declare a PSA Failure? . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 533 Therapy Modalities . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 535 Radical Prostatectomy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 535 3-D Conformal External-Beam Radiation Therapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 535 Interstitial Brachytherapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 536 Combined Modality Treatments . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 536 Hormonal Therapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 536 Biochemical, Recurrence-Free Survival . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 537 Kaplan-Meier Curves . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 537 Comparative Single Institution Survival Results . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 537 Inter-institution and Inter-modality Comparisons . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 538 Actuarial Survival for Low-Risk Patients . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 538 Actuarial Survival for Intermediate-Risk Patients . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 540 Actuarial Survival for High-Risk Patients . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 540 Treatment-related Morbidity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 543 Urinary Morbidity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 544 The Role of Alpha-blockers on IPSS . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 544 Urethral Dosimetry . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 546 Rectal Dosimetry and Morbidity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 546 Erectile Dysfunction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 548 Mechanism of Brachytherapy-Induced Erectile Dysfunction . . . . . . . . . . . . . . . . . . . . . . . . . . . . 548 Radiation Dose to the Prostate Gland . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 548
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Prevalence and Mortality of Prostate Cancer Frequency of Occurrence and Mortality Prostate cancer is one of very few cancers for which there is still a debate not just on how to treat but whether there should be any treatment at all. Today, in a well-screened population, few men diagnosed with prostate cancer should die from that disease. Current screening practice has increased the fraction of prostate cancer which is clinically localized at diagnosis to 90%, which is much higher than rates of locally confined disease found for other cancers with effective population screens—breast, cervix, and colon/rectum at 63%, 55%, and 39%, respectively (Carroll 2005). Nevertheless, the American Cancer Society, using models from the National Cancer Institute, estimates that in 2005 over 30,000 men in the United States will die from prostate cancer and over 230,000 new cases will be diagnosed (Jemal et al., 2005). Among men, prostate cancer accounts for one-third of all new cancers diagnosed—far exceeding lung and bronchus at 13%—but the death rate from prostate cancer, at 10% of all male cancer deaths, is second to lung and bronchus at 31% of all male cancer deaths. In Sweden, active cancer screening is not advocated, but citizens have essentially free access to necessary health care resources if they have symptoms or if they request a test such as prostate-specific antigen (PSA) or a digital rectal examination (DRE). In a prospective analysis of 8887 Swedish men diagnosed with prostate cancer, the disease specific survival rate at 15 years was 44% (Aus et al., 2005). In this largely unscreened population, the median age at diagnosis was 75 years and the median age at death was 80 years. These data represent a worst-case scenario for well-developed countries with high life expectancy. In the Swedish study, treatment with curative intent was given to only 11% of the population, and 37% did not even receive palliative treatment. In that portion of the population with localized cancer for which four fifths were treated curatively, the 10-year and 15-year disease specific survival rates were 93% and 80%, respectively.
Trends Over Time The median age at diagnosis in the United States is about 66 years, but has been decreasing as PSA screening is applied to ever-younger age cohorts. When PSA screening was introduced in the late 1980s, incidence rates increased dramatically, but the rates have subsequently decreased or leveled off for age groups over 60 years (American Cancer Society 2005) and overall the rate is only slightly greater than before the PSA era. Prostate cancer remains extremely rare in men under the age of 40 years. Because of the relatively advanced age at diagnosis and the typically slow progression of disease, prostate cancer patients have an average years of life lost of 6.1 years—the lowest years of life lost of 17 common cancers (Burnet et al., 2005). Brain cancer patients suffer the greatest loss at over 20 years. Cancer statistics show a significant trend for improved 5-year relative prostate cancer survival rates, which have increased from 67% in 1974–1976, to 75% in 1983–1985, to 99% in 1995–2000. Figure 1 plots prostate cancer death rates per 100,000 men. Cancer death rates rose in the late 1980s and peaked in 1992 when the number of deaths exceeded 40,000, and the rate has been declining ever since. Using analytical methods responsive to rapidly changing patterns of care, a continual improvement in survival
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Figure 1. Prostate cancer mortality rate per 100,000 male population. The results are age-adjusted to the distribution of ages in the 2000 census. (Adapted from American Cancer Society, 2005.)
has been documented (Brenner and Arndt 2005). For men with well- or moderately differentiated tumors and a localized or regional stage, their calculated 5-year and 10-year relative survival rates exceeded 100% indicating no excess mortality in that cohort of men, which compromised two thirds of the Surveillance, Epidemiology, and End Results (SEER) study population.
Effect of PSA Screening Although most of us who treat prostate cancer would be happy to accept credit for such increasingly favorable survival results, we should be aware of competing explanations. Screening contributes to diagnosis of disease at earlier stages (Harris and Lohr 2002). Certainly, some men are diagnosed with disease that may never have become clinically significant, but this effect is likely to be small because current incidence rates mimic that of the prescreening era. Earlier diagnosis also means that patients will inevitably live with their disease longer. PSA screening may also selectively enhance the detection of prognostically favorable tumors. Indeed, there has been a reduction in the incidence at diagnosis of metastatic disease, positive lymph nodes, and seminal vesicle involvement, but there has been no significant change in microscopic measures of aggressiveness. The value of PSA screening has been called into question by one of the founders of the test (Stamey et al., 2004) because of a purported loss over time of its positive predictive value. Prostate cancer readily meets two of the criteria for a disease amenable to mass population screening, namely, a high prevalence within the population and a high mortality rate if left untreated (Wilson and Crawford 2004). Therefore, the utility of a prostate cancer-screening test depends on its cost effectiveness. Cost effectiveness requires that the test have reasonably sensitivity (high true positive detection rate) and reasonably specificity (high true negative detection rate). The cost of the test itself should be relatively low, and the societal costs associated with interventions implemented after a positive test (such as morbidity due to biopsy) should be proportionate to years of life saved. For a PSA threshold of 4.0 ng/mL, the sensitivity of PSA is 75% and the specificity is 60% (Wilson and Crawford 2004). The sensitivity of any quantitative diagnostic test may be improved by lowering the threshold, but the specificity will always diminish, meaning that more men would be subjected to unnec-
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essary biopsies. As part of a cancer prevention trial, biopsies were performed on 2950 men who had completed 7 years in the placebo arm and whose PSA had never exceeded 4.0 ng/mL (Thompson et al., 2003). The biopsies were pathologically positive for prostate cancer in 15.2% of the men, and 15% of the cancers detected had a Gleason score of 7 or greater, which is almost certainly an indicator of clinically significant disease. Lowering the PSA threshold that prompts a prostate biopsy to 2.5 ng/mL will detect cancers with more favorable characteristics. In an analysis of PSA and cancer detection rate in over 36,316 biopsies, the detection rate in men with PSAs in the range 2.5 to 4.0 ng/mL was 27.5% while the rate for men in the 4.0 to 10.0 ng/mL range was similar at 30.1% (Gilbert et al., 2005). The effect of changing the threshold from 4.0 to 2.5 ng/mL improved the sensitivity from 87% to 94% while the specificity declined from 20% to 12%. One way to improve the sensitivity without greatly affecting the specificity of the PSA test has been to create age-specific thresholds. Other effective enhancements include determining the PSA velocity in ng/mL/yr, measuring PSA density in ng/mL/cm3 of prostate, measuring free PSA, complex PSA, and their ratios to total PSA, and measuring precursor forms of PSA (Han, Gann, and Catalona 2004). The frequency of screening may also be adjusted for patient risk factors, most notably, the PSA velocity or PSA doubling time. In a large screening population study, men who remained free of cancer over a 10-year period had an average PSA velocity of 0.03 ng/mL/yr (Berger et al., 2005). Also of significance was each total PSA in the previous annual exam. Men with an antecedent PSA of 4 ng/mL and referred to their primary care physician for discretionary follow-up. Men testing positive for either DRE or PSA were 14% of the total, but there was little overlap between the tests with 7.5% positive for DRE and 7.9% above the PSA threshold. Ultimately, 32% of the men with suspicious screening results underwent prostatic biopsy within 1 year of the screen, and 37% of those had biopsy-proven cancer. An abnormal PSA was nearly twice as likely as a positive DRE to result in biopsy, and the yield of positive biopsies was less in DRE positive men (34%) than in PSA suspicious men (44%). Nevertheless, a DRE is considered necessary to detect prostate cancers, which produce little or no PSA.
Clinical Characteristics of Prostate Cancer PSA PSA testing unequivocally revolutionized the diagnosis and treatment of prostate cancer. Prior to 1988 when that test became widely available, diagnosis was usually based on an abnormal digital rectal exam or an incidental pathological finding after some other procedure. Urologists and radiation oncologists had been smugly confident in the efficacy of either surgery or radiation therapy because most prostate cancers
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are inherently indolent and any primary therapy of localized disease will delay the time to progression to frank symptoms such as painful bone metastases. Among the 18% of men in one large series who developed bone metastases, the median time from surgery to that end point was 9 years (Pound et al., 1999). PSA demolished false perceptions of treatment efficacy because the test may also be used as a marker of success or a harbinger of symptomatic failure. A rising or non-nadiring PSA is now referred to as a biochemical failure, and such failures were found to be rife in older surgical and radiation therapy series. The dismal situation prompted the famous conundrum posed by the urologist Willet Whitmore (1990), “Is cure possible in those for whom it is necessary, and is cure necessary in those for whom it is possible?” Because disease detected by PSA screening and pathologically confirmed is very rarely classified as insignificant, the answer to the latter half of the question is yes. The answer to the former part is also yes, but the disease must be loco/regionally confined and the choice and quality of the treatment must be appropriate for success to occur.
Gleason Score Gleason score is one of the most important predictors of biologic behavior, extracapsular extension, and outcome after treatment, and often plays an important role in determining patient management. From microscopic examination of stained prostate tissue, a pathologist assigns a Gleason grade to a specimen based on the glandular architecture from well differentiated (grade 1) to poorly differentiated (grade 5). The grades of the two most prevalent architectural patterns (primary and secondary) are added together to produce the composite Gleason score, e.g., grade 3 + grade 4 = GS 7. Although Gleason scores on prostate needle biopsies often differ by ±1 from those assigned after radical prostatectomy, the differences are most common in lower grades (Bostwick 1994) for which there is an emerging consensus for treatment approach. Of greater concern is the spotty concordance between Gleason scores assigned by community based pathologists and those from prostate cancer specialists (Merrick et al., 1998). Community-based pathologists tend to undergrade Gleason scores relative to their more experienced counterparts. The distribution of Gleason scores within an institution should be similar to those reported in the literature: 10%) of patients diagnosed with Gleason score half of one lobe of the prostate Involves both lobes of the prostate Extracapsular extension of one or both lobes Seminal vesicle invasion Invades bladder neck, muscle, pelvic wall or other
becoming increasingly rare in modern screened populations. In one study by Lee et al. (2000), the rate of positive bone scans was only 1% in men with a PSA 100% mPD) should be avoided. Leibovich and colleagues (2000) have measured the mean distance from the urethra to the nearest foci of cancer to be 3 mm and that 17% of all prostate cancer abuts the urethra. Although it is conceivable that certain segments of the urethra may be more sensitive to radiationinduced morbidity, detailed urethral dosimetry did not substantially improve the ability to predict urinary morbidity (Allen et al., 2005). Radiation doses of 100% to 140% mPD were well tolerated by all segments of the prostatic urethra. It is also possible that the relative insensitivity of segmental urethral doses may have been a result of prophylactic and prolonged alpha-blocker usage. Urethral strictures occur in 4% to 12% of patients and are directly related to over implantation of the periapical region and the use of supplemental external beam radiation therapy (Merrick et al., 2002b, 2005d). Typically, strictures involve the bulbomembranous urethra and are easily managed with dilatation or optical internal urethrotomy. Day 0 CT-based dosimetry has demonstrated that radiation doses to the bulbomebranous urethra are significantly greater in patients with strictures than in those without. With careful attention to implant technique including extensive use of sagittal ultrasonography, it is possible to implant the apex with a 5-mm margin without “dragging” seeds into the bulbomembranous urethral region.
Rectal Dosimetry and Morbidity Rectal complications primarily consist of mild self-limited proctitis (incidence 4% to 12%), and they usually resolve spontaneously (Gelblum and Potters 2000). Bowel function assessment by patient-administered questionnaires has documented that long-term bowel function is minimally affected by brachytherapy (Talcott et al., 2001). Using the RFAS questionnaire (Table 5), Merrick et al. (2003c) found that bowel function worsened slightly and remained elevated. The mean preimplant RFAS was 2.5 and the post implant RFAS was 4 out of the maximum score of 27 (Figure 13). On a retrospective summary question that asked patients to compare their current bowel function to what they recall from before the implant, a large majority perceived no change and those who felt their function had improved or worsened were about equal in number (Figure 14). Rectal dose is important in more severe complications, but it is difficult to compare rectal dosimetry between institutions because the timing of postimplant CT and how the rectum or rectal wall was evaluated varies widely. Waterman and Dicker (1999) reported that the minimum dose that encompassed 10% of the surface area of the rectum (rectal area D10) increased by 68% from day 0 through day 30. Using day 30 CT-based dosimetry, Snyder and colleagues (2001) reported grade II proctitis to be rectal volume dependent for any given dose and that no cases developed more than 36 months following implantation. They documented an 8% rate of grade II proctitis when less than 1.8 cm3 of rectum was exposed to 160 Gy following 125I monotherapy, while the risk increased to 25% when more than 1.8% of the rectum was
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Figure 13. Rectal function assessment score (RFAS) changes over time for 189 patients surveyed in 1999 (open circles) and again in 2002 (filled circles). The dark solid line indicates an insignificant improvement during the 3-year interval (Merrick et al. 2003c).
Figure 14. Patient perception of changes in bowel function relative to preimplant status. Questionnaires were administered to 189 patients in 1999 and again in 2002 (Merrick et al. 2003c).
so exposed. In a prospective randomized trial, the minimum dose received by 5% of the rectum best correlated with brachytherapy-related rectal morbidity (Merrick et al., 2003b). Although it has been assumed that perirectal sources increase rectal morbidity, a limited number of errant perirectal sources did not increase the risk of rectal bleeding provided the overall rectal wall doses were within acceptable values (Mueller et al., 2004). Following brachytherapy, rectal ulceration and fistula formation have occasionally been reported. Although dose is associated with rectal bleeding, Howard et al. (2001) found no correlation between rectal dose and the development of a fistula. They concluded that severe complications might occur in
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an unpredictable manner unrelated to known clinical, treatment, or dosimetric parameters. A reasonable conjecture is that a patient experiencing rectal bleeding sees his family physician, and that visit precipitates invasive rectal procedures such as biopsy, cauterization, laser coagulation, etc. These procedures exacerbate the problem because of the limited healing capacity of highly irradiated tissue. Although no studies have correlated constipation with rectal toxicity, constipation significantly increases the radiation dose to the rectum (Merrick et al., 2000b). For patients prone to constipation, postimplant attention to bowel habits and the use of stool softeners and laxatives for two half-lives of the implanted source will minimize rectal distention and decrease the dose to the anterior rectal wall. Intraoperatively, careful attention to implant technique and ultrasound anatomy can reduce the dose to the anterior rectal wall and minimize bowel dysfunction. Extensive use of both transverse and sagittal ultrasonography to confirm needle placement and the use of multiple ultrasound frequencies helps ensure proper seed placement. Higher transducer frequencies result in clearer definition of anatomy closer to the probe. Posterior row needle placements executed with the 7.5 MHz frequency help ensure that needles are placed within the posterior prostate capsule and not in the rectal wall.
Erectile Dysfunction Erectile dysfunction (ED) is a common sequela of all potentially curative local treatments. The reported rates of post-implant ED vary widely reflecting substantial differences in follow-up, patient selection, implant technique, and the mode of data collection. Early brachytherapy studies reported exceptionally high rates of potency preservation (Wallner, Roy, and Harrison 1996). However, patient-administered quality of life questionnaires differ markedly from data collected by physician interview, especially on sensitive sexual issues (Litwin et al., 1998). Following all treatment approaches, potency preservation rates are significantly lower when patients fill out a form rather than answering oral questions from their treating physician. Although the perception exists that preservation of erectile function is more likely following brachytherapy than other local treatments, the incidence of brachytherapy-induced ED is substantially greater than initially reported (Merrick et al., 2002c, 2005e). Subgroup analyses documented ED in 6% to 90% of patients undergoing brachytherapy with or without supplemental external-beam radiation therapy and/or androgen deprivation therapy. Mechanism of Brachytherapy-Induced Erectile Dysfunction Radiation-induced ED likely represents a multifactorial process including neurogenic compromise, vascular insufficiency, local trauma, and psychogenic causes with microvascular damage representing the most dominant factor (Zelefsky et al., 1999). There are several candidates for structures at risk. Emerging evidence strongly suggests that brachytherapy-related ED is technique-related and may be minimized with careful attention to radioactive source placement. Radiation Dose to the Prostate Gland Conflicting results have been reported regarding the relationship between radiation dose to the prostate gland and the development of brachytherapy-related ED. In multivariate analysis, Stock, Kao, and Stone (2001) reported that a D90 > 160 Gy for 125I implants and >100 Gy for 103Pd implants predicted for brachytherapy-induced ED. However, the absolute differences in potency preservation were minimal— 58% versus 64%, p = 0.02. In contrast, in both retrospective and prospective evaluations, Merrick and colleagues were unable to discern a relationship between radiation dose to the prostate gland and brachytherapy-related ED (Merrick et al., 2002c, 2005e).
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Dose to the Neurovascular Bundles Following radical prostatectomy, ED has been correlated with surgical trauma to the neurovascular bundles. It is conceivable that excessive radiation doses to the neurovascular bundles (NVB) could also represent a potential mechanism of brachytherapy-induced ED (Zelefsky and Eid 1998). Radiation doses to the NVB were calculated using an idealized 2-dimensional geometric model derived from day 0 CTbased images (DiBiase et al, 2000). Using the same methodology in both prospective and retrospective analyses, Merrick et al., (2000c) found no relationship between radiation dose to the NVB and the development of brachytherapy-induced ED. A recent study from Wake Forest University confirmed these findings (Kiteley et al., 2002). Although initial reports have shown no correlation between ED and NVB dose, it is possible that with additional follow-up those doses will be found to contribute to brachytherapy-related ED. Dose to Penile Erectile Bodies The penile erectile bodies—the paired corpora cavernosa and the midline corpus spongiosum—are potential targets for structure-specific radiation-associated ED. Figure 15 is a cartoon of the anatomy. Although the site-specific structure in the proximal penis remains unclear, radiation doses to the corpora cavernosa may be more important than those to the penile bulb because the corpora cavernosa represents the true erectile tissue whereas the corpus spongiosum plays little role in the development or maintenance of erectile rigidity (Mulhall et al., 2002). Detailed reports illustrating the image-based anatomy of the proximal penis have been published by Wallner et al., 2002a). The penile bulb is best visualized on T2-weighted magnetic resonance images and appears as an oval-shaped, hyper-dense midline structure located approximately 10 to 15 mm inferior to the apex of the prostate gland. In both a laboratory rat model and in clinical studies, investigators at the University of California at San Francisco reported that the external-beam radiation dose to the bulb of the penis correlated directly with post-treatment ED (Fisch et al., 2001). A higher incidence of ED was reported with increasing radiation doses to the proximal penis. The authors hypothesized that external-beam radiation directly or indirectly damaged the vascular supply of the erectile tissue as well as the nerves that supply the cavernosal smooth muscles. Subsequent histologic examination of the proximal penile shaft specimens demonstrated that the numbers of nitric oxide synthase-containing nerve fibers per corpus cavernosum were significantly decreased in the irradiated groups.
Figure 15. Sagittal schematic of the prostate and nearby structures.
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Prostate brachytherapy-induced ED was found to be highly dependent on the radiation dose to the proximal penis (Merrick et al., 2001b, 2002d). Using day 0 CT-based dosimetry, maximum potency preservation was dependent on limiting the penile bulb D50 to ≤40% of prescription dose and the proximal crura D50 to ≤50% of the prescription dose (Merrick et al., 2005e). When stratified by these cutpoints, greater than 70% of patients with dose below the cutpoint maintained potency versus only 30% in patients who received higher doses (Figure 16). In contrast, a patient series from Kiteley et al., (2002) failed to establish a relationship between radiation dose to the proximal penis and the development of brachytherapy-related ED. Ultimately, the proximal penis may prove to be the most important site-specific structure for the development of radiation therapy-induced ED. In the case of brachytherapy, suboptimal placement (either due to poor planning or poor implantation) of periapical radiation sources results in excessive radiation doses to the proximal penis. Refinements in implant technique to avoid over-aggressive periapical implantation along with extensive use of sagittal ultrasonography during the implant procedure will decrease the radiation dose to the proximal penis and may improve potency preservation. Potency Preservation ED following definitive therapy increases with time after treatment, and series with longer follow-up have uniformly reported lower rates of potency preservation. Erectile function gradually declines with age, and older patients are less likely to maintain erectile function following any definitive procedure, including brachytherapy. In a recent evaluation of low-risk patients on a prospective randomized trial, Merrick and colleagues (2005e) found that 50% of patients maintained potency 4 years following prostate brachytherapy. Preimplant erectile function—measured by the IIEF (Table 4)—represents the strongest predictor of post-treatment erectile function (Merrick et al., 2005e). That study also discovered that men who experienced nocturnal erections prior to treatment were more likely to maintain potency. Supplemental external-beam radiation therapy and androgen deprivation therapy may be detrimental when combined with brachytherapy in their impact on potency preservation, but there have been conflicting results regarding their effect. Merrick et al. (2005e) found no significant difference between the potency preservation rates between men treated with or without 3DCRT, or with or without ADT, or with 125I or 103Pd. Their
Figure 16. Potency preservation stratified by a penile bulb D50 dose of 30% of the prescribed minimal peripheral dose (Merrick et al. 2005e).
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3DCRT and brachytherapy treatment planning were designed to limit the dose to the proximal penis. On the other hand, Potters and colleagues (2001) found that use of ADT was a strong predictor for ED and that men treated with 3DCRT and permanent implant were much less likely to retain potency. Fortunately, the majority of patients with brachytherapy-induced ED respond favorably to sildenafil citrate (Viagra) (Merrick et al., 1999) or the other PDE-5 inhibitors (type 5 phosphodiesterase) vardenafil (Levitra) and tadalafil (Cialis). A 6-year actuarial rate of potency preservation of 92% was reported when potent patients were analyzed together with erectile dysfunction patients who used sildenafil (Merrick et al., 2002c).
Conclusions There is a growing volume of evidence that permanent prostate brachytherapy produces very favorable and durable biochemical survival for patients with early-stage carcinoma and who have either low-, intermediate-, or high-risk features. There is also increasing evidence in support of the trend to rely less on other modalities—3DCRT and ADT—in combination with brachytherapy in favor of a monotherapy approach. Biochemical progression-free survival is now so high in low- and intermediate-risk patients that only a very large, highly powered randomized trial could prove any innovation or competing modality to be superior to current permanent seed brachytherapy. With feedback from post-implant evaluation, brachytherapy planning philosophies and operative techniques have been refined not only to further enhance disease-free survival but also to further reduce ED and urinary and rectal morbidity. Significant urinary and rectal morbidity is becoming increasingly rare, and most patients with brachytherapy-induced ED respond positively to sildenafil.
References Abuzallouf, S., I. Dayes, and H. Lukka. (2004). “Baseline staging of newly diagnosed prostate cancer: a summary of the literature.” J Urol 171:2122–2127. Allen, Z. A., G. S. Merrick, W. M. Butler, K. E. Wallner, B. Kurko, R. L. Anderson, B. C. Murray, R. W. Galbreath. (2005). “Detailed urethral dosimetry in the evaluation of prostate brachytherapy-related urinary morbidity.” Int J Radiat Oncol Biol Phys. In press. American Cancer Society. Cancer Facts and Figures 2005. Atlanta, GA: American Cancer Society, 2005. American Joint Committee on Cancer. AJCC Cancer Staging Manual, 6th Ed. New York: Springer-Verlag, 2002. American Society for Therapeutic Radiology and Oncology Consensus Panel. (1997). “Consensus statement: guidelines for PSA following radiation therapy.” Int J Radiat Oncol Biol Phys 37:1035–1041. Andriole, G. L., D. L. Levin, E. D. Crawford, E. P. Gelmann, P. F. Pinsky, D. Chia, B. S. Kramer, D. Reding, T. R. Church, R. L. Grubb, G. Izmirlian, L. R. Ragard, J. D. Clapp, P. C. Prorok, and J. K. Gohagan. (2005). “Prostate Cancer Screening in the Prostate, Lung, Colorectal and Ovarian (PLCO) Cancer Screening Trial: Findings from the initial screening round of a randomized trial.” J Natl Cancer Inst 97:433–438. Aström, L., D. Pedersen, C. Mercke, S. Holmäng, and K. A. Johansson. (2005). “Long-term outcome of high dose rate brachytherapy in radiotherapy of localised prostate cancer.” Radiother Oncol 74:157–161. Augustin, H., M. Graefen, J. Palisaar, J. Blonski, A. Erbersdobler, F. Daghofer, H. Huland, and P. G. Hammerer. (2003). “Prognostic significance of visible lesions on transrectal ultrasound in impalpable prostate cancers: implications for staging.” J Clin Oncol 21:2860–2868. Aus, G., D. Robinson, J. Rosell, G. Sandblom, and E. Varenhorst. (2005). “Survival in prostate carcinoma—outcomes from a prospective, population-based cohort of 8887 men with up to 15 years of follow-up: Results from three counties in the population-based National Prostate Cancer Registry of Sweden. Cancer 103:943–951. Beard, C. J., M. H. Chen, K. Cote, M. Loffredo, A. A. Renshaw, M. Hurwitz, and A.V. D’Amico. (2004). “Perineural invasion is associated with increased relapse after external beam radiotherapy for men with low-risk prostate cancer and may be a marker for occult, high-grade cancer.” Int J Radiat Oncol Biol Phys 58:19-24.
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DiBiase, S. J., K. Wallner, K. Tralins, and S. Sutlief. (2000). “Brachytherapy radiation doses to the neurovascular bundles.” Int J Radiat Oncol Biol Phys 46:1301–1307. Egevad, L., W. C. Allsbrook, Jr., and J. I. Epstein. (2005). “Current practice of Gleason grading among genitourinary pathologists.” Hum Pathol 36:5–9. Elliott, K., K. Wallner, G. Merrick, and P. Herstein. (2004). “Medical malpractice of prostate brachytherapy.” Brachytherapy 3:231–236. Feigenberg, S. J., A. L. Hanlon, E. M. Horwitz, R. G. Uzzo, D. F. Eisenberg, and A. Pollack. (2005). “What pretreatment prostate-specific antigen level warrants long-term androgen deprivation?” Int J Radiat Oncol Biol Phys 61:1003–1010. Fisch, B. M., B. Pickett, V. Weinberg, M. Roach. (2001). “Dose of radiation received by the bulb of the penis correlates with risk of impotence after three-dimensional conformal radiotherapy for prostate cancer. Urol 57:955–959. Freedland, S. J., W. J. Aronson, G. S. Csathy, C. J. Kane, C. L. Amling, J. C. Presti Jr., F. Dorey, and M. K. Terris. (2003). “Comparison of percentage of total prostate needle biopsy tissue with cancer to percentage of cores with cancer for predicting PSA recurrence after radical prostatectomy: results from the SEARCH database.” Urol 61:742–747. Galalae. R. M., A. Martinez, T. Mate, C. Mitchell, G. Edmundson, N. Nuernberg, S. Eulau, G. Gustafson, M. Gribble, and G. Kovács. (2004). “Long-term outcome by risk factors using conformal highdose-rate brachytherapy (HDR-BT) boost with or without neoadjuvant androgen suppression for localized prostate cancer.” Int J Radiat Oncol Biol Phys 58:1048–1055. Gelblum, D. Y., and L. Potters. (2000). “Rectal complications associated with transperineal interstitial brachytherapy for prostate cancer.” Int J Radiat Oncol Biol Phys 48:119–124. Gilbert, S. M., C. B. Cavallo, H. Kahane, and F. C. Lowe. (2005). “Evidence suggesting PSA cutpoint of 2.5 ng/mL for prompting prostate biopsy: Review of 36,316 biopsies.” Urol 65:549–553. Grann, A., and K. Wallner. (1998). “Prostate brachytherapy in patients with inflammatory bowel disease.” Int J Radiat Oncol Biol Phys 40:135–138. Grossfeld, G. D., J. J. Chang, J. M. Broering, Y. P. Li, D. P. Lubeck, S. C. Flanders, and P. R. Carroll. (2001). “Under staging and under grading in a contemporary series of patients undergoing radical prostatectomy: Results from the Cancer of the Prostate Strategic Urologic Research Endeavor database.” J Urol 165:851–856. Guillonneau, B., H. el-Fettouh, H. Baumert, X. Cathelineau, J. D. Doublet, G. Fromont, and G. Vallancien. (2003). “Laparoscopic radical prostatectomy: oncological evaluation after 1,000 cases at Montsouris Institute.” J Urol 169:1261–1266. Han, M., P. H. Gann, and W. J. Catalona. (2004). “Prostate-specific antigen and screening for prostate cancer.” Med Clin North Am 88:245–265. Hanks, G. E., A. L. Hanlon, B. Epstein, and E. M. Horwitz. (2002). “Dose response in prostate cancer with 8-12 years’ follow-up.” Int J Radiat Oncol Biol Phys 54:427–435. Harris, R., and K. N. Lohr. (2002). “Screening for prostate cancer: an update of the evidence for the U.S. Preventive Services Task Force.” Ann Intern Med 137:917–929. Holm, H. H., N. Juul, J. F. Pedersen, H. Hansen, and I. Strøyer. (1983). “Transperineal 125iodine seed implantation in prostatic cancer guided by transrectal ultrasonography.” J Urol 130:283–286. Holmberg, L., A. Bill-Axelson, F. Helgesen, J. O. Salo, P. Folmerz, M. Haggman, S. O. Andersson, A. Spangberg, C. Busch, S. Nordling, J. Palmgren, H. O. Adami, J. E. Johansson, and B. J. Norlen. (2002). “A randomized trial comparing radical prostatectomy with watchful waiting in early prostate cancer.” N Engl J Med 347:781–789. Horwitz, E. M., H. D. Thames, D. A. Kuban, L. B. Levy, P. A. Kupelian, A. A. Martinez, J. M. Michalski, T. M. Pisansky, H. M. Sandler, W. U. Shipley, M. J. Zelefsky, G. E. Hanks, and A. L. Zietman. (2005). “Definitions of biochemical failure that best predict clinical failure in patients with prostate cancer treated with external beam radiation alone: a multi-institutional pooled analysis.” J Urol 173:797–802. Howard, A., K. Wallner, and B. Han. (2001). “Clinical course and dosimetry of rectal fistulas after prostate brachytherapy.” J Brachytherapy Int 17:37–42. Hughes, S., K. Wallner, G. Merrick, S. Miller, and L. D. True. (2001). “Preexisting histologic evidence of prostatitis is unrelated to postimplant urinary morbidity.” Int J Cancer 96 Suppl:79–82.
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Hull, G. W., F. Rabbani, F. Abbas, T. M. Wheeler, M. W. Kattan, and P. T. Scardino. (2002). “Cancer control with radical prostatectomy alone in 1,000 consecutive patients.” J Urol 167:528–534. Jemal, A., T. Murray, E. Ward, A. Samuels, R. C. Tiwari, A. Ghafoor, E. J. Feuer, and M. J. Thun. (2005). “Cancer statistics, 2005.” CA Cancer J Clin 55:10–30. Jhaveri, F. M., C. D. Zippe, E. A. Klein, and P. A. Kupelian. (1999). “Biochemical failure does not predict overall survival after radical prostatectomy for localized prostate cancer: 10-year results.” Urol 54:884–890. Kattan, M. W. (2003). “Nomograms are superior to staging and risk grouping systems for identifying high-risk patients: preoperative application in prostate cancer.” Curr Opin Urol 13:111–116. Kiteley, R. A., W. R. Lee, A. F. deGuzman, M. Mirzaei, and D. L. McCullough. (2002). “Radiation dose to the neurovascular bundles or penile bulb does not predict erectile dysfunction after prostate brachytherapy.” Brachytherapy 1:90–94. Kundu, S. D., R. L. Grubb, K. A. Roehl, J. A. Antenor, M. Han, and W. J. Catalona. (2005). “Delays in cancer detection using 2 and 4-year screening intervals for prostate cancer screening with initial prostate specific antigen less than 2 ng/ml.” J Urol 173:1116–1120. Kupelian, P. A., J. C. Buchsbaum, C. Patel, M. Elshaikh, C. A. Reddy, C. Zippe, and E. A. Klein. (2002). “Impact of biochemical failure on overall survival after radiation therapy for localized prostate cancer in the PSA era.” Int J Radiat Oncol Biol Phys 52:704–711. Laumann, E. O., A. Paik, and R. C. Rosen. (1999). “Sexual dysfunction in the United States: prevalence and predictions.” JAMA 281:537–544. Lee, N., R. Fawaaz, C. A. Olsson, M. C. Benson, D. P. Petrylak, P. B. Schiff, E. Bagiella, A. Singh, R. D. Ennis. (2000). “Which patients with newly diagnosed prostate cancer need a radionuclide bone scan? An analysis based on 631 patients.” Int J Radiat Oncol Biol Phys 48:1443–1446. Lehrer, S., N. N. Stone, and R. G. Stock. (2005). “Prostate cancer in a large prostate is associated with a decreased prostate specific antigen failure rate after brachytherapy.” J Urol 173:79–81. Leibovich, B. C., M. L. Blute, D. G. Bostwick, T. M. Wilson, T. M. Pisansky, B. J. Davis, D. M. Ramnani, L. Cheng, T. J. Sebo, and H. Zincke. (2000). “Proximity of prostate cancer to the urethra: Implications for minimally invasive ablative therapies.” Urol 56:726–729. Litwin, M. S., D. P. Lubeck, J. M. Henning, and P. R. Carroll. (1998). “Differences in urologist and patient assessments of health related quality of life in men with prostate cancer: Results of the CaPSURE database.” J Urol 159:1988–1992. Merrick, G. S., W. M. Butler, W. H. Farthing, A. T. Dorsey, and E. Adamovich. (1998) “The impact of Gleason score accuracy as a criterion for patient selection in prostate brachytherapy.” J Brachytherapy Int 14:113–121. Merrick, G. S., W. M. Butler, J. H. Lief, R. L. Stipetich, L. J. Abel, and A. T. Dorsey. (1999). “Efficacy of sildenafil citrate in prostate brachytherapy patients with erectile dysfunction.” Urol 53:1112–1116. Merrick, G. S., W. M. Butler, J. H. Lief, and A. T. Dorsey. (2000). “Temporal resolution of urinary morbidity following prostate brachytherapy.” Int J Radiat Oncol Biol Phys 47:121–128. Merrick, G. S., W. M. Butler, A. T. Dorsey, and J. T. Dorsey III. (2000b). “The effect of constipation on rectal dosimetry following prostate brachytherapy.” Med Dosim 25:237–241. Merrick, G. S., W. M. Butler, R. W. Galbreath, J. H. Lief, and J. G. Donzella. (2000c). “A comparison of radiation dose to the neurovascular bundles in men with and without prostate brachytherapy induced erectile dysfunction.” Int J Radiat Oncol Biol Phys 46:1069–1074. Merrick, G. S., W. M. Butler, R. W. Galbreath, J. H. Lief, and E. Adamovich. (2001). “Perineural invasion is not predictive of biochemical outcome following prostate brachytherapy.” Cancer J 7:404–412. Merrick, G. S., W. M. Butler, J. H. Lief, and R. W. Galbreath. (2001a). “Five-year biochemical outcome after prostate brachytherapy for hormone-naive men < or = 62 years of age.” Int J Radiat Oncol Biol Phys 50:1253–1257. Merrick, G. S., K. E. Wallner, W. M. Butler, R. W. Galbreath, J. H. Lief, M. J. Benson. (2001b). “A comparison of radiation dose to the bulb of the penis in men with and without prostate brachytherapy induced erectile dysfunction.” Int J Radiat Oncol Biol Phys 50:597–604. Merrick, G. S., K. E. Wallner, W. M. Butler, J. H. Lief, and R. W. Galbreath. (2002). “Prophylactic versus therapeutic alpha-blockers in the management of urinary morbidity following permanent prostate brachytherapy.” Urol 60:650–655.
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Merrick, G. S., W. M. Butler, K. Wallner, R. W. Galbreath, R. L. Anderson, B. S. Kurko, and J. H. Lief. (2002a). “Permanent prostate brachytherapy-induced morbidity in patients with grade II and III obesity.” Urol60:104–108. Merrick, G. S., W. M. Butler, B. G. Tollenaar, J. H. Lief, R. L. Anderson, B. J. Smeiles, R. W. Galbreath, M. J. Benson. (2002b). “The dosimetry of prostate brachytherapy-induced urethral strictures.” Int J Radiat Oncol Biol Phys 52:461–468. Merrick, G. S., W. M. Butler, R. W. Galbreath, R. L. Stipetich, L. J. Abel, and J. H. Lief. (2002c). “Erectile function after permanent prostate brachytherapy. Int J Radiat Oncol Biol Phys 52:893–902. Merrick, G. S., W. M. Butler, K. E. Wallner, J. H. Lief, R. L. Anderson, B. J. Smeiles, R. W. Galbreath, and M. L. Benson. (2002d). “The importance of radiation doses to the penile bulb vs. crura in the development of post brachytherapy erectile dysfunction.” Int J Radiat Oncol Biol Phys 54:1055–1062. Merrick, G. S., W. M. Butler, K. E. Wallner, J. H. Lief, A. Mulroy, and R. W. Galbreath. (2003). “Prostate specific antigen (PSA) velocity and benign prostate hypertrophy predict for PSA spikes following prostate brachytherapy.” Brachytherapy 2:181–188. Merrick, G. S., W. M. Butler, K. E. Wallner, R. W. Galbreath, B. C. Murray, D. J. Zeroski, and J. H. Lief. (2003a). “Dysuria following permanent prostate brachytherapy.” Int J Radiat Oncol Biol Phys 55:979–985. Merrick, G. S., W. M. Butler, K. E. Wallner, R. W. Galbreath, B. Kurko, and S. Cleavinger. (2003b). “Rectal function following brachytherapy with or without supplemental external beam radiation: Results of two prospective randomized trials.” Brachytherapy 2:147–157. Merrick, G. S., W. M. Butler, K. E. Wallner, A. L. Hines, and Z. Allen. (2003c). “Late rectal function after prostate brachytherapy.” Int J Radiat Oncol Biol Phys 57:42–48. Merrick, G. S., W. M. Butler, K. E. Wallner, R. W. Galbreath, J. H. Lief, and E. Adamovich. (2004). “Prognostic significance of percent positive biopsies in clinically organ-confined prostate cancer treated with permanent prostate brachytherapy with or without supplemental external-beam radiation.” Cancer J 10:54–60. Merrick, G. S., W. M. Butler, K. E. Wallner, R. W. Galbreath, and E. Adamovich. (2004a). “Permanent interstitial brachytherapy for clinically organ-confined high-grade prostate cancer with a pretreatment PSA < 20 ng/mL.” Am J Clin Oncol 27:611–615. Merrick, G. S., W. M. Butler, K. E. Wallner, R. W. Galbreath, Z. Allen, J. H. Lief, and E. Adamovich. (2005). “Influence of body mass index on biochemical outcome after permanent prostate brachytherapy.” Urol 65:95–100. Merrick, G. S., W. M. Butler, K. E. Wallner, R. W. Galbreath, J. H. Lief, Z. Allen, and E. Adamovich. (2005a). “Impact of supplemental external beam radiotherapy and/or androgen deprivation therapy on biochemical outcome after permanent prostate brachytherapy.” Int J Radiat Oncol Biol Phys 61:32–43. Merrick, G. S., W. M. Butler, K. E. Wallner, Z. Allen, J. L. DeFilippo, and E. Adamovich. (2005b). “The role of enzymatic prostatic acid phosphatase in the clinical staging evaluation of patients with newly diagnosed, untreated prostate cancer.” W V Med J. In press. Merrick, G. S., W. M. Butler, K. E. Wallner, Z. A. Allen, R. W. Galbreath, and J. H. Lief. (2005c). “Brachytherapyrelated dysuria.” BJU Int 95:597–602. Merrick, G. S., W. M. Butler, K. E. Wallner, R. W. Galbreath, R. L. Anderson, and Z. A. Allen. (2005d). “Prostate brachytherapy-induced urethral strictures.” Submitted Merrick, G. S., W. M. Butler, K. E. Wallner, R. W. Galbreath, R. L. Anderson, B. S. Kurko, J. H. Lief, Z. A. Allen. (2005e). “Erectile function after prostate brachytherapy.” Int J Radiat Oncol Biol Phys. In press. Merrick, G. S., K. E. Wallner, and W. M. Butler. (2003). “Minimizing prostate brachytherapy-related morbidity.” Urol 62:786–792. Moran, B. J., M. A. Stutz, and M. H. Gurel. (2004). “Prostate brachytherapy can be performed in selected patients after transurethral resection of the prostate.” Int J Radiat Oncol Biol Phys 59:392–396. Mueller, A., K. Wallner, G. Merrick, E. Ford, S. Sutlief, W. Cavanagh, and W. Butler. (2004). Perirectal seeds as a risk factor for prostate brachytherapy-related rectal bleeding.” Int J Radiat Oncol Biol Phys 59:1047–1052. Mulhall, J. P., P. Yonover, A. Sethi, G. Yasuda, and N. Mohideen. (2002). “Radiation exposure to the corporeal bodies during 3-dimensional conformal radiation therapy for prostate cancer.” J Urol 167:539–542. Niehaus, A., G. S. Merrick, W. M. Butler, K. E. Wallner, Z. A. Allen, R. W. Galbreath, and E. Adamovich. (2005). “Relationship between isotope, prostate volume and urinary morbidity following prostate brachytherapy.” Int J Radiat Oncol Biol Phys. Submitted.
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Partin, A. W., L. A. Mangold, D. M. Lamm, P. C. Walsh, J. I. Epstein, and J. D. Pearson. (2001). “Contemporary update of prostate cancer staging nomograms (Partin Tables) for the new millennium.” Urol 58:843–848. Pickett, B., R. K. Ten Haken, J. Kurhanewicz, A. Qayyum, K. Shinohara, B. Fein, and M. Roach. (2004). “Time to metabolic atrophy after permanent prostate seed implantation based on magnetic resonance spectroscopic imaging.” Int J Radiat Oncol Biol Phys 59:665–673. Pollack, A., A. L. Hanlon, B. Movsas, G. E. Hanks, R. Uzzo, and E. M. Horwitz. (2003). “Biochemical failure as a determinant of distant metastasis and death in prostate cancer treated with radiotherapy.” Int J Radiat Oncol Biol Phys 57:19–23. Ponsky, L. E., E. E. Cherullo, R. Starkey, D. Nelson, D. Neumann, and C. D. Zippe. (2002). “Evaluation of preoperative ProstaScint scans in the prediction of nodal disease.” Prostate Cancer Prostatic Dis 5:132–135. Potters, L., T. Torre, P. A. Fearn, S. A. Leibel, and M. W. Kattan. (2001). “Potency after permanent prostate brachytherapy for localized prostate cancer.” Int J Radiat Oncol Biol Phys 50:1235–1242. Pound, C. R., A. W. Partin, M. A. Eisenberger, D. W. Chan, J. D. Pearson, and P. C. Walsh. (1999). “Natural history of progression after PSA elevation following radical prostatectomy.” JAMA 281:1591–1597. Roach, M., 3rd, A. Chen, J. Song, A. Diaz, J. Presti, Jr., and P. Carroll. (2000). “Pretreatment prostate-specific antigen and Gleason score predict the risk of extracapsular extension and the risk of failure following radiotherapy in patients with clinically localized prostate cancer.” Semin Urol Oncol 18:108–114. Sharkey, J., A. Cantor, Z. Solc, W. Huff, S. D. Chovnick, R. J. Behar, R. Perez, J. Otheguy, and R. Rabinowitz. (2005). “(103)Pd brachytherapy versus radical prostatectomy in patients with clinically localized prostate cancer: A 12-year experience from a single group practice.” Brachytherapy 4:34–44. Snyder, K. M., R. G. Stock, S. M. Hong, Y. C. Lo, and N. N. Stone. (2001). “Defining the risk of developing grade 2 proctitis following 125I prostate brachytherapy using a rectal dose-volume histogram analysis.” Int J Radiat Oncol Biol Phys 50:335–341. Stamey, T. A., M. Caldwell, J. E. McNeal, R. Nolley, M. Hemenez, and J. Downs. (2004). “The prostate specific antigen era in the United States is over for prostate cancer: what happened in the last 20 years?” J Urol 172:1297–1301. Stock, R. G., J. Kao, and N. N. Stone. (2001). “Penile erectile function after permanent radioactive seed implantation for treatment of prostate cancer.” J Urol 165:436–439. Stock, R. G., N. N. Stone, and J. A. Cesaretti. (2003). “Prostate-specific antigen bounce after prostate seed implantation for localized prostate cancer: descriptions and implications.” Int J Radiat Oncol Biol Phys 56:448–453. Stone, N. N., R. G. Stock, and P. Unger. (2005). “Intermediate term biochemical-free progression and local control following 125iodine brachytherapy for prostate cancer.” J Urol 173 803–807. Talcott, J. A., J. A. Clark, P. C. Stark, and S. P. Mitchell. (2001). “Long-term treatment related complications of brachytherapy for early prostate cancer: A survey of patients previously treated.” J Urol 166:494–499. Terk, M. D., R. G. Stock, and N. N. Stone. (1998). “Identification of patients at increased risk for prolonged urinary retention following radioactive seed implantation of the prostate.” J Urol 160:1379–1382. Thompson, I. M., P. J. Goodman, C. M. Tangen, M. S. Lucia, G. J. Miller, L. G. Ford, M. M. Lieber, R. D. Cespedes, J. N. Atkins, S. M. Lippman, S. M. Carlin, A. Ryan, C. M. Szczepanek, J. J. Crowley, and C. A. Coltman Jr. (2003). “The influence of finasteride on the development of prostate cancer.” N Engl J Med 349:215–224. Valicenti, R., J. Lu, M. Pilepich, S. Asbell, and D. Grignon. (2000). “Survival advantage from higher-dose radiation therapy for clinically localized prostate cancer treated on the Radiation Therapy Oncology Group trials.” J Clin Oncol 18:2740–2746. Wallner, K. (2002). “MR imaging for prostate cancer staging: Beauty or beast?” Int J Radiat Oncol Biol Phys 52:886–886. Wallner, K., J. Roy, and L. Harrison. (1995). “Dosimetry guidelines to minimize urethral and rectal morbidity following transperineal I-125 prostate brachytherapy.” Int J Radiat Oncol Biol Phys 32:465–471. Wallner, K. E., J. Roy, and L. Harrison. (1996). “Tumor control and morbidity following transperineal iodine 125 implantation for stage T1/T2 prostatic carcinoma.” J Clin Oncol 14:449–453. Wallner, K., G. Merrick, L. True, W. Cavanagh, C. Simpson, and W. Butler. (2002). “I-125 versus Pd-103 for low risk prostate cancer: Morbidity outcomes from a prospective randomized multicenter trial.” Cancer J 8:67–73. Wallner, K. E., G. S. Merrick, M. L. Benson, W. M. Butler, J. Maki, B. G. Tollenaar. (2002a). “Penile bulb imaging.” Int J Radiat Oncol Biol Phys 53:928–933.
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Waterman, F. M., and A. P. Dicker. (1999). “Effect of post-implant edema on rectal dose in prostate brachytherapy.” Int J Radiat Oncol Biol Phys 45:571–576. Whitmore, W. F., Jr. (1990). “Natural history of low-stage prostatic cancer and the impact of early detection.” Urol Clin North Am 17:689–697. Williams, S. G. (2004). “Ambiguities within the ASTRO consensus definition of biochemical failure: never assume all is equal.” Int J Radiat Oncol Biol Phys 58:1083–1092. Wilson, S. S., and E. D. Crawford. (2004). “Screening for prostate cancer.” Clin Prostate Cancer 3:21–25. Zietman, A. L., E. M. Zehr, and W. U. Shipley. (1999). “The long-term effect on PSA values of incidental prostatic irradiation in patients with pelvic malignancies other than prostate cancer.” nt J Radiat Oncol Biol Phys 43:715–718. Zelefsky, M. J., and J. F. Eid. (1998). “Elucidating the etiology of erectile dysfunction after definitive therapy for prostate cancer.” Int J Radiat Oncol Biol Phys 40:129–133. Zelefsky, M. J., and W. F. Whitmore Jr. (1997). “Long-term results of retropubic permanent 125iodine implantation of the prostate for clinically localized prostatic cancer.” J Urol 158:23–29. Zelefsky, M. J., K. E. Wallner, C. C. Ling, A. Raben, T. Hollister, T. Wolfe, A. Grann, P. Gaudin, Z. Fuks, and S. A. Leibel. (1999). “Comparison of the 5-year outcome and morbidity of three-dimensional conformal radiotherapy versus transperineal permanent iodine-125 implantation for early stage prostate cancer.” J Clin Oncol 17:517–522. Zelefsky, M. J., Z. Fuks, M. Hunt,Y. Yamada, C. Marion, C. C. Ling, H. Amols, E. S. Venkatraman, and S. A. Leibel. (2002). “High-dose intensity modulated radiation therapy for prostate cancer: early toxicity and biochemical outcome in 772 patients.” Int J Radiat Oncol Biol Phys 53:1111–1116.
Chapter 29
Permanent Prostate Brachytherapy Treatment Planning Wayne M. Butler, Ph.D. and Gregory S. Merrick, M.D. Schiffler Cancer Center Wheeling Hospital Wheeling, West Virginia Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 559 Target Volume Definition . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 560 Imaging Modalities . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 560 What Is the Target Volume? . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 560 Prostate Only—or Less . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 562 Planning with a Dosimetric Margin . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 562 Explicit Planning Target Volume (PTV) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 563 Should the Seminal Vesicles Be Included? . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 563 Clinical Necessity for Dosimetric Margins . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 564 How Much Margin Is Necessary? . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 564 Creating the PTV . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 564 Symmetry . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 565 Planning Types . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 565 Pre-OR Planning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 567 Intraoperative and Interactive Planning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 567 Nomogram Approaches . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 567 Computer Planning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 568 Dynamic Dose Calculation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 568 Needle and Seed Placement Approaches . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 569 General Guidelines . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 569 Uniform Loading: Idealism vs. Reality . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 570 Manual Loading Algorithm . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 570 Modified Uniform Loading . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 572 How to Cover the Apex . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 572 Addition of Peripheral Needles . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 572 Removal of Selected Central Seeds . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 572 A Sample Plan . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 573 Peripheral Loading . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 576 Computer-Assisted and Optimized Planning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 577 Convergence of Approaches . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 578 Choice of Radionuclide . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 578 Appropriate Seed Strengths . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 578 Preplan Evaluation Criteria . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 579 Review of Contemporary Planning Variations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 581 Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 585 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 585
Introduction The ability to learn and plan is part of human nature, which is a good thing because we certainly have no hard-wired instinct for constructing a brachytherapy implant. What does come in handy in designing permanent prostate brachytherapy implants is our built-in aesthetic sensibility of symmetry, balance, and proportion. Effective treatment planning requires an understanding not only of the characteristics of various brachytherapy sources, but also an understanding of the disease being treated.
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Target Volume Definition Identification of the target volume in prostate brachytherapy initiates the treatment-planning process. Because physicists and dosimetrists rarely have expertise in urogenital anatomy, the radiation oncologist is usually tasked with target delineation; but the task is sometimes delegated to the physics staff. The physics staff may not be charged with outlining the prostate, but they are usually expected to define the planning target volume (PTV). The expansion of the clinical target volume (CTV) to the PTV, whether performed explicitly or implicitly, should be performed initially with guidance from the radiation oncologist and always with insight.
Imaging Modalities Prostate brachytherapy would probably not have become ascendant were it not for the relatively low cost, simplicity, and utility of transrectal ultrasound. Ultrasound systems usually have several ways to export images for treatment planning: hard copy prints which may be digitized or electronic transfer via video tape, stored sets of bitmap or jpeg images, or network DICOM transfer. Quality control checks and measurements should be periodically performed on the ultrasound system to identify image quality degradation before it affects patient scans (Goodsitt et al., 1998). Because the needle template grid is projected onto transverse ultrasound images, proper scaling and geometric accuracy of the projected grid should also be checked by phantom measurements. Magnetic resonance imaging (MRI) provides excellent visualization and is used instead of ultrasound at a few centers (Cormack et al., 2000), but the additional implant time under anesthesia and expense has not been justified by improved patient dosimetry or biochemical control and quality of life outcomes compared with centers performing conventional ultrasound based implants. Computerized tomographic (CT) imaging has also been used for planning and design of customized templates (Roy et al., 1991). Prostate visualization with CT is inferior to that achieved by ultrasound, and the approach has fallen into disuse.
What Is the Target Volume? Most commonly, a transrectal ultrasound volumetric study of the prostate gland is obtained in the ultrasound suite, preferably in the radiation oncology department, prior to the operative procedure. The preimplant volume study should not be performed on patients undergoing hormonal therapy unless the initiation of androgen deprivation therapy was at least 2 months prior. Such agents are cytoreductive, and the initial prostate volume, V0, will be diminished by the number of weeks of androgen deprivation therapy. For seed ordering purposes, we use an empirically derived formula to predict the volume change from combined use of a luteinizing hormone-releasing hormone (LH-RH) agonist and an antiandrogen approximately as: Vfinal =
weeks 7 1 + 2 2
V0
(1)
Equation (1) is plotted in Figure 1. There will be a 30% volume reduction of the prostate after 2 months and about a 40% reduction after 3 months. Implanting a patient during the early course of hormonal therapy will result in a considerable dose escalation as the seeds draw closer together. It is imperative that each patient and his referring physicians be queried regarding the use of hormones. The volume study is acquired with the patient in the dorsal lithotomy position using a transrectal ultrasound probe mounted securely to a stepper and stabilizer unit (Figure 2). The patient is scanned at 3 to 5
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Figure 1. Plot of relative prostate volume as a function of the duration of hormonal therapy. Avodart® is an example of a class of drugs called 5-alpha-reductases which block the conversion of testosterone to its active form, dihydroxytestosterone. Lupron® is a luteinizing hormone-releasing hormone agonist that denies the testicles the hormone necessary to produce testosterone. For total androgen blockade, these drugs are usually combined with an antiandrogen such as Casodex®. Androgen deprivation therapy, if used at all, should be initiated at least 2 months before a permanent prostate implant, and if not used prior to implant, should be delayed after the implant for at least 2 half-lives of the radionuclide because of the inevitable dose escalation in a shrinking prostate.
Figure 2. A floor-stand stabilizer with an adjustable stepper unit attached. The ultrasound probe and needle template are attached to the stepper. Virtually all stepper/stabilizer units allow independent 6-axis probe movement. Table-mounted stabilizers are also in common use, but they are not as rigid as a floor-mounted unit. The floor stand allows the apparatus to be easily wheeled out of the way for post implant cystoscopy and reproducible reinsertion of the probe for any last minute changes. (Illustration courtesy of CMS, Inc.)
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mm intervals starting from the proximal seminal vesicles just superior to the bladder neck/base of the prostate gland and extending to the apex. The vertical midline of the prostate should be in the center of the transverse field from base to apex. The posterior border of the prostate should be under sufficient pressure from the ultrasound probe so that it is flat at midgland, neither concave up nor down, and the border should be positioned at a reference row of template markers or no more than 2 mm below that row. The importance of consistency in volume studies cannot be overstressed. To ensure that no slices were omitted or duplicated, a sagittal image of the prostate should be recorded, and the measuring tools available in the ultrasound software used to determine the distance from the base to the apex. That distance should correspond to the separation between the transverse base and apex images. The path of the urethra should be visualized on each ultrasound slice by either infusing an aerated gel via the penile urethra or by placing a urinary catheter prior to the procedure. It is very important to ensure that the catheter bulb is well within the bladder and that there is no tension applied to the catheter in order to minimize catheter distortion of the prostate. Other structures frequently contoured are the rectum and bladder. Long-term serious brachytherapy-induced urinary and rectal morbidity occur rarely (~1%) in good programs, and that portion of morbidity attributable to high radiation dose can be reduced to very rare status by careful planning and execution. Prostate Only—or Less In institutions where radical prostatectomy is considered the “gold standard” or the prevailing treatment paradigm, the target volume is frequently defined as the prostate only with minimal margins in analogy with surgery. If the goal is to duplicate the surgical experience, then the brachytherapy results will be no better than those reported for surgical series. As illustrated in the previous chapter, the best-reported biochemical survival results following surgery are good, but not as good as the best brachytherapy results. The centers reporting the best brachytherapy results almost all treat with dosimetric margins, and one study reported that a dosimetric margin >3 mm was almost as good a predictor of freedom from biochemical failure as the widely used minimum dose covering 90% of the prostate, D90 (Choi et al., 2004). With currently available technology, an even riskier approach is to attempt the brachytherapy equivalent of a lumpectomy. Using the location of positive biopsies supplemented by tumor sensitive imaging such as MRI spectroscopy or immunoscintigraphy and single-photon emission computed tomography, only cancer foci or prostate zonal anatomy, typically the peripheral zone, are targeted. A sextant needle biopsy sampling, with each core 1 mm in diameter and 2 cm long, samples 2% (Partin et al., 2001). For high-risk patients with two or more unfavorable parameters, the risk of seminal vesicle involvement is >6%. In radical prostatectomy series where the seminal vesicles are taken, seminal vesicle disease extension is often confined to the first centimeter proximal to the prostate (Kestin et al., 2002). For this reason, patients treated with externalbeam radiation therapy (XRT) adjunctive to brachytherapy have their seminal vesicles included in the treatment portals. Some brachytherapists go further and explicitly place seeds in the vesicles as illustrated in Figure 3. However, it is less common that the seminal vesicles are included in either the planning or post-implant PTV.
Figure 3. Transverse ultrasound image of the seminal vesicles with four implanted seeds visible (arrow). Typically, four needles containing two seeds each, 1 cm apart, are used to implant the proximal seminal vesicles of patients treated at the Schiffler Cancer Center.
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Clinical Necessity for Dosimetric Margins In patients with a pretreatment PSA < 10 ng/mL, approximately 50% manifest extracapsular disease at the time of radical prostatectomy (Partin et al., 2001). Increasing PSA, Gleason score, and clinical stage are markers for an increased likelihood that the disease is not organ-confined. Many surgical series have noted an adverse effect on biochemical progression free survival due to additional factors that are also surrogates for increased likelihood and extent of extracapsular disease: increasing percent positive biopsies, perineural invasion, and a 4+3 pattern in Gleason score 7 rather than a 3+4 pattern. Unlike urologic surgeons, brachytherapists are not constrained in their efforts to treat the periprostatic region. In brachytherapy implants designed to eradicate extraprostatic disease, adverse factors are of much less consequence (Merrick et al., 2001, 2002, 2004). It is our contention that treating with adequate dosimetric margins results in the superior survival results seen in the best brachytherapy series compared with the best radical prostatectomy series. How Much Margin Is Necessary? The mean extent of extraprostatic extension in radical prostatectomy specimens has been reported by the Mayo Clinic (Davis et al., 1999) and the Cleveland Clinic (Sohayda et al., 2000) to be 0.5 mm and 1.1 mm, respectively. The maximum extracapsular extension at the Mayo and Cleveland Clinics was 4.4 mm and 10.0 mm, respectively. Both clinics concluded that brachytherapy margins of 5.0 mm would encompass about 99% of cases amenable to local control. One caveat to this conclusion is that the conventional H&E (haematoxylin-eosin) staining used in these studies may not completely define the radial extraprostatic extension. The determination of molecular margins after radical prostatectomy by reverse transcriptase-polymerase chain reaction for prostate-specific antigen has questioned the validity of extracapsular extension and surgical margin assessment by standard staining (Theodorescu, Frierson, and Sikes 1999; Straub et al., 2001). However, even if H&E staining does underestimate the extent of extracapsular disease, it is likely that in contemporary patient populations disease extending beyond 5 mm still occurs rarely. Capsular penetration beyond 5 mm was an extremely significant predictor of failure in univariate analysis of a prostatectomy series and was so strongly correlated with lymph node involvement that it was not an independent predictor in multivariate failure analysis (Stamey et al., 2000). Figure 4 displays biochemical PSA progression-free survival curves for implants with and without margin and with and without XRT. Treating high-risk prostate cancer with brachytherapy combined with XRT is considered standard of care, because omitting XRT will reduce the 8-year biochemical survival from about 85% down to 65%. Omitting XRT and treating less than the entire prostate and the likely extracapsular disease extension is profoundly detrimental—all the patients (n = 19) in the D’Amico study (D’Amico et al., 1998) failed within 3 years. The other data in the survival figure are from Dattoli (2005), Merrick et al., (2005), and Sylvester et al., (2003).
Creating the PTV The auto-margin tools available in modern treatment planning software are sufficient to create the PTV from the defined prostate. Such tools allow specification of separate margins for at least the 4 Cardinal directions in the transverse plane as well as superior and inferior margins. If the superior and inferior margins are set equal to the slice separation, one can duplicate the shape of the prostate base on a superior slice and duplicate the prostate apex on an inferior slice. All other intervening slices will have a PTV at least as large as the prostate on any adjacent slice and will be further expanded by the chosen transverse margins in areas where the prostate on the current slice is larger than the prostate on adjacent slices. By choosing the superior and inferior margins to be slightly greater than the slice separation, the PTV on each slice will be at least equal to the size of the prostate on any adjacent slice with a slight additional
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Figure 4. Biochemical PSA progression-free survival in high-risk patients stratified by the use of planning margins and external beam radiation therapy. Data are from Merrick et al. (2005), Dattoli (2005), Blasko (Sylvester et al., 2003), and D’Amico et al. (1998).
margin. Margin settings of 4 to 6 mm laterally and anteriorly, 0 mm posteriorly, and 6 to 7 mm superiorly and inferiorly, work well with our patients. Planning with margin has been shown to result in similar margins on post implant analysis, as shown in Figure 5 (Merrick et al., 2003). At the base, and to a much lesser degree at the apex, the posterior border of the prostate is separated from the rectum. Posterior margin may be added to those slices, but the margin should not extend below the lowest implant row. The closest approach of the rectum to the prostate is near the apex, and almost all brachytherapy-induced as well as radical prostatectomy induced rectal injuries occur at the apex. Symmetry One deficiency of auto-margin brachytherapy software is the inability to create a more generous margin at either the base or the apical regions as per the brachytherapist’s request. For this, manual intervention is needed. Another problem is how to cope with asymmetry in the prostate volume study. The prostate of the anesthetized patient in the operating room can usually be aligned so that it is bilaterally symmetric. Patient movement or spasms often cause the asymmetry observed in the volume study obtained in the clinic. In anticipation of the likely situation in the operating room and to make treatment planning, needle loading, and plan execution simpler and more robust, the PTV should be manually adjusted to make it bilaterally symmetric.
Planning Types Table 1 lists the type of planning applicable to prostate brachytherapy as defined by the American Brachytherapy Society (ABS) (Nag et al., 2001). It is likely that most centers will continue to plan permanent seed implants before entering the operating room (OR) even if technology advances to the ultimate level of dynamic dose calculation and most centers adopt it. The intervening types are currently available technologies for intraoperative planning.
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Figure 5. Sagittal views of the dosimetric margin from the prostate (shaded interior) to the 100% isodose line (solid curve), the 90% isodose line (dashed curve), and the 75% isodose line (dotted curve). The plots are a composite of 13 implants each of 125I (a) and 103Pd (b). (Reproduced from Am J Clin Oncol, vol 26, “Extracapsular radiation dose distribution after permanent prostate brachytherapy,” G. S. Merrick, W. M. Butler, K. E. Wallner, L. R. Burden, and J. E. Dougherty, pp. e179–e189. © 2003, with permission from Lippincott Williams and Wilkins.)
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Table 1. American Brachytherapy Society (ABS) Nomenclature for Different Types of Planning Used in Permanent Prostate Brachytherapy (Nag et al., 2001). Planning approach
Definition
Preplanning
Creation of a plan outside the operating room (OR) hours, days or weeks before the implant procedure.
Intraoperative
Plan created in the OR. The patient remains stationary between the time of the volume study and the implant procedure.
Interactive
The treatment plan is revised periodically during the implant procedure using image-based feedback of needle position to recalculate dose.
Dynamic dose calculation
Dose distribution continuously updated using deposited seed position feedback.
(Reproduced from Int J Radiat Oncol Biol Phys, vol 51, “Intraoperative planning and evaluation of permanent prostate brachytherapy: Report of the American Brachytherapy Society.” S. Nag, J. P. Ciezki, R. Cormack, S. Doggett, K. DeWyngaert, G. K. Edmundson, R. G. Stock, N. N. Stone, Y. Yu, and M. J. Zelefsky, pp. 1422–1430. © 2001, with permission from Elsevier.)
Pre-OR Planning Among the virtues of planning prior to encountering the patient in the OR is cost effectiveness. With the volume study performed by the brachytherapist no more than a few weeks prior to the implant, there will be negligible change in the prostate, and fewer than 1% of patients exhibit changes from the time of the volume study to the operating room volume capture that necessitate modification of the plan. The plan, therefore, will entail little or no seed waste. Having a plan also saves valuable OR time and allows the use of pre-loaded needles. This planning approach should result in an implant of better dosimetric quality and one that is more resistant to seed loss or misplacement. Taking the time for a thoughtful, insightful consideration of alternatives, optimizing the fit between seed strength and prescribed dose, manually verifying calculations, and second checking assumptions are all worthwhile activities that minimize errors and lead to good outcomes. Mistakes are made most often when established procedures are not followed, and shortcuts are most likely taken when staff is under deadline pressure. Prostate cancer does not require emergency, heroic measures. Foregoing a pre-OR plan treats prostate brachytherapy as emergency medicine, and the attendant time stress on the planning and implant team appears to invite errors. Frequent mismatches between the prostate volume, shape, or position in the pre-OR study and that performed in the OR led some institutions to forego the former for the latter. The anesthetized patient has much less movement artifact, but the root causes of much of the difference seen in studies performed at the two times are differences in technique, apparatus, and patient positioning. It is important that the same individual, whether an ultrasonographer or brachytherapist, set up the patient in both environments and use similar tables and patient leg support in both.
Intraoperative and Interactive Planning Nomogram Approaches Intraoperative prostate brachytherapy planning has a long history, and most early practitioners employed Lowell Anderson’s nomogram (Anderson 1976) or modified versions of it (Stock et al., 1995). In the Stock implementation, the prostate volume was estimated using a urologist’s measurement of length, width, and height (l, w, h, respectively) and determining the ellipsoidal volume from
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Ve =
π 6
⋅l ⋅w⋅h
(2)
An activity nomogram is then used to determine the total seed strength to bring to the OR. In the OR, an ultrasound planimetry determined volume is used to determine the total activity to implant. Dividing by the available individual seed strength gives the total number of seeds to be implanted. Their nomogram calls for 75% of the total strength to be placed on the periphery of the gland. Stock and colleagues accordingly place needles spaced about 1 cm apart around the periphery of the prostate so that each needle lies about 0.5 cm inside the border of the prostate on the largest mid-gland transverse image. Dividing the number of seeds designated for the periphery by the number of needles placed, the number of seeds per peripheral needle is calculated. The spacing between seeds along each needle track is determined by the base to apex length of the gland with the proviso that the center-to-center spacing be ≤1 cm. Interior needles are placed in a similar fashion. These needles are placed about 1 cm from the periphery and 1 cm from each other. Urethral sparing is achieved by requiring that no needle be closer than 0.5 cm to the urethra. Although users of the nomogram approach to intraoperative planning treat only the prostate rather than a PTV that includes margins, there is no reason why the approach could not be modified to treat the PTV. Auto-margining and calculation of the new planning volume would only take a minute. However, current nomograms make no allowance for the effect of prostate shape on total seed strength needed. Treatment planning dose optimization produces significantly different total strengths when the same volume is shaped as a cylindrical, triangular, or rectangular column. Because nonstandard seed spacing is difficult to implement in the OR, nomogram-based planning almost mandates proficiency with the Mick seed applicator. A recent analysis of the post-operative dosimetry results from practitioners at 35 centers trained by Stock and Stone in their “ProSeed” nomogram approach indicates that 90% of patients implanted with a prescribed 160 Gy monotherapy 125I achieved their day 30 dosimetry goal of D90 ≥ 140 Gy (Stone et al., 2005). Computer Planning All the elements of treatment planning can be moved from the clinic to the OR, but that consumes too much valuable OR time for most, so various hybrid approaches have been developed. Beyer, Shapiro, and Puente (2000) save time by using standard needle positions for all patients. Treatment planning is performed based on those default needle positions while the needles are being placed. The next step up the hierarchy is to evaluate the plan based on the actual needle location. This was first reported by Zelefsky and colleagues at Memorial Sloan-Kettering Cancer Center (Zelefsky et al., 2000; Todor et al., 2003). The positions of the needles on ultrasound were captured by the planning system, and seed placement in the needles was determined by in-house optimization software. Users of commercial software can accomplish this by identifying the needle tips on the transverse ultrasound image or the entire needle on sagittal projections. The planning software then either modifies an existing plan to account for a changed needle position or augments a plan being developed sequentially with each additional needle.
Dynamic Dose Calculation Although an approach similar to needle projection discussed above can be applied to identify and localize seeds as they are extruded from each needle, the process is tedious and time-consuming and does not account for perturbations caused by later needle insertions or seed migration or reorientation (Van Gellekom et al., 2004; ;Rivard, Evans, and Kay 2005). Real-time dynamic dose calculation requires automatic seed identification and registration with prostate imaging. Most of the effort on the seed identification
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task has focused on the use of a limited number of x-ray images (Su et al., 2004; Narayanan, Cho, and Marks 2004). The seed coordinates would then be fused to the ultrasound image set. The latter step is proving to be as challenging as the first. Perhaps the breakthrough will come in ultrasound imaging as the sole modality for anatomy and seeds. Trans-urethral ultrasound has been shown to more clearly identify seeds with better spatial resolution and better soft tissue differentiation than trans rectal ultrasound (Holmes et al., 2003).
Needle and Seed Placement Approaches General Guidelines In all treatment planning philosophies, it is considered poor technique to use single seed needles as in Figure 6. If that single seed is intended to address a local dosimetric defect, consider the imaging uncertainties and the seed placement uncertainties. Roberson and colleagues (1997) found the placement uncertainty to be about 3 mm transversely and 5 mm longitudinally. In an implant with 24 needles and 100 seeds, is that single seed worth the trauma of an additional needle? Most immediate post-operative morbidity from prostate brachytherapy is due to needle trauma. As a corollary to our single seed needle advice, the planner should minimize the number of specially loaded needles, i.e., those with other than a standard seed to seed spacing. Increasing the complexity of the plan increases the likelihood of errors not only in loading the needles (or placing seeds via the Mick applicator), but also because the altered loading is designed to modulate the dose in a specific pattern. Consider placement uncertainty above and whether that modulation would be tolerable if shifted cranially or caudally by 5 mm.
Figure 6. Targeting crucial areas with single seeds as in (a) is risky because of seed placement and visualization uncertainties. A more robust approach is shown in (b) where two or three seeds are deployed to cover the areas of interest.
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As a final general rule, be sparing in your use of back-to-back seeds. These are typically used to build up dose at the prostate base and apex. Double sources create contiguous high dose regions that may place the bladder neck and the bulbomembranous urethra at risk. Incontinence may result if the bladder sphincter is damaged, and urethral strictures may occur if the bulbomembranous urethra is overdosed (Merrick et al., 2002a).
Uniform Loading: Idealism vs. Reality Haakon Ragde, John Blasko, and Peter Grimm began transperineal, ultrasound guided prostate brachytherapy in Seattle, Washington in 1985. They used a uniform seed loading approach in which a relatively large number of low strength seeds were evenly distributed throughout the prostate (Sylvester 2005). An ideal prostate brachytherapy implant would take the form of microscopically dispersed particles of a beta emitting or low-energy photon-emitting radionuclide. Of course this is the same ideal as the radiopharmaceutical industry to find a molecular carrier with exquisite binding specificity to cell membrane molecules that are unique to the target tissue. Alas, the radionuclides used in prostate brachytherapy are encased in macroscopic seeds, and even if they were microscopically dispersible, the emitted photons are not of sufficiently low energy for cumulative, long-range effects to be negligible. In any prostate volume in which seeds are spaced at the corners of a 1 cm cubic grid, the central dose will be much higher than the peripheral dose because of such cumulative effects. Central prostate and urethral doses frequently exceeded 300% of the prescribed dose while the coarse 1 cm cubic grid produced a peripheral scalloping of the 100% isodose that would fail to cover the anterior-lateral curvature of the prostate completely. Unacceptably high urinary morbidity led the Seattle group to abandon uniform loading in favor of a modified version within 2 years after the start of their program. The uniform loading approach is not just an historical curiosity, but forms the basis for most manually planned implants today. Figure 7 illustrates a schematic prostate and typical template markings. The template grid rows and columns are spaced 0.5 cm apart, but the markings are 1 cm apart. The rows are labeled with numbers and the columns with letters. The dotted cross hairs on each image are formed by the D-column and the 3-row. The template columns adjacent to the uppercase letter columns are labeled with lower case letters a…f. The unmarked rows are half-integer numbers, 1.5, 2.5, etc. Manual Loading Algorithm The following paragraphs describe a manual needle and seed placement algorithm that starts with uniform loading, then applies modifications. The modifications create a seed distribution that is urethra sparing and relatively homogeneous—for brachytherapy. Such a uniform dose distribution is very tolerant of systematic seed placement errors (Butler et al., 2000). Because the urethra usually aligns along the D-column, most plans avoid that location. To cover the prostate at mid-gland with a uniform 1 cm square grid, one is obligated to use lowercase columns. The most posterior row implanted in Figure 7 is the 1.5 row, and the most anterior is the 5.5 row. The needles in the figure are placed on the intersection of every lowercase column with every half-integer row that intersects the PTV or is no more than 1 or 2 mm outside the PTV on the largest mid-gland image. Seeds are placed along these valid needle tracks at 1 cm increments from the base (0.0 cm plane) to the apex. If the needle is not within the PTV on the base plane but is within the PTV on the 0.5 cm plane, start the seed train on the base plane. If the needle does not intersect the PTV until the 1 cm plane, start the seed train on either the 1 cm plane or the 0.0 cm plane, and continue placing the seeds at 1 cm increments until they either fall on the last plane of the PTV intersected by the needle. Seeds that are more than 5 mm outside the PTV are omitted.
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Figure 7. Schematic of the prostate (thin solid line) and the Wheeling approach to the definition of the planning target volume (PTV, heavy solid line) at the base, midgland and apex of the prostate. The central open circle marks the urethra and the small open or filled circles indicate needle paths and seed positions in a uniform loading approach. The prostate on each slice is expanded 4 to 5 mm to create the PTV. Additional expansion accounts for the size of the prostate on the next larger adjacent slice along with uncertainties in the position of the needle tip and ultrasound transducer plane. Note the bilateral symmetry imposed on the PTV in the midgland slices rather than maintain the skew or asymmetry of the prostate. The PTV is also extended inferior to the apex.
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Modified Uniform Loading How to Cover the Apex The first modification to the uniform seed loading considers variations in how to treat the apex. If the apical PTV ends at a whole cm increment such as 4.0 cm, then there will be no problem with coverage. Note that the prostate apex of the model in Figure 7 occurs at 40 mm but that the apical portion of the PTV extends to 45 mm from the base and contains no seeds, making it dosimetrically cold. How you correct that deficiency depends on the measured sagittal length of the gland and the concurrence of the radiation oncologist on how to proceed. One solution is to extend the uniform loaded seed trains to the 5.0 cm plane. This is the favored option if the sagittal length of the prostate is nearly 4.5 cm long, implying that the 4.5 cm plane just barely missed seeing prostate. Placing seeds on the 5.0 cm plane, therefore, does not create an excessive inferior margin, and the 5.0 cm seed plane will sandwich the apical PTV between two heavily loaded seed planes to provide coverage. On the model prostate, repeating the 4 ¥ 4 seed pattern of the 4.0 plane on the 5.0 plane would be a good start. The 4 corner seeds of that 16-seed pattern would be unnecessary because the area of the PTV on the 4.5 cm plane is small. The other solution for apical PTV coverage is to select a few of the uniform loading pattern needles for back-to-back seed spacing so that some seeds will lie on the 4.5 cm plane. A pattern using this approach is discussed in detail below. Truncating plans at half integer transverse planes was once rare in our practice but has become very common since our discovery of a strong correlation between radiation dose to the bulbomembranous urethra and urinary strictures (Merrick et al., 2002a) and dose to the proximal penis and patient preservation of erectile function (Merrick et al., 2001a, 2002b). Addition of Peripheral Needles The second modification is to place additional needles along the lateral and anterior periphery on template grid positions that are offset 5 mm in the y direction and 0 or 5 mm horizontally from the main uniformly loaded grid. Usually, six or fewer peripheral needles are sufficient to deliver seeds to cover the anterior and lateral curvature of the prostate completely. As with the basic uniform needle pattern, no peripheral needle should be placed in template grid holes that are more than a couple of millimeters outside the PTV on the largest transverse slice. Unlike the basic uniform needle pattern, no peripheral needle should be placed more than 6 mm inside the PTV. For the schematic prostate of Figure 7, uppercase column and integer row coordinates A3, G3, B5, and F5 are used. Along these needle paths, seeds are placed at 1 cm increments on transverse planes that begin with a 5 mm offset from the base plane. Valid seed positions are no more than 0.6 cm inside or 0.5 cm outside the PTV. Although in Figure 7 only one seed is shown within the A3 and G3 needles (at the 2.5 cm plane), the needle would not be used unless there were valid seed positions on either or both the 1.5 cm and 3.5 cm planes. If the 2.5 cm plane is the only valid seed position, then the needle should be moved inward 0.5 cm to pick up more prostate tissue. Removal of Selected Central Seeds The final modification to uniform seed loading is to reduce the linear seed density in the central needles. Urethral sparing is implicit in this action, but explicit urethral sparing is not necessary if the brachytherapist tracks the path of the urethra in the OR. The number of needles to consider for seed removal and the removal pattern is highly dependent on the seed strength and the radionuclide. Low-strength seeds may require removal of only 1 or 2 seeds from the four needles closest to the urethra. High-strength seeds usually require removal of two or more seeds from about a dozen needles. Due to differences in their radial dose functions, 125I implants are more tolerant of large interior holes while 103Pd implants are less tolerant and need more careful attention to prevent a central low-dose region.
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A Sample Plan To illustrate these principles, a plan for a typical prostate, 31 cm3 in this case, is developed and analyzed in Figures 8 through 11. The prostate was first enlarged with 4 to 5 mm margins laterally and anteriorly and 7 mm superiorly and inferiorly to a bilaterally symmetric PTV. Margin was added posteriorly near the base and apex, and the radiation oncologist requested more generous margins at the apex so that the resulting PTV was 61 cm3, an enlargement factor of 1.97, which exceeds our standard enlargement factor of 1.8 for this size prostate. The apical PTV drawn for this patient is shown in Figure 9a based on the measured sagittal length of the gland. Had the prostate been a few mm longer, the alternate ending of Figure 9b would have been employed. The plan was created with uniformly loaded needles on lowercase columns and half-integer rows. On the base (0.0 cm) plane of Figure 8, there is only the PTV, no prostate. Because of the tilt of the prostate/bladder interface, seeds in the two most anterior rows (3.5 and 4.5) would likely end up in the bladder or bladder wall if executed exactly as planned. In the reality of the OR, those needle tips are not
Figure 8. Seed distribution and isodose plan for a 31 cm3 prostate (heavy grey line) after enlargement to a 61 cm3 PTV (heavy dark line). The base is at the upper left (0.0 cm offset) and the slices proceed at 0.5 cm intervals left to right culminating with the apex at lower right (4.0 cm offset). Two alternate ways of treating the apex are illustrated in Figure 9. The 100% and 150% isodose curves are shown, with the 100% as the outermost. Seeds present on a slice are small filled circles, and needle tracks empty on a slice are small open circles. The urethra appears at D3.5 at the base and curves downward to D3.8 at the base. The seed removal pattern from four central needles each on the 2.5 and 3.5 rows was designed to lower V150 to the desired range with urethral sparing as a secondary effect. The 150% isodose still touches the urethra in places, but additional sparing of the urethra was left for the brachytherapist in the OR as shown in Figure 10.
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Figure 9. Alternate treatments of the apex. In both rows (a) and (b), the first image is of the prostate apex, which is the last image in Figure 8. Row (a) was applied to this patient because the measured sagittal length of the prostate was
